Development of Endothelialisation Methods & in Vitro Testing Systems

Development of Endothelialisation Methods & In vitro Testing Systems for Vascular Stents

A thesis submitted for the degree of

Doctor of Philosophy (PhD)

in the subject of Biomedical Engineering

Adim De, MSc

November 2018

Hannover Medical School

International PhD Program ‘Regenerative Sciences’

in Hannover Biomedical Research School (HBRS)

Cardiopulmonary Regenerative Engineering,

NIFE

Acknowledged by the PhD committee and head of Hannover Medical School

President: Prof. Dr. Christopher Baum (until 31st December 2018)

Prof. Dr. Michael Manns (as of 1st January 2019)

Supervisor: Prof. Sotirios Korossis, Hannover Medical School

Co-Supervisors: Prof. Dagmar Wirth, Helmholtz-Zentrum für Infektionsforschung

Prof. Daniel Sedding, Hannover Medical School

External expert: Prof. Robert Thomas, Loughborough University

Internal expert: Prof. Birgit Glasmacher, Leibniz University

Day of public defence: 25th January 2019

PhD project funded by the DFG through a scholarship by the Cluster of Excellence REBIRTH and the DAAD – Graduate School Scholarship Program. Printed with the support of the German Academic Exchange Service.

Abstract Coronary heart disease is currently responsible for a significant percentage of global mortality in developed and developing nations alike. This occurrence takes place despite the advancement in medical technology and improved treatment options. The primary procedure involved in treating coronary heart disease is by means of balloon angioplasty and subsequent stenting. Due to complications with restenosis and stent thrombosis that are associated with current commercial stents, there has been a growing interest in stent research and development in order to eradicate the causes of such clinical events. In order to improve efficacy of the stent development phase, strategic and systematic preclinical ex vivo testing methods are required.

The present work assessed a range of different drugs for dual-functionality pertinent to the needs of drug eluting stents in order to combat restenosis and stent thrombosis. An array of biological testing was conducted on cells treated with the different drugs in vitro in order to identify the therapeutic potential, cytotoxic limitations, migratory and morphological behaviour, as well as the gene expression of the cells. Ferulic acid and exendin-4 were identified as suitable drugs that could be employed for the purpose of drug eluting stents due to their respective behaviour on inhibiting smooth muscle cell (SMC) and promoting endothelial cell (EC) proliferation. Curcumin and Magnolol, drugs that have not currently been employed in commercially available stents and at the tested concentration range did not exhibit appropriate dual-functional capabilities on the cells.

In addition, the work optimised hybridised polymeric coatings consisting of poly(D,L-lactide-co- glycolide) (PLGA) and type I collagen sourced from either rat tail or genetically modified tobacco plants yielding recombinant human collagen, for use as drug eluting mechanisms in stents. Cell viability and metabolic activity demonstrated that increasing the concentration of collagen within the hybrid coating allowed for better cell proliferation. It was also noted that neutralisation of the recombinant human collagen was necessary to improve cell proliferation. Recombinant human collagen – PLGA hybrid blends resulted in the best coatings in terms of EC proliferation.

In order to facilitate the systematic assessment of drug and coating performance, three bioreactor simulation systems were developed, commissioned and evaluated for in vitro testing. These included i) a bespoke multi-station pulsatile flow bioreactor for near-physiological conditioning of small-calibre blood vessels for simulating and assessing vascular stenting in vitro; ii) a static system for assessing EC migration on and endothelialisation of planar stent meshes; iii) a flow chamber providing laminar flow to EC-seeded planar stent meshes for assessing EC migration on and endothelialisation of stent struts. The 3 systems were shown to be meaningful testing platforms for assessing the endothelialisation of stent struts prior to in vivo studies.

Acknowledgements It is safe to say that this thesis is amongst the most challenging experiences of my life thus far and I would not have been able to complete it without the support and assistance of many people. I would like to express my gratefulness to my supervisor, Professor Sotirios Korossis, for giving me the opportunity to take on an incredibly challenging research project. Thank you for the patience, support and guidance throughout this project and for giving me a wide range of creative freedom that allowed me to develop my skills within various disciplines.

A great deal of my work would not have been possible without my colleague, mentor and friend, Michael. You are owed countless beers and I can’t express how grateful I am for all your help and support. To Lucrezia, Daniele and Sabrina, the postdocs that were extremely supportive during my time here. Massive thanks to all my past and present colleagues at LEBAO and NIFE for the hours of dinners, drinks and entertainment. My time at NIFE would not have been the same if it were not for my office mates, Artemis and Panos – plenty of ensued hilarity. My introduction into the group would not have been so seamless without the help of Eirini. I would also like to thank the REBIRTH admin staff: Daniela, Steffi, and Mariam for all the help and assistance throughout the years.

A great deal of my work would not have been completed without the help of several people. Thank you, Dr. Julia Meier and Dr. Wiegmann for providing me with fresh porcine tissue as well as the team at BioMedImplant. To Dennis Kundrat and Samuel Muller for the constant help with the 3D printing of my tissue holder prototypes. To Dr. Nadav Orr for kindly providing and supporting me with the CollPlant collagen. I would like to thank my co-supervisors, Prof. Sedding and Prof. Wirth, for the invaluable feedback and encouragement throughout the course of my PhD.

My sanity would not have stayed intact if it were not for the friends I have made here in Hannover. The social aspect kept our weekends packed. Most importantly, my housemates, old and new – Iratxe, Ivana, Sanela, Julio, Regis and Giacomo – there was never a dull day with the Isern crew. The Pendejos collective never failed to entertain. To many more trips and reunions in the future! I have to include my closest friends, Andrew and Lewis, cause I told them I would. Thank you for being you and visiting me in Germany for many an inebriated night. Big ups to John and Duncan, the new entries to the Catalina Wine Mixer. Kat and Has, long live the Oreo – thanks for always having my back over the last 8 years and to many more adventures ahead!

To my brother, Aamer, for always being there through good times and bad. So much has changed over the last few years and am grateful to have you in my life. Proud of you. Last but not the least, to my parents, Abhijit and Srilata, for their endless love, support and for guiding me to become who I am today.

Table of Contents

ABSTRACT I

ACKNOWLEDGEMENTS II

LIST OF FIGURES VII

LIST OF TABLES XI

NOMENCLATURE XII

INTRODUCTION 1

AIMS AND OBJECTIVES 2

CHAPTER I: LITERATURE REVIEW 3

1.1 CARDIOVASCULAR SYSTEM 3 1.1.1 BLOOD VESSELS 3 1.1.1.1 Haemodynamics 5 1.1.2 ENDOTHELIAL CELLS 6 1.1.3 ENDOTHELIAL PROGENITOR CELLS 7 1.1.4 SMOOTH MUSCLE CELLS 7 1.2 CORONARY ARTERY DISEASE 8 1.2.1 MECHANISM OF RESTENOSIS 10 1.2.2 MECHANISM OF THROMBOSIS 11 1.3 STENTS: CURRENT AND FUTURE STATE 11 1.3.1 IDEAL STENT DESIGN 12 1.3.2 BARE METAL STENTS (BMS) 14 1.3.3 DRUG ELUTING STENTS (DES) 19 1.3.3.1 First Generation DES 19 1.3.3.2 Second Generation DES 20 1.3.4 CURRENT COMMERCIALLY USED DRUGS FOR DES 21 1.3.5 DRUG DELIVERY MECHANISM 24 1.3.6 SURFACE MODIFICATION 26 1.3.7 ORGANIC COATINGS/ADDITIVES 29 1.3.8 BIO-STABLE POLYMERS 30 1.3.9 BIODEGRADABLE POLYMERS 31 1.3.9.1 Poly(lactic-co-glycolic acid) 32

iii

1.3.10 POLYMER DEPOSITION TECHNIQUES 34 1.3.11 BIOMOLECULES 35 1.3.11.1 Collagen 39 1.3.12 SCREENING OF POTENTIAL BI-FUNCTIONAL DRUGS FOR DES APPLICATIONS 41 1.3.12.1 Magnolol 42 1.3.12.2 Curcumin 43 1.3.12.3 Ferulic Acid 44 1.3.12.4 Exendin-4 44 1.4 BIOREACTORS 45 1.4.1 VASCULAR BIOREACTORS 48 1.4.2 PARALLEL PLATE FLOW 52 1.5 LIMITATIONS OF THE CURRENT STATE 53

CHAPTER II: DRUG ASSESSMENT 54

2.1 INTRODUCTION 54 2.2 MATERIALS AND METHODS 54 2.2.1 CELL CULTURE 54 2.2.1.1 Cells and Cell Culture Consumables 54 2.2.1.2 Cell Seeding of Cryopreserved Cells 56 2.2.1.3 Cell Passaging 56 2.2.1.4 Cell Cryopreservation 57 2.2.2 DRUG PREPARATION 57 2.2.3 METABOLIC ACTIVITY ASSAY 57 2.2.4 WOULD HEALING ASSAY 58 2.2.5 GENE EXPRESSION 60 2.2.4.1 Endothelial Cell – Gene Expression Assessment 60 2.2.4.2 Smooth Muscle Cell – Gene Expression Assessment 60 2.2.4.3 RNA Isolation 60 2.2.4.4 cDNA synthesis 61 2.2.4.5 Quantitative-Polymerase Chain Reaction (qPCR) 61 2.2.4.6 Data Analysis for Gene Expression Quantification 62 2.2.6 FLOW CYTOMETRY 63 2.2.7 STATISTICAL ANALYSIS 63 2.3 RESULTS 64 2.3.1 METABOLIC ASSESSMENT OF DRUGS ON ENDOTHELIAL CELLS 64 2.3.1.1 Human Umbilical Vein Endothelial Cells 64 2.3.1.2 Porcine Coronary Artery Endothelial Cells 69 2.3.2 METABOLIC ASSESSMENT OF DRUGS ON SMOOTH MUSCLE CELLS 72 2.3.2.1 Human Coronary Artery Smooth Muscle Cells 72 2.3.2.2 Porcine Coronary Artery Smooth Muscle Cells 77 2.3.3 WOUND HEALING ASSAY 80 2.3.4 GENE EXPRESSION OF CELLS TREATED WITH DRUGS 84 2.3.4.1 Endothelial Cells 84 2.3.4.2 Smooth Muscle Cells 85 2.3.5 ENDOTHELIAL CELL FUNCTION 85

2.4 DISCUSSION 87

CHAPTER III: POLYMER ASSESSMENT 94

3.1 INTRODUCTION 94 3.2 MATERIAL AND METHODS 94 3.2.1 COATING PREPARATION 94 3.2.2 ASSESSMENT OF CELL ATTACHMENT 95 3.2.3 ASSESSMENT OF THE GELATION AND PREPARATION OF RHCOL1 95 3.2.4 ASSESSMENT OF COATING CYTOTOXICITY 96 3.2.5 FOURIER TRANSFORM INFRARED (FTIR) SPECTROSCOPY 97 3.2.2 STATISTICAL ANALYSIS 97 3.3 RESULTS 98 3.3.1 ASSESSMENT OF CELL ATTACHMENT 98 3.3.2 ASSESSMENT OF THE GELATION AND PREPARATION OF RHCOL1 99 3.3.3 ASSESSMENT OF COATING CYTOTOXICITY 101 3.3.4 CHEMICAL CHARACTERISATION 103 3.4 DISCUSSION 108

CHAPTER IV: DEVELOPMENT OF IN VITRO SIMULATION SYSTEMS FOR STENT TESTING 112

4.1 INTRODUCTION 112 4.2 CULTURE CHAMBER SYSTEM 112 4.2.1 STATIC CULTURE SYSTEM 112 4.2.1.1 Design Specification & Conceptual Design 112 4.2.1.2 Manufactured device 114 4.2.3 DYNAMIC SYSTEM 116 4.2.3.1 Design Specification & Conceptual Design 116 4.2.3.2 Prototyping 117 4.2.3.3 Manufactured Device 119 4.3 MULTI-STATION PULSATILE FLOW SYSTEM 119 4.3.1 DESCRIPTION OF SYSTEM DESIGN 119 4.3.1.1 Base Frame 119 4.3.1.2 Tissue Stations 120 4.3.1.3 Linear Actuator 121 4.3.1.4 Heating System 121 4.3.1.5 Chemostat 122 4.3.1.6 Culture Medium Circulation & Bioreactor Setup 123 4.3.2 PARAMETER CONTROL AND DATA ACQUISITION 125 4.3.3 HEATING SYSTEM CALIBRATION 128 4.3.4 MOTOR PROFILE CALCULATIONS 128 4.3.5 PRESSURE PROFILE OPTIMISATION 131 4.3.6 OPTIMISATION OF TISSUE HANDLING AND MOUNTING 133 4.3.6 BIOREACTOR SYSTEM SET UP 134

CHAPTER V: TISSUE VIABILITY & STENT ENDOTHELIALISATION UNDER STATIC & DYNAMIC CONDITIONS 135

5.1 INTRODUCTION 135 5.2 MATERIALS AND METHODS 135 5.2.1 TISSUE PROCUREMENT AND DISSECTION 135 5.2.2 TRANSPORT SOLUTIONS 136 5.2.3 EFFECT OF ANTIBIOTIC TREATMENT ON TISSUE VIABILITY 136 5.2.4 EFFECT OF LONG-TERM STATIC CULTURE ON TISSUE VIABILITY 136 5.2.5 EFFECT OF COLD STORAGE IN TIPROTEC AND ANTIBIOTICS ON TISSUE VIABILITY 136 5.2.6 STATIC CULTURE SYSTEM EXPERIMENTS 137 5.2.7 DYNAMIC CULTURE SYSTEM EXPERIMENTS 137 5.2.7.1 Arterial Tissue Cultures 138 5.2.7.2 Endothelialised Silicone Sheet Cultures 138 5.2.8 VIABILITY ASSESSMENT 139 5.2.8.1 Metabolic Activity 139 5.2.8.1 Fluorescence Imaging 139 5.2.9 STATISTICAL ANALYSIS 140 5.3 RESULTS 140 5.3.1 EFFECT OF ANTIBIOTIC TREATMENT ON TISSUE VIABILITY 140 5.3.3 EFFECT OF LONG TERM STATIC CULTURE ON TISSUE VIABILITY 140 5.3.4 EFFECT OF COLD STORAGE IN TIPROTEC AND ANTIBIOTICS ON TISSUE VIABILITY 142 5.3.5 TISSUE HOLDER EXPERIMENTS 143 5.4 DISCUSSION 150

CHAPTER VI: CONCLUSION AND FUTURE WORK 152

6.1 CONCLUSION 152 6.2 FUTURE RECOMMENDATIONS 154

REFERENCES 156

APPENDIX I 177

APPENDIX II 181

AUTHOR’S STATEMENT OF CONTRIBUTION 183

DECLARATION ERROR! BOOKMARK NOT DEFINED.

List of Figures Figure 1.1 - Various forms of arteries. 3 Figure 1.2 - Various forms of veins. 4 Figure 1.3- Various forms of capillaries. 4 Figure 1.4 - Variation in pressure amongst the different blood vessels. 5 Figure 1.5 - Aortic Pressure Profile and Coronary flow. 6 Figure 1.6 - Expression products of endothelial cells. 7 Figure 1.7 - Phenotypic switching of smooth muscle cells. 8 Figure 1.8 - Coronary artery disease prevalence by sex and age. 9 Figure 1.9 – Development of the atherosclerotic plaque. 9 Figure 1.10 - Neointimal formation on stainless steel bare metal stent. 10 Figure 1.11 - Mechanism of restenosis. 10 Figure 1.12 - Venn diagram illustrating the key aspects described by Ako et al. (2007) 13 Figure 1.13 – Stress-Strain graph of a Cobalt alloy and 316L Stainless Steel Alloy. 16 Figure 1.14 – Stress-strain curves of natural tissue, nitinol and steel. 17 Figure 1.15 - Sirolimus and paclitaxel mechanism. 23 Figure 1.16 - Diagrams of various drug release mechanisms. 24 Figure 1.17 – Representations of various forms of surface modifications. 27 Figure 1.18 – Chemical Structure of PEVA and PBMA. 30 Figure 1.19 – Chemical structure of SIBS. 30 Figure 1.20 – Chemical Structure of PVDF-HFP. 31 Figure 1.21 – Chemical structures of the components of the Biolinx polymer. 31 Figure 1.22 – Chemical structure of PLGA and its respective by-products through hydrolytic degradation. 33 Figure 1.23 – Concept of antibody immobilization onto a polymeric surface. 36 Figure 1.24 – Immobilization of Anti-CD133 antibody procedure. 37 Figure 1.25 - Simple scheme of the Swanson LBD Paradigm. 41 Figure 1.26 - Structure of Magnolol. 42 Figure 1.27 - Structure of Curcumin. 43 Figure 1.28 - Structure of Ferulic Acid. 44 Figure 1.29 – Schematic of the spinner flask bioreactor system. 46 Figure 1.30 – Schematic of the rotating wall bioreactor. 47 Figure 1.31 – Schematic of a model perfusion flow system. 48 Figure 1.32 – Schematic of the Niklason vascular bioreactor. 50 Figure 1.33 – Schematics of parallel plate flow systems used in current research. 52 Figure 2.1 - The mechanism of the WST-8 assay. 58 Figure 2.2 - Setup of the Axiovision Z1 Microscope with the incubation chamber used for the Scratch Assay 59 Figure 2.3 – Mean metabolic activity of HUVECs treated with different concentrations of Paclitaxel and Everolimus for 24h, 48h, 72h and 144h. 65 Figure 2.4 - Mean metabolic activity of HUVECs treated with different concentrations of curcumin for 24h, 48h, 72h and 144h. 67 Figure 2.5 - Mean metabolic activity of HUVECs treated with different concentrations of magnolol for 24h, 48h, 72h and 144h. 67 Figure 2.6 - Mean metabolic activity of HUVECs treated with different concentrations of ferulic acid for 24h, 48h, 72h and 144h. 68 Figure 2.7 - Mean metabolic activity of HUVECs treated with different concentrations of Exendin-4 for 24h, 48h, 72h and 144h. 68 Figure 2.8 - Mean metabolic activity of PCAECs treated with different concentrations of curcumin for 18h, 24h, 36h, 48h, 72h and 144h. 70 Figure 2.9 - Mean metabolic activity of PCAECs treated with different concentrations of magnolol for 18h, 24h, 36h, 48h, 72h and 144h. 71

vii

Figure 2.10 - Mean metabolic activity of PCAECs treated with different concentrations of ferulic acid for 18h, 24h, 36h, 48h, 72h and 144h. 71 Figure 2.11 - Mean metabolic activity of PCAECs treated with different concentrations of Exendin-4 for 18h, 24h, 36h, 48h, 72h and 144h. 72 Figure 2.12 - Mean metabolic activity of HCASMCs treated with different concentrations of Paclitaxel and Everolimus for 24h, 48h, 72h and 144h. 73 Figure 2.13 - Mean metabolic activity of HCASMCs treated with different concentrations of curcumin for 24h, 48h, 72h and 144h. 75 Figure 2.14 - Mean metabolic activity of HCASMCs treated with different concentrations of magnolol for 24h, 48h, 72h and 144h. 75 Figure 2.15 - Mean metabolic activity of HCASMCs treated with different concentrations of Ferulic Acid for 24h, 48h, 72h and 144h. 76 Figure 2.16 - Mean metabolic activity of HCASMCs treated with different concentrations of Exendin-4 for 24h, 48h, 72h and 144h. 76 Figure 2.17 - Mean metabolic activity of PCASMCs treated with different concentrations of curcumin for 18h, 24h, 36h, 48h, 72h and 144h. 78 Figure 2.18 - Mean metabolic activity of PCASMCs treated with different concentrations of magnolol for 18h, 24h, 36h, 48h, 72h and 144h. 79 Figure 2.19 - Mean metabolic activity of PCASMCs treated with different concentrations of ferulic acid for 18h, 24h, 36h, 48h, 72h and 144h. 79 Figure 2.20 - Mean metabolic activity of PCASMCs treated with different concentrations of Exendin-4 for 18h, 24h, 36h, 48h, 72h and 144h. 80 Figure 2.21 – Percentage reduction of wound area achieved by PCAECs under different drug exposure over a 48h period. 82 Figure 2.22 - Fluorescence microscopy of calcein-stained PCAEC monolayers, subjected to the scratch assay, after 48h exposure to different drug dosages. 82 Figure 2.23 - Percentage reduction of wound area achieved by PCASMCs under different drug exposures over a 48h period. 83 Figure 2.24 - Fluorescence microscopy of calcein-stained PCSMCs, subjected to the scratch assay, after 48h exposure to different drug dosages. 83 Figure 2.25 – Fold change in CDH5 (A), CD31(B) and NOS3 (C) expression in PCAECs after 144h exposure to Exendin-4 10nM (Ex10), Exendin-4 100nM (Ex100), Ferulic Acid 50µM (FA50) and Ferulic Acid 500 µM (FA500), compared to the untreated control. 84 Figure 2.26 – Fold change in MYH11, SMTN, ACTA and CNN1 expression in PCASMCs after 24h, 48h and 72h exposure to Ferulic Acid 50µM (A), Ferulic Acid 500 µM (B), Exendin 10nM (C) and Exendin 100nM (D), compared to the untreated control. 86 Figure 2.27 - Flow cytometry results of PCAECs cultured in to Exendin 10nM (Ex10), Exendin 100nM (Ex100), Ferulic Acid 50µM (FA50) and Ferulic Acid 500 µM (FA500). 86 Figure 2.28 – Model of migration mechanism of endothelial cells. 90 Figure 2.29 - Expression levels of SMC-associated genes for the two phenotypes. 91 Figure 2.30 - Signalling pathway demonstrating the antioxidaive potential of exendin-4. 93 Figure 3.1 – Mean metabolic activity of PCAECs seeded onto the different PLGA-collagen blends after (A) 30min and (B) 1h of culture. 99 Figure 3.2 - Calcein staining of PCAECs onto the different PLGA-collagen blends after 72h culture. 100 Figure 3.3 – Mean metabolic activity of PCAECs seeded on rHCol1 prepared under different gelation temperatures, with and without neutralisation. 100 Figure 3.4 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rtCol1 blends. 102 Figure 3.5 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rHCol1 (non-neutralised) blends. 102 Figure 3.6 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rHCol1 (neutralised) blends. 103 viii

Figure 3.7 - FTIR spectra of PLGA, rtCol1, rHCol1, and their respective blends (20% (w/w)) with PLGA 105 Figure 3.8 - FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (brown) and high (blue) concentrations of Ferulic Acid 106 Figure 3.9 - FTIR spectra of rtCol1-PLGA (20% (w/w)) blend coating with low (blue), medium (violet) and high (red) concentrations of Ferulic Acid 106 Figure 3.10- FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (violet) and high (blue) concentrations of Exendin-4 107 Figure 3.11- FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (brown) and high (blue) concentrations of Exendin-4 107 Figure 3.12 - Observation of a double peak within the rHCol1 (blue) and rtCol1 (red) samples associated with the amide I band. 111 Figure 4.1 – Concept of in situ endothelialisation. 113 Figure 4.2 – Schematics of the planar tissue-stent mating holder of the static system, featuring the individual components. 114 Figure 4.3 – Schematic of the flat stent mesh 114 Figure 4.4 – Differences in cut quality on different strut widths 115 Figure 4.5 - Manufactured tissue holder with sample tissue and mesh 115 Figure 4.6 - Silicone mould for the mesh holder 116 Figure 4.7 - Schematics of the flow chamber of the dynamic system for the planar tissue-stent system, featuring the top (red) and bottom (blue) flow unit, the tissue-stent mating holder, the stent mesh and the upper (pink) and lower (grey) silicone gaskets. 117 Figure 4.8 – 3D printed components (from left to right: Top flow unit, top silicone gasket, bottom silicone gasket and the bottom flow unit) of the flow chamber (top) and the assembled prototype with the mesh holder adopted from the static system (bottom). 118 Figure 4.9 - Region of leakage observed in the prototype 118 Figure 4.10 - Final Design of the Flow Chamber System 118 Figure 4.11 – Assembled and the manufactured Flow Chamber System 119 Figure 4.12 - Schematic of the multi-station pulsatile flow bioreactor 120 Figure 4.13 - Schematic of the tissue chamber with the attached compliance chamber 121 Figure 4.14 – Heating system components. Left: Control unit; right: Heating pads. 122 Figure 4.15 - Chemostat system including the control unit and the chemostat vessel. 123 Figure 4.16 - Diagram illustrating the bioreactor system set up. 124 Figure 4.17 – Full setup of the bioreactor system with two stations 124 Figure 4.18 - Flow path within the tissue chamber 125 Figure 4.19 - Experimental and factory values (dotted) of the flow rate generated by the peristaltic pump. 126 Figure 4.20 - Marked 1/4” – 28 UNF hole for PRESS-S-000 pressure transducer fitting (right). 126 Figure 4.21 – Positioning (black arrow) of the distal pressure transducer (PRESS-N-025; right). 126 Figure 4.22 - Schematic of the pressure transducer calibration set up 127 Figure 4.23 – Standard curve of raw voltage (mV) vs pressure (mmHg) for 2 sensors of each type. 128 Figure 4.24 - Representative aortic pressure waveform. 129 Figure 4.25 - Exemplary calculated actuator motion profile during simulated systole and diastole. 131 Figure 4.26 - Pressure waveform at 1 Hz with and without peripheral resistance. 132 Figure 4.27 - Pressure profile using altered input parameters with peripheral resistance. 132 Figure 4.28 - Superimposed physiological aortic pressure (red) and bioreactor pressure profiles. 133 Figure 4.29 - Sutured carotid artery onto luer barbed fittings 133 Figure 4.30 - Loose Suture causing slippage of artery. 134 Figure 4.31 - Autoclavable zip-ties keeping the tissue in place 134 Figure 5.1 - Segment of excised porcine aortic arch. 135 Figure 5.2 - Exploded Flow Chamber System Render with endothelialised sheet (left) and excised carotid artery (right). 137 Figure 5.3 - Metabolic activity of porcine carotid artery treated in PBS containing the antibiotic cocktail. 140 ix

Figure 5.4 - Metabolic activity of tissue samples cultured in DMEM 10% for 7 days. 141 Figure 5.5 - Live/Dead staining of tissue samples at day 0 and Live/Nuclei Staining of tissue samples at day 7. 141 Figure 5.6 - Metabolic activity of tissues cultured for 7 days. 142 Figure 5.7 - Metabolic activity of tissues cultured for 11 days. 142 Figure 5.8 - Metabolic activity of tissues cultured for 7 days after 10 hours of Tiprotec storage at 4°C. 143 Figure 5.9 – Live (green)/dead (red) staining of carotid artery segments stored in Tiprotec and cultured for 7 days. 143 Figure 5.10 – Nuclei (blue) and viable cells (green) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 2 days. 144 Figure 5.11 - Nuclei (blue) and viable cells (green) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 4 days. 145 Figure 5.12 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 3 days. 146 Figure 5.13 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 7 days. 146 Figure 5.14 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 10 days. 147 Figure 5.15 - Metabolic Activity of Porcine Carotid Arteries Cultured under Static and Flow Conditions (Left) and the Weight-Normalised Version (Right). 148 Figure 5.16 - Viable cell staining of PCAEC-seeded silicone sheet cultured with the uncoated SS stent for 3 days under flow. 149 Figure 5.17 – Viable cell staining of PCAEC-seeded silicone sheet cultured with the uncoated SS stent mesh for 3 days under static conditions. 149 Figure 5.18 - Viable cell staining of PCAEC-seeded silicone sheet cultured with the SS mesh, coated with PLGA 20% (w/w) rhCol1 and 1.8% (w/w) Ferulic Acid for 3 days. 149 Figure 0.1 - Univessel ready for sterilisation 179 Figure 0.2 - User interface of the chemostat system 180

List of Tables TABLE 1.1 - GENERAL IDEAL STENT PROPERTY CONSIDERATIONS 12 TABLE 1.2 - PROPERTIES OF DIFFERENT STAINLESS STEEL ALLOYS 15 TABLE 1.3– PROPERTIES OF DIFFERENT COBALT-CHROMIUM ALLOYS 16 TABLE 1.4 - PROPERTIES OF NITINOL PHASES 18 TABLE 1.5 – PROPERTIES OF MISCELLANEOUS MATERIALS 18 TABLE 1.6 – COMPARISON OF THE FIRST GENERATION DES 20 TABLE 1.7 – COMPARISON OF THE SECOND GENERATION DRUG-ELUTING STENTS 21 TABLE 1.8 – COMPARISON TABLE OF THE DRUGS USED IN DES 24 TABLE 1.9 - FACTORS AFFECTING DRUG RELEASE FROM POLYMERS 25 TABLE 1.10 - OPTIMUM PROPERTIES FOR POLYMER COATINGS FOR STENTS 30 TABLE 1.11 - COMPARISON OF ANIMAL-DERIVED COLLAGEN AND RECOMBINANT COLLAGEN 40 TABLE 1.12 – CONTROL PARAMETERS CRITICAL FOR VASCULAR BIOREACTOR DESIGN 48 TABLE 2.1 - CELLS USED IN THE STUDY 56 TABLE 2.2 - SOLUTIONS AND MEDIA USED FOR CELL CULTURE 56 TABLE 2.3 - DRUGS USED IN THE STUDY 57 TABLE 2.5 - DRUG CONCENTRATIONS FOR QPCR OF PCAECS 60 TABLE 2.6 - DRUG CONCENTRATIONS FOR PCASMC TIME-DEPENDANT QPCR 60 TABLE 2.7- DNA OLIGONUCLEOTIDE PRIMERS SEQUENCES FOR RT-PCR 61 TABLE 3.1 - SAMPLE AND CONSTITUENT LIST FOR A SINGLE WELL COATING (IN 24 WELL PLATE). 95 TABLE 3.2 - SAMPLE AND CONSTITUENT LIST FOR A SINGLE WELL COATING (IN 48 WELL PLATE) 96 TABLE 4.1 - ISMATEC DATA SHEET VALUES 125 TABLE 4.3 - MOTOR PROFILE PARAMETERS AND CRITICAL VALUES 130 TABLE 0.1 - INPUT SETTINGS FOR CONDITIONING CULTURE MEDIUM USING THE CHEMOSTAT 180

Nomenclature

∆∆Ct Delta-Delta Ct °C Degrees Celsius µM Micromolar 3D Three-Dimensional ANOVA Analysis of Variance BMS Bare Metal Stent cDNA Complimentary Deoxyribonucleic Acid Ct Threshold Cycle DES Drug Eluting Stent DMEM Dulbecco's Modified Eagle Medium DMSO Dimethyl Sulfoxide DO Dissolved Oxygen EC Endothelial Cell EGM-2 Endothelial Growth Medium - 2 eNOS Endothelial Nitric Oxide Synthase EPC Endothelial Progenitor Cell Ex Exendin-4 FA Ferulic Acid FTIR Fourier Transform Infrared GLP-1 Glucagon-like Peptide 1 HCASMC Human Coronary Artery SMC HFIP 1,1,1,3,3-hexafluori-2-propanol HUVEC Human Umbilical Vein EC LDL Low Density Lipoprotein mM Milimolar mmHg Millimetre Mercury NO Nitric Oxide P/S Penicillin/Streptomycin PCAEC Porcine Coronary Artery EC PCASMC Porcine Coronary Artery SMC PCI Percutaneous Coronary Intervention PCR Polymerase Chain Reaction PLGA Poly(lactic-co-glycolic acid) Pn Passage (number) qPCR Real-Time PCR rHCol1 Recombinant Human Collagen Type I RNA Ribonucleic Acid rtCol1 Rat Tail Collagen Type I SMC Smooth Muscle Cell SMDM Smooth Muscle Differentiation Medium SMGM Smooth Muscle Growth Medium SS Stainless Steel VEGF Vascular Endothelial Growth Factor WST-8 Water Soluble Tetrazolium Salt

τw Wall Shear Stress

xii

Introduction The progression of stent development has been a major field of research and development since the introduction of vascular stents. Moreover, the progressive nature of cardiovascular disease has also led to the increase in the number of patients that suffer from stent-related complications inherited in the current commercially available stents, including in-stent restenosis and thrombosis [1]. The current shift in the aging demographic has led to the increase in patients requiring stent intervention due to the increase in population suffering from coronary heart disease. It is estimated that by 2030, the cardiac mortality rate will rise to 22.2 million [2]. With more than 3 million stents implanted per year [3], even a seemingly insignificant thrombosis complication rate of 5% can be the cause of thousands of deaths (approximately 1-2% of total stented patients) per year amongst drug eluting stent implants [4]. There is a principle need to assess stent coatings and therapeutic options within a systematic in vitro/ex vivo platform prior to in vivo and clinical studies in order to minimise these risk rates.

The work presented in this thesis comprises drug and polymer assessments in addition to the development of a platform to assess the capacity of the stent coating for re-endothelialisation. This multi-disciplinary research focusses on the proof-of-concept principles of developing a platform with the perspective to identify potential therapeutic options to improve stent performance through the alleviation of thrombosis and in-stent restenosis.

Chapter I describes the relevant biological background regarding coronary heart disease and percutaneous coronary intervention. A literature review of current research pertaining to stent development, drugs and coatings as well as bioreactor technology is a focal point of the chapter. Chapter II describes the work conducted on cell-drug interactions using the different therapies that were screened from literature review in Chapter I. This chapter details the effect of these drugs on endothelial cells and smooth muscle cells with regards to metabolic activity, gene expression, migration rate and endothelial functionality. Chapter III addresses the optimisation and the development of the hybrid polymer stent coating. Endothelial cell adhesion and proliferation were presented in addition to chemical characterisation of the hybrid coatings. Chapter IV highlights the development and commissioning stages of the three bioreactor systems as well as the incorporation of a chemostat. Chapter V presents the optimisation and study of porcine carotid tissue utilised in the aforementioned testing platforms from Chapter IV. Various proof of concept studies were assessed and described. Chapter VI concludes the key observations of this research and the specific observation in regard to the primary aims of the work. Recommendation for future work was also provided in order to tighten the gap between research and an effective stent testing platform.

Aims and Objectives The aim of this research was to develop a systematic methodology and tools for assessing vascular stent coating strategies that can offer dual functionality, including both inhibition of smooth muscle cell proliferation and promotion of endothelialisation. The specific objectives of the work were to:

1. assess different drugs for their potential incorporation in drug eluting stents that would: a. promote stent endothelialisation; b. inhibit smooth muscle cell migration and proliferation. 2. identify and optimise a polymeric coating that would: a. allow for rapid cell adhesion with no cytotoxic effects; b. demonstrate the capacity for drug release. 3. commission a bespoke multi-station bioreactor for use as an in vitro near-physiological stent testing platform by: a. optimising user interface and handling; b. ensuring a fully controllable and adaptable system for ex vivo tissue studies. 4. design, develop and test a static and dynamic system for stent testing that would: a. maintain tissue viability; b. support stent endothelialisation.

Chapter I: Literature Review The development of the platform requires a thorough understanding of a variety of subjects, much of which is highlighted in the literature review chapter including but not limited to: cardiovascular system and pathology, current state of stent research and bioreactor development and design. The aims and objectives of the study are outlined at the end of this chapter.

1.1 Cardiovascular System The cardiovascular system is a closed system comprised of the heart, the blood and a complex network of blood vessels. The system serves as a means of blood transport, protection and regulation of the body.

1.1.1 Blood Vessels The transport system of blood vessels is differentiated into 3 classifications; arteries, veins and capillaries. The arteries transport blood away from the heart, the veins transport blood to the heart and capillaries are the intermediate vessels connecting the arteries and veins. The tissue structure of each vessel varies due to different functional purpose.

The high output pressure of blood from the heart requires arteries to have thick walls. The arteries close to the heart have an abundance of elastic fibres (observable in Figure 1.1). The elasticity allows regulation of blood pressure within the vessels. The arteries located further from the heart exhibit thicker layers of smooth muscle cells within the tunica media, as well as a lower percentage of elastic fibres. These muscular arteries are capable of vasoconstriction, which consequently limits the elastic potential of the artery. The blood pressure at this point has been dampened and therefore arterial expansivity is not a primary factor. The arterioles are micrometre scale arteries that lead to capillaries and have a similar cross-sectional set up as the larger arteries but on a much smaller scale. Their scaled- down structure allows blood flow resistance.

Figure 1.1 - Various forms of arteries. Adopted and modified from Saladin (2011)[5]

Figure 1.2 - Various forms of veins. Adopted and modified from Saladin (2011) [5]

Figure 1.3- Various forms of capillaries. Adopted and modified from Saladin (2011) [5]

Veins have large lumens with relatively thin walls. Certain veins contain valves, which regulate the direction of blood flow, which is an issue due to the lowered blood pressure and gravitational force. Similar to the arterioles, venules are smaller scaled (8-100µm) versions of veins that join the capillary beds and are structured in a way that a group of venules come together to eventually form a vein. The variations of veins are presented in Figure 1.2. Capillaries are micro-scaled vessels (illustrated in Figure 1.3) with a lumen diameter of 5-10µm and are used to supply blood directly to the tissue. Gases, metabolites and nutrients are diffused through the thin capillary walls. The cellular wall itself is only one cell thick.

The vessel wall of the arteries and veins are composted of; the tunica intima, the tunica media and the tunica adventitia. The intimal layer has direct contact to blood and therefore, flow-related shear stress. It is composed of simple squamous epithelium (endothelium) atop a basement membrane and loose connective tissue. The endothelium is the permeable barrier that controls chemical exchange into and out of the blood stream. The endothelium also restricts platelet and blood cell adhesion unless a

wound healing cascade is activated. The tunica media is the thick layer which consists of collagen, smooth muscle cells and occasionally elastic tissue. The media supports and prevents the vessel from rupture due to blood pressure. It is also capable of vasomotion, the ability to widen or narrow the vessel. Vessel classification governs the amount of muscle or elastic tissue within the vessel. The outermost layer, the tunica adventitia is primarily comprised of collagen fibres, with some bands of elastic fibres and also contains some smooth muscle fibres in veins. It is the thickest layer in veins and even in certain large arteries. This purpose of this outer layer is to prevent vessel movement through connections to surrounding tissue [5][6].

1.1.1.1 Haemodynamics Haemodynamics defines the physical properties of blood as a fluid. It is the relationship between pressures, flow, fluid velocity, fluid viscosity and resistance which ultimately results in the observable arterial pressure and cardiac output. The different distances, sizes and branching points of blood vessels would mean different pressure profiles between them. As shown in Figure 1.4, the arteries, due to its close proximity and direct path of the blood flow leaving the heart results in a high pressure range. Changes in physiological blood flow will cause the blood vessels to adapt in order to maintain suitable conditions. The circumferential and axial stresses on the blood vessels are governed by the pressure within the structure. The contraction (systole) and relaxation (diastole) of the heart results in these cyclic pressure changes throughout the cardiovascular system. The physiological pressure profile of arteries is graphed in Figure 1.5. Aortic pressure at physiological conditions ranges between 80mmHg and 120mmHg. A key characteristic of blood flow in arteries is the strong pulsatile nature and the elasticity of the artery causes vascular recoil upon the differences in pressure [7].

Figure 1.4 - Variation in pressure amongst the different blood vessels. Adopted from Elad (2004) [8]

Figure 1.5 - Aortic Pressure Profile and Coronary flow. Adopted from Crystal and Pagel (2018) [9]

The vascular system follows Darcy’s law which is “Q = ∆P/R”, where Q is flow, ∆P is pressure differential and R is resistance. This is usually coupled with Q = VA, where V is fluid velocity and A is cross sectional area of the blood vessel. These equations demonstrate the relationship between these various parameters and that major changes to any of these parameters would effect the stresses exerted on the vascular tissue, resulting in cellular changes through mechanotransduction in order to compensate for the variation in hemodynamics.

1.1.2 Endothelial Cells Endothelial cells (ECs) are the cellular component of the endothelial monolayer which lines all blood vessels in the body with a diameter of 8-12µm. With over ten trillion cells, they can be considered an organ, weighing in at approximately 1 kilogram. The EC formation process begins with the splanchnopleuric mesoderm, which are transformed into mesenchymal cells, which are subsequently differentiated into the hemangioblast. They are transformed into pre-endothelial cells which could either differentiate into the hematopoietic cell line or endothelial cells. Studies have shown that endothelial cells can also differentiate into mesenchymal cells or intimal smooth muscle cells. It has been shown that there is a possibility of great phenotypic variation, based on the individual, the vessel type and position of the vessel. A varied phenotype could mean exhibiting different responses to the same stimuli. This could cause complications when studying the cell line in a laboratory setting. The semi-permeable ECs barrier controls molecular transfer between the blood and the vessel wall. They also secrete various products that influence the underlying smooth muscle cells as well as the elements in circulating blood. These cells are also highly relevant in terms of maintaining non-thrombogenic interfaces through autocrine, paracrine and endocrine secretion, which regulate various factors. Figure 1.6 lists the various expression products of endothelial cells [10].

Figure 1.6 - Expression products of endothelial cells. Adopted from Sumpio et al. (2002) [10]

1.1.3 Endothelial Progenitor Cells Endothelial progenitor cells (EPC), derived from bone marrow are known to be critical in the process of postnatal vasculogenesis and are sourced from bone marrow. They have the ability to migrate, proliferate and differentiate into mature ECs, which would subsequently be crucial for the development of vascular structures. During tissue damage, EPCs migrate to the site, not only to promote neovasculature formation, but also release cytokines and growth factors in order to encourage pre-existing ECs to contribute to vascular regeneration [11].

Identifying EPCs is posed as a challenging task as they express CD34+ and vascular endothelial growth factor receptor-2 (VEGFR-2+) markers, which is also observable on mature ECs. The CD133 marker for stem cells is more specific to EPCs but the expression decreases as differentiation progression occurs. It has been shown that EPC functionality and circulating population decreases in patients who suffer(ed) from coronary artery disease and strokes [12,13]. This would limit the rate of endothelialisation and vascular regeneration in diseased patients if the primary objective was to attract circulating EPCs.

1.1.4 Smooth Muscle Cells Smooth muscle cells (SMCs) are non-striated mononuclear cells that are found in the vascular wall, primarily, the tunica media. With an adjustable structure and functionality based on the physiological environment, they allow the vessel to withstand high pressures of blood flow as well as being crucial to the arterial repair procedure. SMCs exists as either contractile quiescent or proliferative synthetic states, along with some intermediate phenotypes [14]. The phenotypic switch is initiated by environmental and extracellular factors. SMCs are heavily subject to mechanical forces and have the capacity to adapt to these changing forces via mechanotransduction. Figure 1.7 highlights key initiators that influence the phenotypic changes from contractile to synthetic and vice versa.

Figure 1.7 - Phenotypic switching of smooth muscle cells. Adopted from Al-Shehabi et al. (2016) [15]

1.2 Coronary Artery Disease Coronary Artery Disease (CAD) is the leading cause of mortality globally where the highest risk lies in patients over the age of 60 as shown in Figure 1.8 [16]. The condition is based on several risk factors including smoking, diabetes, diet, old age, gender and genetic issues. Since coronary arteries supply oxygenated blood to the cardiac muscles, any obstruction of the coronary arteries would limit the nutrition supplied to the associated region of muscle and could potentially lead to cell death and failure of the muscle.

Atherosclerosis is classed as an inflammatory disease that primarily influences cardiovascular diseases such as myocardial infarction, strokes and heart failure. The first event in atherosclerosis is the attachment of leukocytes to the endothelium and subsequently intruding into the intima. It is categorized as inflammatory due to the abundance of immune competent cells at the lesion sites. The high population of macrophages develops into foam cells, containing modified low-density lipoprotein (LDL). The release of certain growth factors from the inflammatory response causes the SMCs to dedifferentiate into the synthetic phenotype and migrate to the intima. A fibrous cap grows covers the atherosclerotic plaque. Dying and dead cells derived from the foam cells accumulate at the lesion site. A major concern with this phenomenon is the rupture of the fibrous cap, causing the interaction of various components initiating thrombosis, which could lead to infarctions [17]. Figure 1.9 illustrates the initiation and development of the plaque leading from a normal artery through to the stages of thrombosis [18]. The occlusion of the blood vessel caused by the plaque reduces blood flow and hence, nutrient and oxygen supply to the cardiac muscle which can lead to heart attacks and eventually death.

There are several treatment options for atherosclerosis dependent on the severity of the disease. For onset symptoms of the disease, lifestyle changes are necessary which encompasses dietary changes, physical activity and weight management. Other aspects such as stress management and smoking are also important. Medication such as blood thinners may also be provided to slow down the plaque build-up but can come with complications. If the plaque develops to a life threatening stage, surgery 8

is the only option. Coronary artery bypass grafting (CABG) is a highly invasive procedure that bypasses the occluded artery using a blood vessel obtained from an artery obtained from the chest or a vein from the leg. Percutaneous coronary intervention or angioplasty is a non-surgical procedure where a catheter with a deflated balloon is maneuvered from the femoral artery to the lesion site in the heart. The balloon is then inflated in order to break away the plaque or push it against the sides of the artery wall in order to enlarge the arterial lumen. In order to avoid recoil of the artery wall, a stent can be expanded at the site to keep the vessel wall open. Although highly effective, stents are associated with certain complications, namely, restenosis and thrombosis.

Figure 1.8 - Coronary artery disease prevalence by sex and age. Adopted from Mozaffarian et al (2016) [16]

Figure 1.9 – Development of the atherosclerotic plaque. Adopted from Libby (2001) [18]

1.2.1 Mechanism of Restenosis Restenosis is a complication that arises due to neointimal hyperplasia caused by proliferation of vascular smooth muscle cells which would obstruct the blood flow within the vessel wall. Figure 1.10 demonstrates the thick neointima following 6 weeks of stenting using a bare metal stent (BMS). The endothelial cell monolayer is crucial in preventing this formation through their anti-inflammatory properties. During coronary angioplasty, there is inevitable damage to the endothelium and causing platelet activation, cytokine and growth factor release which eventually causes fibrinogen deposition. This results in a decrease of thrombomodulin due to a variety of cascades and leads to the proliferation and migration of the underlying smooth muscle cells. The excessive ECM production causes the neointimal formation resulting in restenosis. Figure 1.11 illustrates a simple pathway of the mechanism involved in restenosis [19].

15x 30x 100x

Figure 1.10 - Neointimal formation on stainless steel bare metal stent. Adopted and modified from Tesfamariam (2008) [19]

Figure 1.11 - Mechanism of restenosis. Adopted from Tesfamariam (2008) [19]

1.2.2 Mechanism of Thrombosis Although the development of stents has led to reduced incidence rate of restenosis, the complication of in-stent thrombosis arose. Stent thrombosis is the build-up of a thrombus (blood clot) in the blood vessel and causes an occlusion, thus minimizing oxygenated blood flow to the target organ. Reports have indicated that cases with stent thrombosis have shown death in 20-48% and major myocardial infarctions in 60-70% of the patients [20].

The causes of in-stent thrombosis depends on the patient (medical history and medication usage), the lesion itself (dimensions, bifurcation, pre-stent stenosis %) and the PCI procedure (stent characteristics, DES usage, residual thrombus, etc.) [21]. This complication is prominent in patients with DES due to the drugs used. The drugs (paclitaxel and rapamycin) inhibit smooth muscle cell proliferation, but simultaneously restrict the proliferation of endothelial cells. Certain drugs on the DES could demonstrate pro-thrombogenic effects. Following stent placement, the regular pathway leads to re-endothelialisation of the damaged endothelium through proliferation and migration of intact neighbouring tissue. With the DES, this endothelialisation stage is suppressed due to the mechanism of the drug. The exposure of the underlying layer to the blood flow induces a mechanism that attracts platelets to the injury site and resulting in eventual aggregation and the formation of a thrombus.

1.3 Stents: Current and Future State Coronary angioplasty presents the complication of restenosis (closure of vessel) and has been reported to have a 40% chance of occurrence [22]. There was a need to reduce this high rate of incidence and therefore the use of stents has been introduced in conjunction with the angioplasty procedure in order to physically hold the vessel open. Although restenosis still occurs, this rate has then dropped down to 25% [23].

The medical term “stent”, originally derived from Charles T. Stent, a dentist in the 19th century who developed a dental impression material had the definition slowly modified through the last century. The “stent” material was used to affix skin grafts during the First World War. The word transitioned into the meaning of “any kind of non-biological support used to give shape or form to biological tissue”. Although, the stent was established as a tool for percutaneous insertion, the word itself did not appear in literature until Dotter published an article reporting on expandable nitinol stent grafting in 1983 [24].

The patent regarding balloon-expandable stents was filed in 1985 by Julio Palmaz and had been downsized for use as a coronary stent. The first FDA approved coronary stents were the Cesare Gianturco and Gary Roubin developed stents under Cook Inc. It is stated that by 1999, more than 80% of all PCIs used the balloon expandable stent resulting in a high level of success [25].

After cardiac catheterization/angiography and the need for a percutaneous coronary intervention has been established, a catheter with a deflated balloon would be inserted through the femoral or radial artery. After reaching the region of the coronary blockage, the balloon would be expanded in order to increase the diameter of the blood vessel by pushing the plaque into the arterial wall, and hence allowing improved blood flow to the cardiac muscles. With more severe blockages, coupling of a stent with the angioplasty procedure would be used, where a crimped stent covers the balloon and the balloon as well as the stent is expanded at the site of blockage. The balloon is then deflated and withdrawn, and therefore leaving the expanded stent in place.

1.3.1 Ideal Stent Design As coronary stents can be split into various categories, designing an ideal stent cannot account for the qualities for all subtypes as conflicting properties exist. There are, however, certain key features that stents should incorporate into the design. There are several reviews that indicate the different qualities of an optimal stent. Table 1.1 lists parameters that need to be considered when developing a stent.

Ako et al. (2007) [26] based their review of stent design criteria on 3 underlying principles; deliverability, efficacy and safety. Aspects of their criteria may show contradictory properties but should have compatible aspects from each category and these certain aspects are shown in Figure 1.12.

Table 1.1 - General ideal stent property considerations (derived from Regar et al. (2001) [27])

General Stent Features Complete access to any vessel in the coronary system High Trackability High Flexibility Distributed Radial strength (prevent arterial collapse) Good Rigidity Minimal plaque protrusion through the struts Side branching capability Reduce inflammatory response Inhibit neo-intimal hyperplasia Thrombo-resistant Should not inhibit wound healing Biocompatible Biodegradability/bio-absorbable – avoid late stent thrombosis Therapeutic tool (Drug Delivery capabilities)

Figure 1.12 - Venn diagram illustrating the key aspects described by Ako et al. (2007)

The primary challenge for an optimal stent is the ability to be delivered with ease without neglecting the physical properties and without compromising on the therapeutic aspects; inhibiting smooth muscle proliferation whilst maintaining thrombosis resistance and anti-inflammatory properties. Endothelialisation should also be unaffected [28].

The properties of the material govern the designs of the stent. Most importantly, Young’s modulus, yield strength, ultimate tensile strength and strain. Yield strength is the required stress in order to initiate plastic deformation of a material. At any force below the yield strength, defines elastic deformation. The ultimate tensile strength is the property that defines the maximum stress the material can withstand before failure. The dynamic properties of a material include the Young’s Modulus and strain, which are used to define the stiffness of the material. Young’s Modulus is the stress-strain relationship and the higher the value indicates that there is a high stress to strain relationship, demonstrating that there will be a smaller elongation and therefore being classed as a “stiff” material. A general stent material would have a linear stress-strain relationship up until the point of yield strength, at which point, the material will start to plastically deform. The elongation of the material defines the percentage of deformation it can endure before complete failure.

As the material is the base aspect, the next stage is the actual three-dimensional structure of the stent. The properties of the actual stent are based on hoop strength, radial resistive force and chronic outward force. Hoop strength is the required compressive force in order to collapse an expanded stent, which effectively is the force that surpasses the yield strength of the stent and they yield strength is said to provide the stent with the adequate structural integrity. The expandable balloon should provide a force greater than the hoop strength of the collapsed stent in order to undergo the required plastic deformation. Self-expanding stents use the radial resistive force and the chronic outward force

properties for their deployment and design. Hoop Strength is not applicable to self-expanding stents. The radial resistive force is the amount of resistance the stent exerts against an external radial compressive force (radial arterial force). The chronic outward force is the force created by the stent against the vessel wall as it is expanding to the desired size [29].

1.3.2 Bare Metal Stents (BMS) Bare Metal Stents were the first stage in combatting the issues observed with coronary angioplasty. Initially, they were primarily composed of either stainless steel or cobalt chromium alloys and later, nitinol (a nickel-titanium alloy). Due to the required flexibility, biocompatibility, corrosion resistance, expandability and radiopacity of the material, the options available become very limited, and the cost effectiveness needs to be considered as well.

The deployment mechanisms of these stents are either self-expandable or balloon-expandable, where the latter method requires the material to undergo plastic deformation and must have the ability to preserve the shape and position once expanded. The self-expandable stents requires a certain degree of elasticity in order to be compressed during the delivery stage and then expanded once it has reached the desired location [30].

The stainless steel 316L alloy has been the most commonly used metal platform for stents due to the model properties it exhibits. 316L encompasses the ideal properties of strength, corrosion resistance and ductility due to the stable austenitic form [31]. Despite being used as a popular stent material, there are several disadvantages to using SS. The poor fluoroscopic visibility and MRI compatibility due to the strong ferromagnetic nature of SS makes it difficult to track within the body. This can be countered by adding markers to the stent (Gold, platinum, tantalum) at certain points of the stent [32]. Table 1.2 displays the various compositions of stainless steel alloys that could be used for stents. The relatively high nickel content is worth noting as there are significantly high populations (20% of women and 4% of men) that are allergic to nickel and could show signs of contact dermatitis. A high nickel percentage has shown teratogenic and carcinogenic capabilities. It is considered to be safe as long as corrosion does not occur and high concentrations of the ion do not interact with the tissue.

It had been shown that the inflammatory response caused by stainless steel, which leads to a fibro- proliferative response around the implant can also be observed in restenotic tissue when the arteries have been stented [31]. Studies have shown that the combined nature of blood, proteins and salts coupled with mechanical stress can influence the release of the sensitizing metal ions, where both organic and inorganic species can bind to the ions [33]. Studies have been conducted on identifying a correlation between in-stent restenosis and metal ion release from BMS and although the studies were not wholesome enough to provide an absolute conclusion, the studies do indicate there is evidence of

metal ion release into the vascular wall [34,35]. Saito et al. (2009) [36] conducted a study that resulted in strong data suggesting that nickel ion release from the 316L SS alloy has a direct influence in chronic refractory in-stent restenosis and that this information should lead to development of improved biocompatible materials for long term safety.

This led to the developmental concept of a nickel-free stainless steel stent using nitrogen as the alloying element and has been reported to show a better corrosion resistance. The properties, of a Ni- free SS alloy (F2229) can be seen in Table 1.2 and shows higher yield strength and ultimate tensile strength as well as a marginally better elongation percentage [37]. An in vivo study by Fujiu et al. (2012) [38] was the first to use Ni-free stainless steel as a stent. Their results show a significant reduction in neointima formation compared to the 316L BMS and therefore a clear conclusion that Ni2+ ion release contributes to activation in SMCs by stimulating inflammatory processes. This would overcome any possibility of hypersensitivity in nickel-allergic patients as well as reduce the rate/extent of restenosis caused by BMS. Research by Inoue et al. (2014) [39] also realized the potential of nickel- free high nitrogen stainless steel as a stent platform and demonstrates that endothelialisation was greater on the Ni-Free SS compared to SS316L. They also showed that the inflammatory response exhibited was almost 5 times greater in the SS316L than in the Ni-Free SS. This reaffirms the possibility of using a nickel-free stent material in order to combat the high rates of restenosis seen in conventional BMSs.

Table 1.2 - Properties of different Stainless Steel alloys (obtained from Poncin & Proft (2003) [37])

Metals Composition Density Young's Yield UTS Elongation (g/cm3) Modulus Strength (MPa) (%) (GPa) (MPa) SS 316L Fe65.5/Cr18/Ni14/Mo2.5 7.95 193 340 670 48

SS F1586 Fe63/Cr21/Ni10/Mn3.5/M 7.9 195 430 740 35 o2.5 SS F1314 Fe61/Cr22/Ni13/Mn5 7.88 193 448 827 45

SS (Ni- Fe55/Mn23/Cr21/Mo1 7.63 190 607 931 49 Free) F2229

Research into stent material development progressed and since cobalt alloys have already been well established for previous implantable devices since the 1930’s, it had been considered as a stent material due to its properties and biocompatibility. The trends in materials used in bare metal stents have shown a transition to Co-Cr alloys from stainless steel. Table 1.3 highlights the properties of the Cobalt-Chromium alloys used in stents along with their compositions. Comparing the properties of the commonly used SS316l alloy with that of Co-Cr alloys, it is apparent that the mechanical properties are 15

superior to that of SS. These properties help counter some of the limitations observed within SS stents. The ferrous nature of stainless steel corresponds to the poor MRI-compatibility and thus, the negligible Fe content of the cobalt chromium alloys provides a higher radiopacity. The improved density allows for designing stents with thinner struts with a good level of radiopacity. The strength of the Co-Cr alloys also delivers a significant advantage over SS alloys as thinner struts can be used while ensuring sufficient radial strength. Figure 1.13 clearly shows the improved mechanical properties through the stress-strain graph [40]. Thinner struts are a desired quality for stents as it has shown reduced SMC proliferation and therefore lower rates of in-stent restenosis [41]. The deployed stent would damage the endothelial lining, exposing the underlying SMC layer; hence why thinner struts are idealized as there would be a less damage, resulting in a lower ISR rate. Thinner struts also allow for improved flexibility when reaching the point of lesion. The MP35N alloy shows a much higher Ni content than the other alloys (including SS) and thus could cause the nickel based hypersensitivity and inflammation issue mentioned earlier.

Table 1.3– Properties of different Cobalt-Chromium alloys (sourced from Poncin & Proft (2003) [30].

Alloy Composition Density Young's Yield UTS Elongation (g/cm3) Modulus Strength (MPa) (%) (GPa) (MPa) L605 Cr20/W15/Ni10/Co55 9.1 243 1000 950 50 MP35N Cr20/Ni35/Mo10/Co35 8.43 221 414 930 45 Phynox Co40/Cr20/Fe16/Ni15/ 8.3 221 450 950 45 Mo-Mn7

Figure 1.13 – Stress-Strain graph of a Cobalt alloy and 316L Stainless Steel Alloy. Adopted from Medtronic (2003) [40]

The nickel-titanium alloy, termed Nitinol originated from the Naval Ordnance Laboratory in the late- 1950, discovered by William Buehler and Frederick Wang. The material was only revealed publically in 1962. The unique shape memory properties drove the commercial interest of the material in various technological fields [42]. The observed shape memory ability of nitinol makes it ideal for the concept of self-expandable stents, where they are compressed into the delivery catheter and released in the target location and allowed to expand against the wall of the artery. The superelastic nature (with a flexibility of 10-20 times greater than stainless steel) makes it a desirable material option for stents [43]. Medical grade nitinol consists of 55wt% nickel and 45wt% titanium.

Organic material within the body exhibit a very different elastic behaviour as seen on a stress-strain curve, when compared to stent material such as stainless steel and cobalt alloys. Figure 1.14 illustrates the elastic behaviour in the form of a stress-strain curve of natural tissue in comparison to nitinol as well as steel with nitinol [44]. The elastic deformation for steel is limited to 1% strain as opposed to the 10% strain achievable by natural tissue. As the stress is released, the strain is recovered at lower stresses. This mechanism is defined as the elastic hysteresis. Nitinol portrays a similar behaviour and due to the fact that the material can recover elastically from more than 10% strain, it can be classed as superelastic. Another important characteristic of nitinol is the ability to recover any deflection thermally if the temperature is changed during the load cycle and this ability is termed as “Shape memory”. For nitinol, superelasticity can occur up to 50°C above transition temperature. If the temperature exceeds this range, nitinol will start to lose the recoverability and will behave like a regular metal. The mechanism itself is called thermoelastic martensitic transformation.

Figure 1.14 – Stress-strain curves of natural tissue, nitinol and steel. Adopted from Cohen-Inbar (2012) [44]

Below the transition temperature, the material is in a martensitic phase, which can be deformed easily, after sufficient heating above the transition temperature, the material would recover its original shape and a phase transformation would occur from martensitic to austenitic [45]. The massive shift in properties due to the phase change can be seen in Table 1.4.

These unique properties allow for the production of self-expandable stents and are manufactured slightly larger than the target vessel diameter. They are crimped and placed into the catheter at a temperature on or lower than room temperature. Once the stent reaches the transition temperature 17

of 30°C, it will attempt to expand, but in order to prevent early expansion, it is fixed within the sheath until the target location is reached. It is then released and expands into the vessel wall.

As mentioned previously, the nickel content of a material may be a significant factor regarding inflammation but Nitinol is a self-passivating material as it forms a stable oxide layer which protects the bulk material from corrosion. The high nickel content in Nitinol has brought about concerns, but studies have reaffirmed that the nickel ion release is significantly lower than the daily dietary limit [46]. These studies were generalized and were not stent/vascular specific.

Table 1.4 - Properties of Nitinol phases (obtained from (Mani, et al., 2007) [47])

Nitinol Density Young's Modulus Yield Strength UTS (MPa) Phase (g/cm3) (GPa) (MPa) Austenite 6.7 83 195-690 895/1900* Martensite 28-41 70-140 *895MPa – Fully Annealed / 1900MPa – Work Hardened

Table 1.5 lists properties of metals that have been used in permanent implants and when these refractory metal properties are compared to that of the cobalt chromium or the stainless steel alloys; we observe a mismatch of material properties. Perhaps if these metals were developed into functional alloys, they might have a part to play in the stent industry. As it stands, the commercially available BMS consist of stainless steel and cobalt chromium alloys. Although not as common, the platinum alloy based stent is gaining traction

Table 1.5 – Properties of miscellaneous materials (obtained from (Mani et al., 2007) [47]

Metals Density Young's Yield Strength UTS Elongation (g/cm3) Modulus (GPa) (MPa) (MPa) (%) Tantalum 16.6 186 170 205 25 Titanium 4.5 107 200 300 30 Niobium 8.57 103 105 195 25 Tungsten 19.3 411 3000 3126 3 Molybdenum 10.2 324 1386 1540 15

Studies have indicated that the vascular response following BMS implantation, closely mimicked the wound-healing response, where the initial steps include platelet adhesion and thrombus formation, followed by fibrin plug formation and inflammatory cellular infiltration. T-lymphocyte penetration occurs after the 2 week mark and can occur for several months. Initial vascular healing occurring within the first 2-4 weeks is proceeded by SMC proliferation and matrix deposition. The study indicates complete re-endothelialisation and vascular repair three to four months following BMS implantation. 6-12 months later, the neointimal formation is at a maximum level, where neointimal volume decrease occurs due to collagen replacement [48].

Although Bare Metal Stents were a major contribution to the cardiovascular field, developments were needed in order to reduce the rate of restenosis even further. This led to the creation of Drug Eluting Stents (DES), which was used to counter neointimal hyperplasia and SMC proliferation using an embedded pharmaceutical approach.

1.3.3 Drug Eluting Stents (DES) The development of DES was propelled by the need for countering the issue with in-stent restenosis and thus, due to positive results, has shown the increase in regulatory bodies approving drug eluting stents into clinical use. The DES are based on three aspects; the stent platform, the stent coating and the drug itself.

The stent platform is usually a regular bare metal stent, and much like BMS design development, DES have switched from stainless steel to cobalt chromium alloys, due to the thinner strut design capabilities. There is constant development and progress in identifying a stronger, more anti-restenotic material for the DES platform. The stent platform must have all the sufficient mechanical properties mentioned with regards to the BMS design. Biodegradable materials are also considered for the stent platform.

The stent coating is the intermediate layer that facilitates the drug adhesion and delivery from the stent platform into the surrounding environments. This mediating layer controls the mechanism of release and the majority of DES uses synthetic polymers. These polymers are mixed with the chosen drug to form a drug-polymer matrix and placed over the stent platform. The drug release has been stated to be primarily through a diffusion mechanism. Biocompatible polymers have been used and resulted in minimizing thrombus formation. Studies are also trying to incorporate bioabsorbable polymers for drug release.

As the stent is inserted into the target site, a certain level of damage occurs followed by a healing response. Platelet agglomeration and thrombus formation occurs, followed by cell recruitment, where these cells produce certain factors which induce SMC activation. This activation causes proliferation and migration of SMCs into the intimal layer and thus causing ISR. The ideal pharmaceutical intervention applied to the stent must inhibit the SMC proliferation [49]. The drugs used for DES are identified in further detail later. Synthetic drugs are not always the solution as some studies shows that bioactive agents/biomolecules can be used to inhibit SMC proliferation/encourage EC proliferation.

1.3.3.1 First Generation DES The first series of DES to be approved was later termed as 1st generation drug-eluting stents. These consist of two different stents; the Cypher sirolimus-eluting stent (SES) (Cordis Corporation, Miami

Lakes, Florida) and the TAXUS paclitaxel-eluting stent (PES) (Boston Scientific Corporation, Natick, Massachusetts) released in 2003 and 2004 respectively.

Cypher consists of an electropolished 316L stainless steel platform, coated with an inert polymer (parylene C) and coated with a mixture of polyethylene-co-vinyl acetate (PEVA) and pol n-butyl methacrylate (PBMA) along with the dose of the drug, sirolimus. A final coating of the polymer mixture without the drug is applied in order to prevent rapid immediate leaching of the drug once implanted [47]. TAXUS stents uses a poly(styrene-b-isobutylene-b-styrene) (SIBS) copolymer mixed directly with the drug, paclitaxel, without the use of any intermediate or covering layer and also uses the 316L SS alloy as the platform material [50]. Table 1.6 displays the key specifications of the two first generation DES [51]. The drug release observed for the TAXUS stent indicates that SIBS is a more stable polymer and therefore controlling the release of the paclitaxel as compared to Cypher. Studies have tried to judge the “better” or more efficient stent but a superior product cannot be outright identified. They are both highly efficient in certain settings [52,53]. However, one particular study shows there was a greater propensity in restenosis rate and late thrombosis in TAXUS PES when compared to the Cypher stent [54]. This idea is reaffirmed by several other studies and randomized controlled trials [55–57].

Table 1.6 – Comparison of the first generation DES (obtained from Iqbal et al. (2013) [51]

Cypher TAXUS Manufacturer Cordis Boston Scientific Express Platform Bx-Velocity Express Design

Material SS 316L SS 316L Strut Thickness (µm) 140 132 Polymer PEVA,PMBA SIBS Drug Sirolimus Paclitaxel Drug concentration (µg/cm2) 140 100 Drug Release (4 weeks) 80% <10%

1.3.3.2 Second Generation DES The first generation of DES have used a stainless steel platform and as stent research developed, studies have shown that decreasing strut thickness would lower arterial damage and endothelialisation rates. In addition, first generation DES exhibited cases of local polymer and drug toxicities. The thick polymer coating also resulted in polymer delamination and non-uniform drug distribution [58]. The cobalt-chromium alloy used in BMS showed positive results with regards to maintaining stent strength whilst improving rate of EC coverage through the use of reduced strut thickness [59]. Clinical studies have also proven that a thin-strut structure lowers incidences of clinical restenosis than the thicker counterparts [60]. This information led to the development of the 2nd generation drug-eluting stents.

The Endeavor (Medtronic) zotarolimus eluting stent received the CE and FDA approval in 2006 and 2008 respectively. It uses a drug, based on sirolimus mixed with a phosphoylcholine polymer on a cobalt chromium alloy platform and releases 95% of the drug within 15 days of placement [61]. Another 2nd generation DES commonly used was the Xience-V stent, composed of a different cobalt chromium based alloy, allowing it to have thinner struts and thus improving re-endothelialisation rates. Table 1.7 summarises key differences between the two key 2nd generation stents [51].

Table 1.7 – Comparison of the second generation drug-eluting stents (obtained from Iqbal et al. (2013) [51]

Endeavor Xience-V Manufacturer Medtronic Abbott Vascular Platform Driver Vision Design

Material CoCr MP35N CoCr L605 Strut Thickness (µm) 91 81 Polymer Phosphorylcholine PBMA,PVDF-HFP Drug Zotarolimus Everolimus Drug concentration (µg/cm2) 100 100 Drug Release (4 weeks) 100% 80%

1.3.4 Current commercially used Drugs for DES The primary aspect in countering the restenotic effect seen with bare metal stents is the use of drugs. There are four categories of pharmacological agents that would aim at preventing or reducing the rate of in-stent restenosis; anti-neoplastics, immunosuppresives, migration inhibitors and enhanced healing factors.

Anti-neoplastic agents are primarily used to treat cancer by aiming to inhibit any tumour growth. They promote stabilization of tubulin into microtubules, and this stabilization would interfere with the cellular growth and therefore inhibit proliferation. Migration inhibitors inhibit the matrix metalloproteinase enzymes which prevents matrix degradation. A key factor required for cell migration. Enhanced healing factors are a class of growth factors and miscellaneous compounds that are specific for endothelial cell healing. Stents utilising these compounds are still under development. Immunosuppressive drugs are targeted towards prevention of an immune response against transplanted organs. They halt the DNA synthesis process which in turn stops the cell cycle route. Drugs in the -limus family are part of the immunosuppressive group.

Sirolimus (also known as rapamycin) was originally obtained from a micro-organism found in soil on Easter Island and was initially used as an anti-fungal antibiotic [62]. The lipophilic quality of the drug

ensures that release of sirolimus into the bloodstream is limited during stent placement and the mechanism of drug release favours the release directly into the tissue. Apart from the antibiotic role it possesses, Sirolimus is also an anti-proliferative and an immunosuppressant [63]. In vivo and in vitro Studies have shown that Sirolimus inhibits the proliferation of smooth muscle cells and therefore making it a viable candidate for drug eluting stent in order to prevent tissue hyperplasia [64]. The pharmaceutical mechanism of Sirolimus is initiated by binding to a cytosolic receptor called FK binding protein 12 (FKBP-12) and creates a complex which binds to the mammalian target of rapamycin (mTOR) which is a key growth factor and signal transduction kinase. This binding would inhibit the mTOR activation and would cause a rise in p27Kip1 levels and a reduction in cyclin-dependent kinase (CDK) complexes. The CDK–cyclin complexes are crucial to ribosomal phosphorylation and protein synthesis, which therefore leads to the interruption of the cell-cycle at the G1-S phase and therefore inhibits cell proliferation [65]. The mechanism is illustrated in Figure 1.15. Additionally, FKBP-12 upregulation seen in neointimal SMCs reaffirms the antirestenotic claims of sirolimus and also exhibits anti-inflammatory effects on the artery wall [66,67]. Tacrolimus, another immunosuppressant only binds to FKBP-12 and would halt the cell cycle through calcineurin inhibition [68].

Although the end result is the same, the antineoplastic drug, Paclitaxel uses a different mechanism to that of Sirolimus. Paclitaxel was identified in the extracts of yew trees in 1971 and is classed as a powerful microtubule stabilizer [69]. The pharmaceutical agent inhibits mitogen-activated protein kinase, which is an enzyme used for microtubule depolymerisation. Paclitaxel also promotes microtubule assembly and the effect of both pathways creates superstable microtubules. The abnormal stability interferes with other microtubule-dependent structures that are needed for mitosis and therefore directly stops the cell-cycle at the M-phase [70]. Figure 1.15 visualizes the basic mechanism of the effect of paclitaxel on VSMC proliferation. Studies have shown that sirolimus-eluting stents are marginally superior to the paclitaxel-eluting stents due to the lower rates of restenosis observed [71].

Figure 1.15 - Sirolimus and paclitaxel mechanism. Adopted from Fattori and Piva (2003) [72]

Zotarolimus – The phosphorylcholine coating has been shown to resist fibrinogen adsorption as well as reduced monocyte and platelet activation. The rapid drug release effect seen with the Endeavor Zotarolimus-eluting stent (ZES) has been reported to show signs of a reduction in local toxicity [73]. Studies have shown that patients with ZES exhibited higher revascularization, as compared to SES, which showed higher very late stent thrombosis risks [74].

Everolimus is an immunosuppressive agent derived from Sirolimus and much like the original agent, also inhibits mTOR. The primary difference between the two is half-life, where sirolimus is 60 hours, compared to the 30 hours of everolimus [75]. An article in The Lancet has stated that the cobalt chromium based, thin strut, Everolimus eluting stent, and with a fluorpolymer is considered the optimum product in terms of clinical performance. They reported that the permanent polymer present on the stent may cause improper vascular healing which could lead to late ST and neoatherosclerosis [76]. A possible way to combat the risk of late ST would be to use a biodegradable stent. A randomized controlled trial between Biolimus and Everolimus eluting stents displayed that following a one year trial showed that both DES have a low rate of target lesion revascularization and stent thrombosis [77]. Table 1.8 displays the outcome of clinical studies between a series of drug eluting stents. Based on this data, the most effective drug commercially used is Everolimus.

Table 1.8 – Comparison table of the drugs used in DES

Drug Sirolimus Paclitaxel Zotarolimus Everolimus Everolimus [78] Everolimus [79] Zotarolimus Sirolimus [80] Zotarolimus [81] Paclitaxel Sirolimus [71] 1.3.5 Drug Delivery Mechanism Figure 1.16 illustrates the various possible strategies involved in drug release from stents [82]. Currently, the majority of stents use a polymer coating, but there are certain products that use non- polymer based drug release. The time dependency of drug release kinetics is crucial, subject to the therapy proposed for the patient and therefore there is no ideal drug release strategy. The release rate and dosage volume are major consideration when designing a drug eluting stent.

Figure 1.16 - Diagrams of various drug release mechanisms. Adopted from Sousa et al. (2003) [82]

The Taxus stent uses a diffusive approach, where a non-biodegradable polymer (poly(styrene-b- isobutylene-b-styrene) triblock copolymer) is used as the paclitaxel housing system and the drug is released by diffusion [83]. The Cypher stent consists of drug free “membrane” layer of poly (n-butyl methacrylate) (PBMA) which controls the release from the drug containing layer (blend of poly (ethylene-co-vinyl acetate) PEVA and PBMA). Sirolimus would diffuse from the medium to the membrane layer and undergoes an initial burst release which slows down over time [84]. The Endeavour stent uses phosphorlycholine (PC) which allows for rapid Zotarlimus release, with total elution by week 2 after implant [85]. The Xience V stent uses a 2 layer coating consisting of a polymer base layer and a drug reservoir polymer layer. PBMA was used for the base adhesive layer followed by

poly(vinylidene fluoride-co-hexafluoropropylene) (PVDF-HFP) mixed with everolimus. As with the other durable polymer coatings on stents, the drug would be released via molecular diffusion [86].

The Infinnium stent is a multicoated stent, with several layers of different polymers in order to precisely control paclitaxel release. A top layer of polyvinylpyrrolidone (PVP) with no drug embedment is used as a protective layer against premature drug release. The second layer uses poly L-lactide-co- caprolactone which undergoes quick degradation and exhibits a fast drug release of 33%. The subsequent layer releases 30% paclitaxel at a medium release rate using poly DL-Lactide-co-Glycolide (75/25) and the basal layer uses a 50/50 poly L-Lactide for the remaining 36% at a slow release rate [87].

Table 1.9 - Factors affecting drug release from polymers (Derived from Raval et al. (2010) [88])

Basic properties of drug Effects Hydrophobicity/hydrophilicity Affects protein binding, local drug conc., aqueous solubility and tissue retention characteristics Diffusion characteristics Release kinetics Solubility within polymer Release kinetics Solubility within release media ↑ Solubility = ↑ drug release rate Properties of polymer used Thermal properties Degradation, drug release, hydrophobicity, drug solubility in biodegradable polymers Degree of crystallinity Affects drug solubility and water penetration within biodegradable polymers. Degradation and drug release. Biodegradability properties (if applicable) Degradation behaviour and time Processing parameters Coating process Coating properties and drug elution Solvent properties Residual solvent effects and merging of coating layers affects release kinetics Coating Design Drug:polymer ratio Drug carrying capacity of polymer and drug elution rate Thickness and composition of coating Drug diffusion through polymer Drug distribution within polymer and Initial burst effect and dissolution mechanisms concentration Microstructure of coating Exhibits process conditions ฀฀drug delivery kinetics Hydrophobicity of polymer Lowering diffusion phenomenon = drug kinetic regulation Drug-Free thickness (top layer) Lowering diffusion phenomenon = drug kinetic regulation Mechanical properties of coating/film Coating integrity, could cause inflammation/thrombosis and obstruct drug uptake Stent design Affects drug dose differentiation within target region

There are a variety of factors that could influence the kinetics of drug release from polymer encapsulation (summarised in Table 1.9). The property of the drug, the polymer and the coating parameters are key influencers that define how the drug would be released and should be of key consideration when designing drug releasing coatings. Despite the solid development into drug eluting stents to combat the in-stent restenosis issue, complications have presented themselves associated with DES.

There have been cases that reported the chance of the polymer or the drug causing localized vessel wall inflammation. The report shows inflammatory cells infiltration into the layers of the coronary artery as well as around the stent strut. The frequency of the occurrence is still only considered as rare. Acute stent thrombosis would occur within the first 24 hours after stent placement and is often caused by poor stent expansion and uncovered dissections. Subacute stent thrombosis would occur between 2 and 30 days after stent deployment and is considered a rare occurrence. Although rare, the end results of stent thrombosis are usually fatal. Studies have indicated that the major cause is poor patient compliance with antiplatelet therapy or insufficient stent expansion. Late stent thrombosis would occur 30 days after implantation and is now considered a well-known complication associated DES and is said to be caused due to poor or delayed endothelialisation. Stent under-expansion is not common but can be linked to in-stent restenosis and is crucial for optimal stent expansion to avoid such complications [89].

1.3.6 Surface Modification In the medical device industry, the need for surface modification is a crucial aspect. It is difficult to find a single material that would contribute to all the specific needs of the device and therefore in order to maintain certain mechanical properties of the device, surface modification would need to take place in order to provide biocompatibility and therapeutic aspects. The bulk properties would remain intact whilst the surface would allow for any required biological interactions such as cell adhesion, proliferation, anti-inflammatory responses, drug deliverability and others. Figure 1.17 visualizes the major categories of surface modification currently used for treatment of medical implants/devices.

Roughening is semi-randomised category of surface modification in order to increase the material’s surface area, which in turn would restrict cell movement but would improve cell attachment. Surface roughening would cause surface changes without altering the direct chemical properties. The review by Govindarajan and Shandas has detailed that both metallic and polymer substrate roughening can cause an affect cell attachment [90].

Figure 1.17 – Representations of various forms of surface modifications. Adopted from Govindarajan et al. (2014) [90]

A study by Khang et al. (2008) [91] compared flat, micron and sub-micron roughness titanium surfaces using different e-beam current densities and deposition thickness. Their studies demonstrate that the surface energy is altered without changing the chemistry of the surface and furthermore identified that the sub-micron roughness showed a substantial increase (400%) in the adhesion density of endothelial cells. The nanometre roughness also showed an increase (50%) when compared to the flat, smooth titanium surface. If properly integrated into stent technology, the adaptation of the material’s surface without altering the chemistry directly would greatly influence the device. The only obstacle would be to identify a means of capturing the cells onto the roughened surface under physiological flow conditions. Another study by Samaroo et al. (2008) [92] attempted a sub-micron and nanometre roughness on nitinol surfaces and also reaffirms that the EC density and monolayer formation had improved. Roughening techniques such as sputtering [93] and microblasting [94] with reactive ion etching have been used and have both shown improved cell attachment. However, TiN/TiO2 sputtering has shown lower endothelial nitric oxide synthase (eNOS) expression which has been stated to lead to EC dysfunction. The study also stated that metals seem to exhibit signs of lower eNOS expression levels. This information suggests the need for non-metal - cell interactions.

A study that involved creating a nanometre rough surface on polymer substrates (polyurethane) using grafting of different length polymeric chains also demonstrated enhanced EC adhesion and growth [95]. Oxygen/argon plasma deposition onto polymer substrates has also been shown to increase surface roughness and hydrophobicity, both of which would improve cell attachment. Similar to surface roughening, patterning involves no direct chemical alteration of the surface, but it does provide an organized, structured pattern to the surface. Certain patterning techniques are used to improve endothelial cell attachment which would promote anti-thrombotic effects and vessel wall healing. Since endothelial cells align themselves with the direction of blood flow, a directionally based

patterned surface could possibly enhance EC function. Research by Lu et al. (2008) [96] developed grooves (750nm to 100µm) on Ti surfaces using plasma dry etching micromachining techniques and have shown to promote EC function.

Polymer surface patterning is a lot more popular due to the variety of techniques available. Lithography can be used to form dots, pillars, grooves, ridges etc. and have shown to directly influence the way the cells attach and spread across the surface. Photolithography has also been studied and has demonstrated formation of surface patterning using photo irradiation. Polymer printing methods have also been a major focus in the material processing industry and have demonstrated highly detailed surface patterning. Techniques such as nanoimprinting, robotic deposition, ink-jet printing amongst others have been used [97].

The chemical modification category aims at modifying the surface to a certain extent as not to affect the bulk mechanical properties of the structure and can be adapted to both metals and polymers. Not only does this type of modification target improvement of biocompatibility of certain materials, but it also aims to improve certain mechanical properties. Plasma immersion ion implantation and deposition has been used for both metals and polymers and have shown to have improved blood compatibility [60,98]. Passivation is another technique that has been used for surface modification and a study uses it to remove exogenous iron from a nitinol surface through chemical dissolution. This forms a nanometre level thickness of homogenous titanium dioxide later is formed over the nitinol. The same group also worked with electro-polishing which uses an ionic dissolution with an electric current which alters the chemical properties of the stent. The process removes metal ions from the surface (imperfections), leaving a titan surface, which is easily susceptible to oxidizing into TiO2 [99]. Nitinol has also been studied under hydrothermal treatment in order to reduce nickel release from the metal surface [100].

Polymers have the propensity for more chemical modifications, depending on the groups available within the structure. Ammonia plasma treatment has been shown to improve EC adhesion and growth through acid-amine interactions at the surface [101]. Free radical surfaces of polymers have shown the ability to encourage protein attachment which could improve biocompatibility [90]. A highly relevant study suggests that grafting of polyethylene glycol (PEG) monoacrylates to polycarbonate urethane films exhibited hydrophilicity and therefore resisted platelet adhesion [102]. PEG has already been established as a key biocompatible material for implants and this platelet adhesion resistance could show a high potential in stent technology in order to combat thrombosis.

Surface coatings would modify the material’s surface without directly influencing the chemical composition but rather are attached physically. Dimethyl sulfoxide (DMSO) has been shown to be

nontoxic and inhibited tissue factor expression and inhibited VSMC proliferation and migration as well as prevented thrombotic occlusion [103]. DMSO could be applied to the surface through wet coating/solvent coating. Dip coating is a very popular technique and can be used to form super- hydrophobic surface to prevent clotting. Plasma vapour deposition or cathodic arc deposition can also be used to lay polymers onto various substrates.

Porous surfaces are a viable candidate to house certain volumes of drugs within the stent without the need for an intermediate link between the two such as a polymer matrix or a linkage. Prolonged etching of a surface has been known to form pores on polymer surfaces, as well as using photolithography and sandblasting. Nanoscale pores have been seen on aluminium coatings and stainless steel in acidic conditions which can be used for drug delivery [90]. The main challenge here is careful control of the drug release rate from the pores.

The use of pharmaceuticals or biomolecules on a medical implant relies on the type of “carrier” used. This is greatly influenced by functionalizing the surface of a normally inert material in order for interactions between the surface and the (bio-) pharmaceutical can occur. Covalent links are the most common form of interaction due to its stability. The versatility of polymers makes it an ideal candidate for a functionalized surface. Plasma treatment allows for the formation of reactive groups and is known to be more reproducible, with fewer side effects than wet chemical methods [104]. This highly configurable surface modification could be used for stents in order to improve endothelial cell attachment or prevent thrombosis.

1.3.7 Organic Coatings/Additives Apart from the bulk material used for stents, the defining factor in the biological sense relies on the surface parameters. The stent surface is the point of direct contact to the vessel wall and hence would determine the stent’s efficacy and the healing properties associated with the device. The surface parameters could include anything from mechanical modifications to biodegradable polymer coatings. The polymer coatings are designed to house pharmaceuticals or to provide an intermediate layer to allow for biomolecule anchorage as well as improve biocompatibility of the stent. Table 1.10 highlights major necessities for polymer coatings for use on the stent [89]. A research highlight based on a randomized controlled 2-year outcome trial observed that biodegradable polymer DES did not lower the incidence rate of very late stent thrombosis (VLST) when compared to durable polymer DES. Duarte suggests that VLST may not in fact be associated with the use of durable polymer stents [105]. On the other hand, a meta-analysis of randomized trials conducted by Zhu et al (2015) [106] suggested the opposite; as biodegradable polymer DES exhibited a lower risk of very late stent thrombosis as compared to the durable polymer based stent.

Table 1.10 - Optimum properties for polymer coatings for stents (derived from Moyer et al. (2007) [89]

Low profile – Easy stent deliverability Conform to vessel shape Withstand stent expansion without disturbing the coating Damage resistant within artery/during stent delivery Optimum surface area – Drug delivery Avoids localized inflammatory responses/hypersensitivity/fibrosis 1.3.8 Bio-stable Polymers Polyethylene Vinyl Acetate (PEVA) – poly n-butyl Methacrylate (PBMA) (structure in Figure 1.18) blend has been described as bio-stable and non-erodible. Since PBMA is the direct surface interfacing polymer used in the Cypher stent, haemolysis and thrombogenicity related analysis had been conducted and had passed several of these tests, enabling FDA approval. The PEVA+PBMA blend/Sirolimus ratio used is 2:1

Figure 1.18 – Chemical Structure of PEVA and PBMA. Adopted from ASM International [107]

Poly(styrene-b-isobutylene-b-styrene) (SIBS) (structure in Figure 1.19) is a hydrophobic thermoplastic triblock copolymer with a Mn of 80k-130k g/mol and has elastomeric properties. These elastomeric properties allow it to deploy without any mechanical disruption on the TAXUS DES [108]. A complication with SIBS is the high adhesive characteristics, requiring high forces needed for the withdrawal of the deflated balloon when deployed in curved areas [109].

Figure 1.19 – Chemical structure of SIBS. Adopted from Fittipaldi et al. (2015) [110]

Phosphorylcholine is a biocompatible polymer that has been shown to be non-thrombogenic due to the identical structure of the predominant phospholipid head group seen in the red blood cell membrane [111]. Poly(vinylidene fluoride-co-hexafluoropropylene) (PVDF – HFP) (structure in Figure 1.20 is an advanced biocompatible copolymer. The polymer has a hydrophobic surface which causes a response called fluoropassivation which reduces throbogenicity and fibrin adhesion and therefore improves the healing process [112].

Figure 1.20 – Chemical Structure of PVDF-HFP. Adopted from Hwang et al. (2010) [113]

The BioLinx polymer is comprised of a blend of 3 polymers; polyvinyl pyrrolidnone (PVP), C10 and C19 at a 64/27/10% composition respectively. The chemical structure of the individual components are illustrated in Figure 1.21. PVP is a hydrophilic polymer which is aimed at improving biocompatibility and to allow a burst drug release during the early stages. The C10 polymer is based on butyl methacrylate and provides hydrophobicity in order to contain the drug, zotarolimus. It ensures slow and sustained release. The C19 polymer is synthesized from hydrophobic hexyl methacrylate, hydrophilic vinyl pyrrolidinone and vinyl acetate and is used to improve biocompatibility (hydrophilic component) and drug release control (hydrophobic component) [112,114]. This information demonstrates the value of surface wettability with regards to biocompatibility and drug release profiles.

Figure 1.21 – Chemical structures of the components of the Biolinx polymer. Adopted and Modified from Cheng et al. (2009) [115]

1.3.9 Biodegradable Polymers Biodegradable polymers are of key importance in the field of tissue engineering and particularly within the drug delivery aspect. The versatility of synthetic polymers allow for highly controlled degradability. The key aspect that needs to be addressed is the level of biocompatibility as well as by-products of

degradation. Polyester-based polymers are extremely popular due to their susceptibility to modification of properties [116]. Copolymers and polymer blends allow for hybrid properties and ideally combining polymers with EC proliferative and SMC inhibitory capabilities will be essential for stent coatings.

A study by Busch et al. (2014) [117] examined the interaction of ECs, SMCs and platelets on various biodegradable polymer surfaces, using established polymers (PEVA and PBMA) as a comparison study. The research looked at PLLA, P(3HB), P(4HB) and a PLLA/P(4HB) blend under both static and flow conditions, where EC proliferation was less than the control (polyolefin coverslip) for the PLLA, PEVA, PBMA and P(4HB) samples. The PLLA/P(4HB) blend showed a 100% increase in EC proliferation compared to the control and also shows strong adhesion and structural organization. On the P(3HB) the EC proliferation is shown to be quite high but cell clusters are formed as opposed to cellular monolayers. Their data also suggests that the (4HB) monomer could be responsible for the antithrombogenic properties observed from low platelet adhesion counts. However, these properties only exist when the P(4HB) component is blended with PLLA and has yet to be understood. The established polymers, PEVA/PBMA showed poor EC proliferation rates but also showed lower signs of platelet adhesion and aggregation than the other polymers.

1.3.9.1 Poly(lactic-co-glycolic acid) Poly(lactic-co-glycolic acid) (PLGA) is a co-polymer comprising of poly(lactic acid) (PLA) and poly(glycolic acid) (PGA). Its biocompatibility and degradation kinetics has been attributed to suitability for drug delivery and hence has been well established amongst the biomedical research community. It has been approved by the Food and Drug Administration (FDA) for use in drug formulation and medical devices.

PLGA can be broken down by hydrolytic degradation to its respective polymers/monomers: glycolic acid and lactic acid as well as an acidic oligomer and a hydrogen ion. The chemical structure and breakdown can be visualised in Figure 1.22 [118]. The monomers can be cleared by natural pathways, where glycolic acid is converted to metabolites and lactic acid is removed through tricarboxylic acid cycle [119]. It is critical to understand the physio-chemical properties of PLGA in order to customise the polymer for controlled drug delivery systems. PGA is hydrophilic in nature with a highly crystalline structure and can undergo relatively fast degradation. PLA on the other hand exhibits a similar structure to PGA but due to the methyl group, is more hydrophobic and an increase in LA percentage within PLGA increases degradation time. The degradation rates are governed by the percentage of lactide and glycolide groups within the polymer (m:n %). The crystallinity of PGA is lost in PLGA but studies have reported that amorphous polymers are better suited for drug delivery applications due to the homogenous release of active species in the polymer [120].

Figure 1.22 – Chemical structure of PLGA and its respective by-products through hydrolytic degradation. Adopted from Talujah et al. (2007) [118]

PLGA undergoes bulk degradation hydrolytically in a stepwise manner. Hydration causes water to penetrate the amorphous regions of the polymer by attacking van der Waals forces and hydrogen bonding. This process lowers the glass transition temperature. The second stage calls for the cleaving of the covalent bond which causes the drop in the polymeric molecular weight. Progressive degradation then causes the carboxylic end groups to autocatalyse the degradation process and the backbone bonds are cleaved, lowering structural integrity of the bulk. The final step is the solubilisation, where the smaller polymeric fragments are cleaved to soluble end products and through the means of the Krebs cycle are released in the form of CO2 and H2O [121]. The degradation rate is controlled by numerous factors: molecular weight, lactide and glycolide composition, stereo- chemistry, functional end groups, bulk geometry, surrounding medium properties and chemical derivatisation [122].

Due to the nature of PLGA being an optimal tool for drug delivery, there have been a vast assortment of studies identifying drug release, ideal drug concentrations and mathematical models identifying the kinematics of drug dispersion [123]. The drug loading percentage is a crucial parameter that would have both, biological effects on the patient as well as a physiochemical effect on the polymer coating and thus, needs to be carefully optimised based on the properties and safe dosage levels of the drug. Qi et al. uses a 1-3 wt% doxorubicin as the drug loading scheme within electrospun PLGA fibres [124].

The crucial aspect involved with biodegradable polymers is the rate of degradation as that directly contributes to the rate of drug release. There are also various ways a polymer could degrade,

depending on the chemical structure as well as its bulk structure. Synthetic polymers would usually degrade hydrolytically, whereas natural polymers, such as polysaccharides and proteins are degraded via enzymatic reactions [125].

Passive hydrolysis is the most common method of degradation for synthetic polymers and can be influenced by various factors. The type of chemical bond that governs the polymer backbone is the most crucial factor, but can also be influenced by neighbouring functional groups as steric and electronic hindrances can affect the water infiltration. The hydrophilicity of end groups or the backbone are also crucial factors as an increased hydrophilicity would allow for further water penetration into the polymer. The pH of the environment around the polymer is also an important parameter due to the catalysis. Depending on the polymer, optimal degradation can occur anywhere on the pH scale. Adding a new monomer to the polymer chain would also affect the chemical properties of the polymer as glass transition temperature and crystallinity could be altered. A high crystallinity would decrease the rate of hydrolysis. The way the polymer hydrolytically degrades could be classed as bulk erosion or surface erosion, where bulk erosion keeps the size of the polymer but loses strength and structural integrity over time. Surface erosion causes polymers to keep their original shape but reduces in size. Surface erosion is more predictable (due to linear loss of mass at surface) and therefore is desirable for drug release as it is dependent on the rate of degradation.

1.3.10 Polymer Deposition Techniques The deposition of polymers, both durable and biodegradable can be performed using numerous methods. Simple techniques such as spray coating, dip coating, spin coating and solvent casting can be employed. These coatings generally require the polymer to be dissolved in a compatible organic solvent.

Dip coating is a strategy that involves dipping the substrate into a polymeric solution and allowed to dry after. This method provides an effective and homogenous coating of smooth thin films onto the substrate. The thickness of the film can be governed by adjusting dip cycles, concentration of polymer within the solution as well as hold time. This method provides a cheap and effective solution for polymer deposition [126]. However, depending on the shape and structure of the substrate, this method may use an excessive volume of the polymer-solvent solution and may be ineffective for complex smaller shapes.

Spin coating is better suited to planar surfaces where the substrate is mounted onto a spinning platform and the polymer-solvent solution is applied to the surface. The spinning of the platform causes the polymer to spread and coat the substrate evenly. The thickness of the coating can be

adjusted by varying the speed of the platform or the viscosity of the solution. Studies have suggested that the coated polymer may exhibit anisotropic orientation due to the spin [127].

Spray coating is the most commonly used deposition technique with regard to stent coating, based on a simple system consisting of spray nozzle and a pump. The spray nozzle shoots microdroplets of the polymer solution onto the substrate on a rotating mandrel (if a stent is to be coated). For homogenous coating, the nozzle moves along the surface during the spraying process and is often coated multiple times until the required thickness has been achieved. A key issue is that the spray cloud exiting the nozzle has a higher droplet concentration directly longitudinal to the nozzle, whereas the outer regions of the spray cloud host less polymer droplets. Commercially available stents use spray coating technologies for their coating procedure [128]. This coating method also demonstrates higher evaporation rates due to the rapid dispersion of the liquid to a spray cloud containing micro-droplets. This could pose problematic if the concentration of the polymer in the solvent is too high, causing rapid deposition before the spray has reached the substrate surface.

Solvent casting is a simple procedure relying solely on the substrate shape to ensure near-homogenous solution coverage of the substrate and the evaporation of the polymer over time. The coating thickness can be calculated using the surface area of the coating space and the volume of polymer used in the solution [129]. This low cost, quick process allows for homogenous coating of the substrate and has been actively utilised as a coating technique.

1.3.11 Biomolecules The cell surface antigen, CD34 has been shown to exist on the surface of endothelial cells, endothelial progenitor cells and hematopoietic progenitor cells. Bone marrow-derived endothelial progenitor cells have been shown to have the capacity to differentiate into mature endothelial cells [130]. Additionally, studies have indicated that circulating EPCs have the ability to relocate to the site of vascular injury, resulting in the promotion of endothelialisation [131]. This has led to the possibility of capturing EPCs directly from the bloodstream using anti-CD34 antibodies and allowing them to differentiate into a functional endothelial cell monolayer. The anti-CD34 antibody is classed as an igG2a antibody targeted towards class III epitopes of the CD34 membrane protein (observed on progenitor cells). Currently, only one company commercially (Genous Combo/Bio-engineered stents, Orbus Medical Technologies) employs this strategy. The antibodies are covalently attached to the metal surface via a polysaccharide base matrix. It is reported that the antibody concentration attached to the stent surface is 1µg/cm2 [132]. Depending on the primary material, there are various approaches in immobilizing the Anti-CD34 antibody including; peptide linkages to biopolymer, physisorption, covalent bonding using glutaraldehyde and research is still being conducted on identifying site-selective immobilization

methods [133]. Figure 1.23 provides a simple example of a technique used to immobilize the antibody to a POSS-PCU polymer using EDC-NHS as the chemical linker between the two amino groups [134].

Figure 1.23 – Concept of antibody immobilization onto a polymeric surface. Adopted from Tan et al. (2013) [134]

Although the anti-CD34 antibody is highly sought after in terms of a tool to enhance endothelialisation, there are known limitations with their usage. A study by Vasa et al. demonstrated that EPC functionality is impaired within patients with coronary artery disease. Furthermore, they also observed lower population of circulating EPCs [135]. It has also been reported that one in 250 CD34+ cells are endothelial progenitor cells [136]. In addition, CD34+ cells have shown the potential to differentiate into vascular smooth muscle cells [137], which would highly contradict the use of the CD34+ antibody as a stent strategy.

CD133 is considered as an alternative marker to use as studies have stated that it is more specific than CD34 markers. A study conducted by Li et al. evaluated the endothelialisation of immobilized anti- CD133 antibody. Figure 1.24 highlights the preparation technique used in their study, where the stainless steel stent was covered with a copolymer film, activated by (N-ethyl-N‘-(3- dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS). This allowed for the peptide bonding with the antibody. Their in-vivo studies showed complete endothelialisation in both their CD-133 stents as well as BMS, but more distinct cell arrays were observed in the CD133 stents. Their study was not clear, but showed that the EPC capture technology was possible with the use of anti-CD133 antibodies. There was no comparison with CD34, which was their initial reasoning for attempting an alternative EPC capture approach [138]. A previous study by Sedaghat et al. showed similar results as well as the occurrence of unspecific binding of mononuclear cells [139].

Figure 1.24 – Immobilization of Anti-CD133 antibody procedure. Adopted from Li et al. (2014) [138]

The Arginine-Glycine-Aspartate (RGD) tri-amino peptide sequence is known to enhance cell adhesion to synthetic surfaces. This sequence can be found in extracellular matrix proteins and bind to integrin receptors [140]. It has been reported than EC function is maintained by integrins. ECs have a certain integrin receptor that binds ECM proteins and is mediated by the RGD integrin-binding motif [141]. There have been several studies aimed at linking the peptide to various polymers; PEG hydrogels, self- assembled monolayers, RGD-methacrylate terpolymer [142] using approaches such as surface etching and polymer embedding [143]. Most of this research has been adapted towards vascular grafting in order to enhance endothelial cell adhesion. A study by Joner et al. successfully allowed adhesion of the cyclic RGD peptides to a nitinol stent without the need for a polymer intermediate phase. They used tetraphosphonate as an anchor to the titanium dioxide stent surface in order to irreversibly link the peptide. The study did show the capacity of improved re-endothelialisation using this concept in an in vivo setting.

The Arg-Glu-Asp-Val (REDV) peptide is found within alternatively spliced CS5 fibronectin domain and can be specifically recognized by integrin α4β1. The integrin α4β1 has been said to be selectively expressed on proliferating ECs. The REDV peptide has been reported to selectively promote endothelial cell adhesion and spreading over SMCs and platelets [144]. The recent indications of REDV peptide for enhanced endothelialisation have shown a great rise in research aiming to incorporate the peptide onto various substrates: successful covalent binding to glycophase glass and Polyethyleneterephthalate (PET) – mediated with polyethylene glycol (PEG) [145], poly(poly(ethylene glcol) monomethacrylate) (P(PEGMA))-modified titanium surfaces [146], Covalent attachment of the peptide to a zwitterionic poly (CBMA-r- BMA) surface has shown limited SMC proliferative and adhesion capacity and could be used on various substrates [147], self-assembled REDV-heparin onto a Ti surface with an intermediate polylysine layer [148].

A study shows a comparison of RGD, YIGSR and REDV peptides coated onto a polyvinylamine (PVAm)- coated PET using PEG as a spacer. Their results contrasted from previous work, where the REDV grafted

PEG-PET film demonstrated a lack of EC adhesion. The RGD peptide showed improved HUVEC adhesion as well as platelet adhesion as expected. The REDV peptide exhibited the expected lack of platelet adhesion, however, their study shows a lack of any adhesive effect on the HUVECs. The study suggests that this could be due to the density range used but the reason for the poor adhesion requires further insight into the exact signalling pathway. This study alone is not a suitable reason to disregard REDV peptides as a potential therapeutic strategy as several other research groups have shown the high value and efficiency of the peptide. In this instance, the YIGSR peptide is revealed to be the most promising sequence with regards to EC adhesion and minimising platelet adhesion [149].

The laminin β1-derived tyrosine-isoleucine-glycine-serine-arginine (YIGSR) peptide, unlike the RGD peptide, does not interact with integrin cell receptors, but with the laminin binding protein instead [150]. A study by Taite et al. (2008) [151] synthesized a polyurethane-PEG copolymer with incorporation of the YIGSR peptide with the capacity to generate Nitric Oxide. They observed a decrease in platelet adhesion and SMC growth, whilst simultaneously stimulating EC proliferation. The study also showed that the material itself did not encourage SMC adhesion. Another study using polyurethaneurea with PEG-YIGSR identified the enhanced endothelialisation effect of the peptide, through higher cell densities as well as hydroxyproline production, which indicates ECM production [152]. The studies have shown the versatility of the peptide coupled with various polymeric mediators with positive results. This could be extremely beneficial in terms of rapid endothelialisation. However, YIGSR has been claimed to be non-specific and is not exclusively selective towards ECs.

A strategy proposed by de Torre et al. (2015) [153] studies elastin-like recombinamers (ELR) covered stents which encompasses the ability of ELRs to include various bioactive sequences (REDV and RGD) for cell adhesion and proliferation. ELRS have been studied heavily for different applications due to their biocompatibility and versatility. In this study the ELRs were attached to the stent using click chemistry. The RGD-ELR stent exhibited quicker endothelialisation after 15 days of dynamic culture compared to the REDV-ELR stent. The study also demonstrated the high hemocompatibility of the ELR- modified stent, suggesting that this strategy could be a viable option as an intermediate phase for biomolecular anchorage. However, in vivo studies have not yet been performed and may exhibit different results due to various physiological parameters.

Omentin or intelectin-1 is an abundant adipokine expressed in human visceral fat tissue and has been observed to show lower concentrations in patients suffering from obesity-linked disorder. Studies have also observed the reduction of omentin concentration in patients with coronary artery disease and carotid atherosclerosis. Omentin has been reported to be anti-inflammatory in vascular ECs and SMCs, but the exact mechanism is still uncertain. Kazama et al. (2014) [154] reported inhibited platelet derived growth factor-BB-induced NADPH oxidase activation which in turn inhibited TNF-α induced

SMC inflammatory response. They concluded that omentin prevents SMC migration through the activation of p38/HSP27 proteins, which could lead to prevention of neointimal hyperplasia. This is reaffirmed by another study indicating the prevention of VSMC proliferation and neointimal growth following arterial injury through an AMPK-dependent pathway. The study has also reported improvement of endothelial function [155]. These early-stage studies shows that omentin could be a viable option for stent-based therapies in order to reduce restenosis, but it is still clear that further studies are required in order to evaluate the possibility of using omentin in vivo.

Although EPCs have been suggested to have various issues regarding cell population, late EPCs have a longer lifespan with improved proliferative and differentiation potency. A study uses stents coated with anti-vascular endothelial-cadherin (VE-cadherin) as it is expressed solely on late EPCs and exhibits crucial intracellular signalling. The stents were coated using silanized and PEGylated prior to antibody adsorption. The study compared the commercially available OrbusNeich anti-CD34 coated stent with the newly coated anti-VE-cadherin stent and observed improved re-endothelialisation and less restenosis in their in vivo results [156]. This does not divert away from the known limitations observed with endothelial progenitor cells with regards to inversely proportional functionality with age and disease severity of the patient.

1.3.11.1 Collagen Collagen has been gaining serious traction within the last few years in the medical field due to improvements in materials science. Collagen is a highly abundant family of fibrous proteins within the body as it is a major component of the extracellular matrix and makes up around 25% of the dry weight of mammals. The standard collagen format can be arranged into a triple helix structure from three polypeptide chains and exists in a multitude of conformations, each with different α chain geometries and combinations. Although there are 29 different types of collagen, only a few are currently used for tissue engineering purposes. Type I collagen is the current gold standard as it represents about 90% of the total collagen in the body [157].

Type I collagen is also regarded highly amongst the alternative natural polymers as only a small population have humoral immunity against it and a possible allergic reaction can be verified prior by conducting a simple test [158]. The availability of collagen will always remain high as it can be extracted from almost any living animal. With regards to tissue engineering, the most common sources are: rat tail, bovine skin and tendons and porcine skin.

The degradation of collagen is dependent on the method of crosslinking used or if specific agents applied. Matrix metalloproteinase (MMP) is a class of enzymes that causes the degradation of the collagen [159]. Vascular endothelial cells release these MMPs in order to regulate and remodel the

ECM [160]. Collagen has also been regarded as a valuable tool for drug delivery applications due to its degradability and biocompatibility. The drugs can be physically encapsulated or attached through hydrogen or covalent bonding [161]. Chen et al. demonstrated sustained release of sirolimus from a multi-layer type I collagen coated stent. The study used genipin as a crosslinking agent to further increase degradation time [162].

Collagen-cell interactions have been shown to be a key element in vivo as well as in vitro and have been thoroughly examined to show that collagen can enhance the proliferation and adhesion of different cell types due to its bioactive status. It also promotes further proliferation, migration and differentiation of cells [163]. A study by Relou et al. demonstrated that collagen, fibronectin and gelatin coatings on tissue culture plastic have all shown optimal endothelial cell growth [164].

The primary advantages of using collagen as a biomaterial are: i) no issue with toxicity or foreign body response, ii) biologically active at the molecular level and iii) naturally degraded and can be varied upon cross-linking modulation. These properties make the material a desirable choice for several tissue engineering and wound healing applications. However, natural polymers such as collagen do have their limitations: i) possible immunological reaction, ii) high variability and iii) complex structure and thus has limited methods of manipulation [165].

To overcome some of these limitations, the use of recombinant human collagen (rHCol) was becoming more popular. The batch variability and the possibility of non-human proteins possessing immunological agents of animal derived collagen have led to the production of recombinant human collagen obtained from non-animal sources such as yeast, milk and tobacco through gene transfection [166]. Table 1.11 highlights the primary differences of various crucial collagen-based parameters between animal-derived and recombinant collagen.

Table 1.11 - Comparison of animal-derived collagen and recombinant collagen (derived from Yang et al. (2004) [166]

Parameter Animal-derived collagen Recombinant Safety Prion and virus contamination risk No prion/virus risk Source Animal carcasses and tissue Genetically equivalent expression systems Variation Batch-batch variation with possible mixture Reproducible and of high purity – of other collagen types predictable performance Structure Multimeric derivation. Degree of crosslinking Monomeric tissue derivation dependent on animal age and tissue Processing Harsh treatment for collagen extraction Mild conditions for bioprocessing Customisation Limited Engineer-able to specific characteristics

The aforementioned benefits of rHCol make it highly attractive as a biomaterial for tissue engineering purposes. Studies have shown the capacity of recombinant collagen to perform in a similar manner to native collagen in terms of gelation and structural formation of membranes, gels and sponges. Studies have demonstrated cell adhesion, growth and differentiation of various cell types on rHCol-based scaffolds [166]. A type I rhCol wound dressing sheet was seeded with endothelial cells amongst other cell types and the study observed superior biological performance with regards to cell proliferation [167].

Although there is a lack of studies describing the potential of rHCol as a stent coating, the described benefits and results demonstrate a strong hypothesis for rHCol to replace native collagen as a biomaterial for stent coatings.

1.3.12 Screening of Potential Bi-functional drugs for DES applications The pursuit of drug discovery usually begins with 5,000 to 10,000 compounds and is a multi-million euro process, which can take over a decade for a drug to be introduced to market [168]. The lack of a HTS setup and a compound library led to the use of a literature-based discovery (LBD) to identify targeted therapeutic approaches for potential use within a drug eluting stent. The LBD paradigm was first introduced by Don. R. Swanson in 1986 [169]. The concept of LBD is based on the extraction of data from previously published literature to identify new characteristic connections of the knowledge for a new biomedical hypothesis. A schematic of the paradigm is illustrated in Figure 1.25 highlighting the principle. This diagram describes the process of developing a hypothesis between two unrelated published articles through the means of intermediate connections.

Figure 1.25 - Simple scheme of the Swanson LBD Paradigm. Adopted from Hristovski et al. (2013) [170]

With regards to drug eluting stents, literature was scanned using Swanson’s LBD method for therapeutic options that provide a positive or maintained effect on the endothelial cells whilst potentially limiting the extent of smooth muscle cell proliferation.

1.3.12.1 Magnolol The bark of the Mangolia has been utilised for its antibiotic, antioxidative and anti-diabetic properties as a traditional Chinese medicine. It has been reported to be an important treatment for cardiovascular disorders. Mangolol, a polyphenol, is the active compound isolated from the bark and the chemical structure can be visualised in Figure 1.26.

Figure 1.26 - Structure of Magnolol. Adopted from Wu et al. (2006) [171]

Several studies have looked at the effect of magnolol on vascular smooth muscle cells. Chen et al. observed that magnolol induces an apoptotic effect on the SMCs and hence can limit the proliferative capacity in a dose-dependent manner through the downregulation of a mitochondrion- associated protein, Bcl-2 [172]. Inhibition of vascular cell adhesion molecule (VCAM-1) and monocyte chemoattractant protein (MCP-1) expression was also demonstrated, which has been claimed to be valuable for the treatment of atherosclerosis and restenosis. SMCs have the ability to release these inflammatory molecules, which in turn promote the process of atherosclerosis [173]. Smooth muscle cells treated with Magnolol at concentrations of 200µM demonstrated a reduction in SMCs shifting to the S-phase (DNA synthesis) and can be attributed to NF-κB downregulation and an upregulation of caspase-3 [174]. NF-κB expression has been shown to be higher in atherosclerotic SMCs and therefore suppression of this protein can lead to the reduction in SMC proliferation. Kim et al. also observed the inhibitory effects of Magnolol at 5-20µM on TNF-α-induced SMCs through reduction of extracellular signal-regulated kinases 1/2 (ERK1/2) activity. ERK1/2 is linked to cell growth and migration [175].

Chen et al. treated human aortic endothelial cells with up to 60µM of magnolol and observed a drastic reduction in viability after reaching the 30µM threshold. They also observed the inhibition of TNF-α induced NF-κB activation [176]. Magnolol was also shown to protect vascular endothelial cells by inhibiting intrinsic apoptosis onset by oxidised-LDL. The drug removed intracellular reactive oxygen species (ROS) and thus the NF-κB activation was inhibited as previously reported in 2002 by Chen et al. [177]. This suggests a trend in the potential of magnolol as a clinical tool to prevent atherosclerosis.

1.3.12.2 Curcumin Curcumin is a polyphenolic compound found in the root of the turmeric plant. Research over the past decades has shown that it possesses anti-bacterial, anti-fungal, anti-viral, anti-inflammatory, anti- proliferative and anti-atherosclerotic properties. It is also known as diferuloylmethane and its structure is illustrated in Figure 1.27.

Figure 1.27 - Structure of Curcumin. Adopted from Hatcher et al. (2008) [178]

Studies have indicated that curcumin has the capacity to inhibit the proliferative state of vascular smooth muscle cells. Chen et al. reports that 100 µM of curcumin induces cell apoptosis amongst vascular smooth muscle cells. This could potentially be a means to eliminate the excessive proliferation of cells observed during the progression of restenosis. The anti-proliferative effect of curcumin on vascular smooth muscle cells was identified and showed no effect on cell viability when a concentration of 1-10 µM was used. Significant reduction in cell population was observed when 100 µM was used, indicating cell cytotoxicity. Their study also demonstrates that using 10µM displays an anti-proliferative effect on the smooth muscle cells, without cells entering into the apoptotic phase [179].

The anti-oxidative profile of curcumin allows it to neutralise reactive oxygen species such as nitric oxide, superoxide and hydroxyl radicals, which are influential factors towards damage to endothelial cells through oxidative stress. Han el al. demonstrated the curcumin’s capacity to induce autophagy at concentrations up to 50 µM and observed improved cell viability of cells exposed to hydrogen peroxide. Concentrations over 10µM showed no added benefit on H2O2-treated cells without reducing viability [180].

An interesting study by Guo et al. (2015) [181] demonstrated the effect of curcumin on rapamycin (anti-proliferative drug) – treated endothelial cells. Curcumin was shown to negate the apoptotic and anti-migratory effects seen when treated only with rapamycin. The study shows no observable difference when compared to their PBS control when only using curcumin, up to a concentration of 40µM. Increasing concentrations of curcumin on the rapamycin-treated cells showed cell recovery up to a dosage maximum of 30µM.

1.3.12.3 Ferulic Acid Ferulic Acid (4-hydroxy-3-methoxycinnamic acid) is a phenolic acid present at high concentrations in the cell walls of various plants (structure displayed in Figure 1.28. Plants containing ferulic acid have been used in Chinese medicine for centuries for treating strokes, hypertension, cardiac arrhythmias and other cardiovascular issues [182,183]

Figure 1.28 - Structure of Ferulic Acid. Adopted from Mori et al. (2013) [184]

Wang et al. assessed the effect of ferulic acid at various concentrations on the proliferation of HUVECs. They observed improved cell proliferation within the 0.5-50µM concentration range through means of a cell counting assay. They also deduced the increase in cells in the S phase and a decrease of cells in the G1 phase, which indicates the possible transition of the HUVECs from a quiescent state to a proliferative state. This was reaffirmed by the upregulation of cyclin D1 and VEGF mRNA levels [185]. Nicholson et al. analysed the effect of polyphenolic compounds on the gene expression in HUVECs and observed that eNOS gene expression was unchanged after 24h with a concentration of 0.1µM [186]. Another study also demonstrated the lack of eNOS expression change using a concentration range of 1-33µM [187].

An early study assessed a single concentration of around 283 µM with no avail to the inhibition of SMC proliferation [188]. Ferulic Acid treatment on aortic smooth muscle cells induced by angiotensin II using a concentration range of 0 – 160 µM was assessed. Inhibition of cell proliferation was observed in a dose-dependent manner with 80µM demonstrating maximum inhibition [189]. Vascular SMC proliferation caused by angiotension II induction is largely due to the mitogen-activated protein kinases pathway (MAPK). Their study described the effects of ferulic acid on these MAPK pathways (ERK1/2, JNK and p38) and demonstrated the inhibition of the proliferative state of vascular SMCs caused by angiotension II through the blocking ERK1/2 and JNK activation. This suggests that ferulic acid might be a strategic inhibitive tool to reduce the excessive proliferation of smooth muscle cells in stented arteries.

1.3.12.4 Exendin-4 Exendin-4 was initially observed to be an optimal therapy for type II diabetes due to its role as a glucagon-like peptide 1 (GLP-1) receptor agonist and hence, its ability to lower blood glucose through

the stimulation of insulin release from the pancreas [190]. There have been several studies on the cardiovascular effect of GLP-1 including modulating blood pressure, improving the survival rate of myocardial infarction in mice, produces anti-inflammatory effects in endothelial cells [191]. Exendin-4 is a 39 amino acid with the sequence: His-Gly-Glu-Gly-Thr-Phe-Thr-Ser-Asp-Leu-Ser-Lys-Gln-Met-Glu- Glu-Glu-Ala-Val-Arg-Leu-Phe-Ile-GluTrp-Leu-Lys-Asn-Gly-Gly-Pro-Ser-Ser-Gly-Ala-Pro-Pro-Pro-Ser-

NH2.

Hirata et al. (2013) [192] observed a significant reduction of neointimal formation upon vascular injury in mice administered with exendin-4. They also saw supressed proliferation of PDGF (platelet-derived growth factor) -BB induced smooth muscle cells in vitro. Another study identified that exendin-4 inhibited the proliferation of angiotensin II-induced rat aortic SMCs at maximum inhibition using an exendin-4 concentration of 10nM [193].

Exendin-4 was also shown to increase eNOS activity and protein levels in HUVECs upon 5nM/L treatment for 48h at 10nM [194]. A study by Erdogdu et al. (2010) [195] demonstrates the promotion of endothelial cell proliferation of human coronary artery SMCs with 10nM treatment of Exendin-4. The promotion of endothelial cell proliferation invoked by exendin-4 was attributed to the PKA- PI3k/Akt-eNOS pathway through the GLP-1 receptor-dependent mechanism. A correlation between exendin-4 dosage and eNOS activation was observed and hence the increase in NO production. NO is a critical factor reported to be a key regulator of endothelial cell growth and function, directly associated with the P13K-dependent pathway [196]. The same pathway activated by vascular endothelial growth factor [197] .

1.4 Bioreactors With the constant advancements in tissue engineering for clinical applications, several studies have demonstrated the value of the importance of mechanical stresses under flow conditions on the proliferation of cells and hence, the development of the whole tissue. The principle of tissue engineering relies on three factors: i) cells, ii) scaffolds and iii) a signal – mechanical, chemical or electrical. The mechanical signal is where a bioreactor is needed. Bioreactor systems are an approach that subjects the biological samples to a dynamic environment. A bioreactor is generally used as a system to simulate physiological cues for the conditioning, development and assessment of cells, tissue and scaffolds as well as organs in vitro [198]. Bioreactors should fulfil one of the following functions: i) conditioning and maintenance of culture medium, ii) even cell distribution onto a 3D scaffold, iii) efficient mass transfer to the 3D tissue, iv) assessment of the formation of the developing tissue and v) exposure of the tissue to physical stimuli. The primary goal is a suitable system that develops functional tissue for the purpose of in vivo translation.

Watanabe and Okada [199] observed the importance of temperature on the growth of cultured mammalian cells. Cell viability, growth, metabolism and protein synthesis are all affected by changes in temperature. Temperatures below 37° reduces cell growth and viability. The body temperature can rise to 43°C or drop down to 32°C but can be fatal at these extremes. Bioreactor systems are typically designed to either fit entirely within an incubator where the temperature is set at 37°C. Larger systems will have a heat exchanger adapted into it within the flow system of the bioreactor.

There are several different types of bioreactors where each system provides a specific effect on the cell and are highly application-based. The systems all rely on surrounding/filling the sample with media. A spinner flask bioreactor uses the concept of suspending the scaffolds within the media and use a magnetic stirring system to mix the media and provide nutrient and waste transport to the fixed samples as visualised in Figure 1.29. This simple system has an array of issues such as flow turbulences, inadequate mass transport and therefore heterogeneous cell distribution.

Figure 1.29 – Schematic of the spinner flask bioreactor system. Adopted from Yeatts and Fisher (2011) [200]

Rotating wall bioreactors is a NASA based system, originally intended to minimise forces on cell cultures during aircraft launch. This system allows for free movement of the scaffold in the respective media. The inner and outer walls of the cylindrical can be rotated at constant speeds. A simple scheme of the system is shown in Figure 1.30. The speeds and direction of wall rotation compensate for the downward gravitational forces and result in a suspended scaffold. The system has to be stopped temporarily for exchanging the culture media.

Figure 1.30 – Schematic of the rotating wall bioreactor. Adopted from Chen and Hu (2006) [201]

A compression bioreactor is a commonly used system for bone and cartilage tissue engineering. The system provides dynamic loading, similar to a physiological system. The system relies on a motor providing a linear motion in accordance to controllable displacements provided by an input system. The compressive nature also causes fluid flow throughout the scaffold [202]. Similar to the compression bioreactors, certain applications require the addition of strain to their samples. Tendons, ligaments and bone are amongst the first tissues that come to mind that might benefit from the application of tensile strain to the scaffold. A hydrostatic pressure bioreactor utilizes a pressurise-able chamber and an actuator-controlled piston that allows compresses the fluid in order to provide a mechanical stimuli to pre-seeded scaffold [203].

Flow perfusion bioreactors use a pump based system to circulate medium through the culture vessel and has been reported to provide homogenous cell distribution. The system allows for fresh feeding of circulating media from a reservoir and enables removal of waste products. Most perfusion systems are designed to be very application-specific and are mostly bespoke. A model perfusion system is illustrated in Figure 1.31. The fluid flow system tends to get complicated as it is essential to optimise a balance of adequate medium flow rate without causing a negative effect on the cells and newly synthesised ECM and thus fluid shear stress needs to be accounted for [204]. The thickness and choice of scaffold also plays an influential role in the performance of the tissue regeneration. Thick samples with poor porosity will not allow for appropriate levels of mass transport to supply the entirety of the tissue with nourishment [205].

Figure 1.31 – Schematic of a model perfusion flow system. Adopted from Barron et al. (2003) [206]

1.4.1 Vascular Bioreactors As stated previously, the biomechanical cues for tissue culture are tissue-specific. Vascular bioreactors would replicate the physiological cardiac output which is represented by the stroke volume. In an adult, the stroke volume ranges from 70 – 100 ml/beat and with a heart rate range of 75 to a 100 bpm. Depending on what vessel the vascular bioreactor is working with, an appropriate volume/minute can be deduced and can correspond with the pump flow rate. Lyons and Pandit (2004) [207] underline the key physiological parameters involved when designing a vascular bioreactor. Table 1.12 notes these key factors.

Within the vascular system it is imperative that blood is perfused with oxygen at the same rate it is consumed. If the oxygen concentration drops too low, hypoxic induced apoptosis occurs and the exposure of the cellular contents to neighbouring tissue may be fatal. On the other hand, if oxygen is supplied faster than it is consumed, the dissolved oxygen concentration rises and can impair cell viability and proliferation [208].

Table 1.12 – Control parameters critical for vascular bioreactor design (derived from Lyons and Pandit (2004) [207])

Biochemical Biomechanical

pO2 and pCO2 Shear Stress pH Volume 37°C Flow Rate 100% Humidity Compliance Waste Products Pressure Nutrition Resistance

The balance between acids and bases within the system is as critical as it is in physiology for the purpose of homeostasis. This is usually maintained with a pH between 7.35 and 7.45. A CO2-

bicarbonate buffer system manages the acidic-basic homeostasis in all living systems where CO2 reacts with H2O to produce carbonic acid which transforms rapidly to bicarbonate and hydrogen ions. An imbalance within this reaction causes pH shifts but is regulated by the respiratory system. In a closed bioreactor system, the culture medium typically contains this bicarbonate buffer system through the use of sodium bicarbonate. As cells metabolise CO2 is released into the media, causing an acidic shift.

The bicarbonate ions dissociated from sodium bicarbonate readily react with the dissolved CO2. In order to ensure a balance between the bicarbonate ions and the CO2, manufacturers of culture medium provide an indication of what percentage of sparged CO2 is needed.

In vasculature, endothelial cells sense and respond to shear stress and adapt accordingly. Endothelial cells induce neointimal thickening when decreased flow rates and thus decreased shear stress is sensed. Arteries will adapt to ensure wall shear stresses of 15 dynes/cm2. If the shear stress is too high or if turbulence is observed, cells can be damaged [209]. Physiological pressure of the cardiac output holds steady between the 80mmHg and 120mmHg on heart relaxation and contraction, respectively. Since the resistance and vascular network is too complicated to replicate in an artificial system for the purpose of pressure regulation, a simplified clamping method is used to adjust the pressure. The level of clamping of the fluid line correlates to the pressure within the system. This also depends on tubing thickness and diameter. The compliance chamber is a critical component of perfusion bioreactor systems required to dampen the noise generated from the peristaltic pump. A compliance chamber is a chamber filled with air and culture medium and adds elasticity to the closed bioreactor system [210].

The iconic vascular bioreactor was developed by Niklason et al. in 1999. Their system consisted of a medium reservoir connected to 4 vessel chambers in parallel sat on a magnetic stir plate with fluid flow (using a peristaltic pump) passing through a compliance chamber prior to the vessels as seen in Figure 1.32. The system aimed at synthesising functional arteries grown using smooth muscle cells and endothelial cells seeded onto polyglycolic acid scaffolds [211]. The system does not incorporate pressure profile measurements for the artery development but displayed their results relative to wall stress. The lack of physiological mechanical cues might slow down the process of the development of the arterial graft.

Figure 1.32 – Schematic of the Niklason vascular bioreactor. Adopted from Niklason et al. (1999) [211]

Mundargi et al. (2015) [212] developed a sensor based perfusion vascular bioreactor that allows for real-time monitoring of tissue viability using resorufin-based fluorescence. The system was connected to a customised plate within a fluorimeter and observed the effect of various oxygen levels on the viability of the tissue in real time. They use DMEM with 10% FBS and 2% P/S as their culture medium throughout the experiment. Their studies demonstrate observed denudation of the endothelium after 72h of dynamic culture and could potentially be attributed to the shear stress applied on the lumen as well as the decreasing oxygen concentration. This study demonstrates the potential to assess pharmaceutical compounds and stents within a system that allows for real-time monitoring of tissue viability. There are many issues that need to be overcome before this can be utilised in a practical aspect. The ability to maintain viability is critical and their study states the consistent decrease of viability to 30% after 72h, which does not exhibit therapeutic testing conditions required.

The constant development and research into stent technology has not demonstrated a correlative representation of studies on bioreactor systems aimed at assessing the efficacy of these stents/coatings. Wang et al. (2016) [213] assessed the degradation of magnesium wire (as a stent model) within excised porcine aortas. The study used DMEM with 10% FBS and 1% penicillin/streptomycin as their standard culture medium with a flow rate of 100ml/min under standard incubator conditions for 5 days. The authors suggest that the media used is physiologically closer than other commonly used media. The bioreactor used was a commercially available system (Lume Gen) consisting of the vessel chamber, pumps, flow control channel and a hertz oscillator and the entire system was autoclavable. Their study focused primarily on the comparison between the degradation in vivo compared to that of the ex vivo bioreactor model and less regarding the viability and maintenance of the tissue within the aforementioned system.

Yazdani and Berry (2009) [214] developed a vascular bioreactor for the purpose of stent testing with the goal of evaluating smooth muscle cell proliferation. They used carotid arteries perfused with DMEM with 10% FBS and 1% antibiotics/antimitotics. Their system allowed for catheterisation so that

the fixed artery can be stented. They performed contractile measurements induced by potassium chloride as a means of assessing functionality of arterial ring segments. Upon 7 days of conditioning within their bioreactor, they observed that the cultured tissue contracted 45% of that of the native fresh artery. Although their study does demonstrate the feasibility to assess proliferation using a BrdU assay, but makes no mention at the potential for the endothelial cells to heal. A key aspect would also be to check the viability of the tissue after the 7 days of conditioning. This study show that unstented areas do not show a proliferative capacity when compared to the stented regions and that the stents in question had different mechanical properties and thus would require different radial forces exerted onto the artery, potentially encouraging SMC proliferation. The bioreactor system utilised a custom program to generate the flow and pressure waveform through the artery. Although they achieved a regular pressure profile within the range of 80mmHg and 120mmHg, heavy noise was observed within their pressure readings. This could be attributed to the lack of a compliance chamber.

A study by Punchard et al. (2009) [215] developed a tubular model to assess the cellular and molecular changes to endothelial cells (HUVEC) subject to stent deployment. They coated silicone tubes with fibronectin and seeded at a high cell density of 171,500 cells/cm2. A commercially available stent, the Liberté, was deployed and subject to dynamic conditioning within a bioreactor for 24h. Their bioreactor system exhibited key physiological parameters onto the stented cell seeded structure: Radial distention (5%), pressure (40/120mmHg) and a flow with 10 dynes/cm2.The system used several aforementioned and commonly used techniques to provide these mechanical forces. The cells were dynamically conditioned in endothelial growth media (EGM-2). Although the study does identify changes in gene expression changes amongst the endothelial cells after the 24h, there was no indication as to the rate of cell migration onto the struts of the stent. The concept allows for a simplified system to assess the effect of stented arteries under certain near-physiological haemodynamic conditions.

The mechanical cues exhibited by the fluid flow within the bioreactor are critical to the behaviour of endothelial and smooth muscle cells. The endothelium faces direct fluid interactions and the cardiac cycle and blood flow results in cyclic strain, shear stress and circumferential deformations. Endothelial cells switch morphology from a cobblestone pattern to elongated structures that correspond in the direction of flow [216]. An increasing shear stress has been shown to decrease endothelial cell proliferation and inducing cell apoptosis and highlights the importance of maintaining a balanced physiological shear stress within a bioreactor system. Additionally, a lack of haemodynamic forces have also demonstrated induction of vascular endothelial cell apoptosis [217].

The medial layer consisting of smooth muscle cells are also directly exposed to the haemodynamic environment. The underexposure to these physiological forces can cause the smooth muscle cells to

dedifferentiate from their quiescent contractile state to the proliferative synthetic state [218]. It is critical that fluid flow characteristics are considered when developing bioreactor systems as cell phenotype and functionality is directly influenced by these parameters.

1.4.2 Parallel Plate Flow Due to the importance of laminar shear stress on the endothelium, parallel plate flow systems have been developed to assess various parameters under these dynamic settings. The system is designed so shear stress can be modulated by adjusting the fluid flow rate or the gap width of the flow path [219]. The shear stress in such a system can be calculated using the equation:

6푄 휏 = 휇 × in which τw is the wall shear stress applied to the cell monolayer푤ℎ , µ the viscosity (0.007 g/cm/s at 37° for culture media), Q is the flow rate, w is the width of the flow path and h being the gap width. The equation assumes a Newtonian fluid. Parallel plate flow chambers are usually designed in a bespoke manner based on the format of the seeded cells. The principle relies on the inlet flow passing fluid through a thin flow domain created by a silicone gasket. Figure 1.33 displays two schematics of such systems used for assessing endothelial cell monolayers under shear stresses. The seeded cell layer faces the fluid flow region and is packed tightly using appropriate O-rings to avoid for leakages. Some designs also include a vacuum port [220,221].

Figure 1.33 – Schematics of parallel plate flow systems used in current research. Adopted from Murayama et al. (1986) and Chung et al. (2003) [220,221]

The complexity of the cell response to haemodynamic forces is very difficult to recreate and thus by creating a system that provides accurate shear rates on the cells, allows for improved understanding of cell adhesion, arterial wall remodelling, mechanisms of mechanotransduction and associative gene products under physiological shear stresses [221].

1.5 Limitations of the Current State The anti-proliferative drugs available commercially have a negative effect on the endothelial cells within the stented artery where re-intervention and anti-platelet therapies are required to ensure there is no occlusion of the lumen due to thrombosis. Although, there are many promising strategies, from dual layered stents to endothelial progenitor cell capture in order to maximise endothelialisation whilst minimising smooth muscle cell proliferation, there are many pitfalls with current strategies. In order to identify an ideal therapeutic candidate, an effective methodical assessment workflow is required.

There are several studies utilising bioreactor systems to generate tissue engineered vascular grafts providing crucial physiological haemodynamic conditions in order to promote functional development of the tissue. There have been a few studies aiming to assess the effect of stents on arteries/arterial models but miss a critical aspect of stent development. Due to the nature of stent-related pathologies, primarily, the aspect of current drug eluting stents inhibiting the re-endothelialisation of the damaged lumen, it is imperative to aim to accelerate the re-endothelialisation process. There is a lack of research focused on assessing the efficacy of stent therapeutics and coatings with regards to endothelial cell adhesion, migration and proliferation onto the stent struts. Rapid endothelialisation would minimise neointimal hyperplasia due to the inflammatory response as well as ensure minimal platelet adhesion resulting in the eradication of a thrombotic response. Several studies have also monitored the response of novel stent designs and therapeutics in an in vivo setting. The importance of the three R principal is vital for translational research. Developing a near physiological in vitro system that could provide an indication whether a certain drug or coating is more effective than another will not only reduce animals for surgery but will also cut down on cost and time. A systematic platform for the assessment of various stent coatings is an important tool to identify optimal strategies and is crucial to the future development of stents.

Chapter II: Drug Assessment

2.1 Introduction Ever since the development of the drug eluting stent, there has been an exponential rise in research targeting potential drugs and therapies to combat the pathological drawbacks observed amongst the current commercially available stents. Industries have utilised their resources and the advances in laboratory robotics for this purpose and have performed high-throughput screening (HTS) of drugs and have created complex computational screening models to identify potentially effective drugs [222]. On the basis of the present study, literature screening was performed to identify candidate agents for inhibiting or maintaining SMC proliferation whilst enhancing EC proliferation. The knowledge of drug concentration and dosage time in an in vitro model setting is often used as a gauge for scaling up to a suitable in vivo drug concentration without surpassing the cytotoxicity limit of the drug. Owing to this, the candidate agents were assessed in terms of their effect on the proliferative states, cell phenotypes, migratory patterns and function of SMCs and ECs. The cells were assessed in order to identify whether the compounds exhibited a dual-regulatory action on inhibiting SMC and promoting EC proliferation.

2.2 Materials and Methods This section describes the materials, equipment and methodologies used for the assessment of the cellular response to the different compounds in terms of cytotoxicity, proliferation, migration and gene expression.

2.2.1 Cell Culture All cell culture and incubation stages were carried in a temperature-and gas-regulated incubator

(Heracell VIOS 160i Incubator, Thermo Scientific). CO2 was set to 5% (v/v) and the temperature was maintained at 37°C. The culture methods described in this chapter were used throughout this project unless otherwise stated in subsequent chapters.

2.2.1.1 Cells and Cell Culture Consumables The cells studied in the work present in this chapter and subsequent chapters are highlighted in

Table 2.1. The solutions used in the cell culture protocol are listed in Error! Reference source not found..

Table 2.1 - Cells used in the study

Cell Species and Tissue Source Source Abbreviation Endothelial Cells Porcine Coronary Artery Cell Application – P300-05 PCAECs Endothelial Cells Human Umbilical Vein Primary - Previous isolation HUVECs Smooth Muscle Cells Porcine Coronary Artery Cell Application – P350-05 PCASMCs Smooth Muscle Cells Human Coronary Artery Cell Application – 350-05a HCASMCs

Table 2.2 - Solutions and media used for cell culture

Solution Source Abbreviation Porcine Smooth Muscle Cell Cell Application (P310-500) SMBM Basal Medium Porcine Smooth Muscle Cell Cell Application (P311-GS) SMGM (supplemented into Growth Supplement SMBM) Porcine Smooth Muscle Cell Cell Application (P311I-250) SMDM Induction Medium Endothelial Basal Media Lonza (CC-3156) EBM-2 Endothelial Growth Medium Lonza (CC-3162) EGM-2 BulletKit Dulbeco’s Modified Eagle Gibco (21969035) DMEM Medium Phosphate-Buffered Saline Gibco (70011) PBS Penicillin/Streptomycin Lonza (17-602) P/S Trypsin/EDTA Biochrom (L2153) T/E Dimethyl Sulfoxide Applichem (A3672) DMSO Foetal Bovine Serum Biochrom (S 0615) FBS CASYton Solution OLS (5651808) N/A

2.2.1.2 Cell Seeding of Cryopreserved Cells Frozen cryovials were removed from liquid nitrogen storage and dipped into a 37°C water bath (1083, GFL, Burgwedel, Germany) until a small ice crystal was visible within the cryovials. The cryovials were sprayed with 70% ethanol and transferred into to a class II laminar flow cabinet. The contents of the vials were pipetted into a sterile falcon tube containing 25ml of warm growth media (SMGM or EGM- 2 containing 1% P/S) and transferred to 175cm2 culture flasks. The cells were incubated at 37°C and

5% (v/v) CO2 for 24h before the flasks were washed three times with PBS and the culture media was replenished.

2.2.1.3 Cell Passaging Once the cells had reached 80% confluency, as evidenced under light microscopy, the culture media was aspirated and the culture flasks were washed gently with warm PBS three times to ensure maximum removal of dead cells. The PBS was aspirated and 6ml of warmed trypsin/EDTA was added to the flasks. The flasks were tilted to ensure complete coverage of the culture surface before incubation at 37°C and 5% (v/v) CO2. The flasks were gently tapped to ensure the detachment of cells from the culture surface. The cells were monitored under phase contrast microscopy to ensure

adequate cell detachment prior to the addition of 18ml of complete media (DMEM with 10% (v/v) FBS and 1% (v/v) P/S) to deactivate the trypsin. The solutions were pipetted up and down a few times to reduce any cell agglomerates. The cell suspension was centrifuged at 300 g for 5 min. The supernatant was removed and the cell pellet was re-suspended in 1ml PBS. Cells were counted using a CASY Cell Counter (OLS, Bremen, Germany). 20µl of the cell suspension was added to 10ml of CASY solution. The counting solution was placed in the automatic cell counter and the value of viable cells per ml was recorded. Cells were also plated into 175cm2 flasks with pre-warmed media and used in subsequent experiments or stored for later use. All cell incubations were performed at 37°C and 5% (v/v) CO2 in humidified air.

2.2.1.4 Cell Cryopreservation The cryopreservation medium consisted of 90% (v/v) foetal bovine serum and 10% (v/v) dimethyl sulfoxide (DMSO). This solution was warmed prior to the addition of cells. After counting the cells using the CASY Cell Counter, appropriate volumes of cell suspension to account for 1.5 million cells were pipetted into a 15ml falcon tube. The falcon tubes were centrifuged for 5 min at 300 g and the supernatant was discarded. 1.5 ml of pre-warmed cryopreservation solution was added to the cell pellet and pipetted to suspend the cells. The cell suspension was added to labelled cryovials and placed in a freezing container (Mr. Frosty Freezer, Thermo Scientific), consisting of 100% isopropyl alcohol. The freezing container was placed in a -80°C freezer for 24 hours to allow for controlled cooling at - 1°C/min, before transferring the cryovials into liquid nitrogen for long term storage.

2.2.2 Drug Preparation The drugs assessed in this work were supplied in solid form and are listed in Table 2.3. The compounds were dissolved in DMSO under sterile conditions, according to the manufacturers’ guidelines. The compound solutions were then agitated sufficiently to allow for complete solubilisation of the solute. Table 2.3 - Drugs used in the study

Drug Company Molecular Dissolved in Product No. Weight Exendin-4 Enzo Life Sciences 4186.57 DMSO ENZ-PRT111 Ferulic Acid Sigma-Aldrich 194.18 DMSO 128708 Magnolol LKT Labs 266.34 DMSO M0125 Curcumin Tocris 368.38 DMSO 2841 Paclitaxel AdipoGen 853.9 DMSO AG-CN2-0045 Everolimus Cayman 958.2 DMSO 11597 2.2.3 Metabolic Activity Assay A cell metabolic activity assessment assay was adopted in order to assess the potential cytotoxic effects and proliferation-induction capacity of the selected compounds. The Cell Counting Kit – 8 (CCK- 8; PromoCell) was used to assess the metabolic activity of the cells following drug treatment. The assay

utilised a water-soluble tetrazolium salt (WST-8), which resulted in its reduction to a soluble formazan dye through mitochondrial dehydrogenases (Figure 2.1). The dye intensity was directly proportional to the number of living cells.

Figure 2.1 - The mechanism of the WST-8 assay. Adopted from Held (2009) [223]

Cells were seeded in 96-well plates at 5 104 cells/well. Each well plate was associated with a dosage time (24h, 48h, 72h or 144h). Following× seeding, all plates were maintained at 37°C and 5% (v/v) CO2 in the respective growth medium for 4h to allow for cell attachment, prior to adding the required drug concentrations for the respective drug compound used, and incubated for 24h, 48h, 72h or 144h. After incubation, the drug-medium solution was aspirated and the wells were washed with DPBS (14040- 091; Gibco) three times to ensure minimal compound residual that might affected absorbance. 10% WST-8 solution in Hanks Buffered Saline Solution (HBSS) (14025-092; Gibco) was added to each well. The plates were then incubated for 4h and the WST-8 solutions were transferred to a fresh 96-well plate for assessment (with technical triplicates). The plates were loaded into a microplate reader (Synergy 2, BioTek, Winoski, USA) and the absorbance was measured at 450nm and 630nm (background). In the calculated results, the background values obtained at 630nm were subtracted from the values obtained at 450nm.

2.2.4 Would Healing Assay PCASMCs and PCAECs were seeded at 8,000 cells/cm2 in 24-well plates and allowed to reach confluency in their respective growth medium. Subsequently, the culture medium was aspirated and a vertical scratch was performed in the middle of the well using a 200µl pipette tip. The wells were then washed with PBS three times before adding a drug-medium combination in accordance to Table 2.4. The plates were then placed within an incubation station attached to a microscope (Axio Observer Z1, Zeiss) for time lapse imaging (Figure 2.2). Sterile water was added to the space between the wells to allow for humid conditions. Each well was assessed individually to ensure appropriate focus was

achieved at a well-defined region of the scratch. Once the appropriate focus was achieved, the axes were logged for each well and the program was run for 48 hours with images taken every 7 minutes.

Table 2.4 - Concentrations of drugs used for the Wound Healing Assay for PCASMC and PCAEC

Drug Concentrations Untreated (Positive Control) N/A Everolimus (Negative Control) 100µM Ferulic Acid 50, 500 µM Exendin-4 10, 100 nM Magnolol 93.9 µM Curcumin 10 µM

Figure 2.2 - Setup of the Axiovision Z1 Microscope with the incubation chamber used for the Scratch Assay

After the image sequences were recorded, 1:1000 Calcein AM (L3224, Invitrogen) was added to each well and incubated for 30min. Single images were taken of the calcein-stained wells under blue excitation (470nm). Calcein acetoxymethyl is a non-fluorescent derivative of calcein that is capable of permeating the membrane of viable cells. Esterases within the cell cleave the acetoxymethyl group and the calcein molecule gives off a green fluorescence. Since esterases only exist within living cells, this makes Calcein AM an ideal tool to visually label viable cells [224]. Cell migration was determined using ImageJ, where the measured area of the wound over several fixed time point images was taken as a percentage of the initial t=0 scratch area using the following equation:

퐴푟푒푎 − 퐴푟푒푎 퐺푎푝 퐶푙표푠푢푟푒 (%) = × 100 퐴푟푒푎 where Area0 was the wound area at t=0 and Areat the wound area at subsequent migration times.

2.2.5 Gene Expression 2.2.4.1 Endothelial Cell – Gene Expression Assessment PCAECs were cultured to P5 in EGM-2 and seeded in 6-well plates at 8,000 cells/cm2 in triplicates. After reaching 80% confluency, the desired concentration of each drug was added to each well triplicate in accordance to Table 2.4. One triplicate group was not treated with any drugs and served as control group. The PCAEC cultures were then cultured in an incubator under standard conditions for 144 hours. The culture solution was replenished after 72h. 144 hour was used to determine the long term sustainability of the endothelial phenotype upon treatment with the drug.

Table 2.4 - Drug concentrations for qPCR of PCAECs

Sample Concentration Untreated N/A Magnolol 93.9µM Curcumin 10µM Everolimus 10µM Ferulic Acid 50µM, 500µM Exendin-4 10nM, 100nM

2.2.4.2 Smooth Muscle Cell – Gene Expression Assessment PCASMCs were cultured to P5 in SMGM and seeded in 6-well plates at 8,000 cells/cm2 in triplicates. After reaching 80% confluency, the growth medium was aspirated and differentiation medium (SMDM) was added to the cells. The cells were kept in SMDM for an additional 4 days, with media change occurring every second day. Subsequently, fresh SMDM with the desired concentration of either ferulic acid or exendin-4 (Table 2.5) was added to each well triplicate. Triplicates were maintained without any drug supplement and served as untreated controls. The triplicates were assessed after either 24, 48, or 72h. Multiple time points were assessed in order to determine potential phenotypic changes over the course of the drug treatment.

Table 2.5 - Drug concentrations for PCASMC Time-Dependant qPCR

Sample Concentration Untreated N/A Ferulic Acid 50µM, 500µM Exendin-4 10nM, 100nM

2.2.4.3 RNA Isolation The NucleoSpin® RNA Mini kit (Macherey-Nagel, Duren, Germany) was used for RNA isolation from the PCAECs and PCASMCs, treated with the drug compounds, in accordance to the manufacturer’s guidlines. The culture medium was discarded and a lysis buffer containing 1:100 β-mercaptoethanol (31350, Gibco) was added to each well to detach the cells. 70% ethanol was added to the cells and was resuspended and transferred to a spin filter column. The spin columns were centrifuged at 11,000g for

30s. The filtrate was discarded and the remaining lysed solution was centrifuged. Membrane desalting buffer was then added to the column and subsequently centrifuged at 11,000g for 1 minute. The filtrate was discarded and a DNAse-reaction solution was added to the column membranes and incubated at 25°C for 15 min. RA2 buffer was added and centrifuged for 30 seconds, followed by RA3 for 30 seconds and again for 2 minutes. The filtrate was discarded after each centrifugation. 50µl of RNase-free water was added directly to the filter and centrifuged for 1 min to elute the RNA. The RNA concentration was determined using a nanodrop spectrophotometer (ND 1000, Thermo Scientific)

2.2.4.4 cDNA synthesis 500ng of the mRNA, calculated according to the concentration obtained from the Nanodrop reading, were used for cDNA synthesis, and RNase free-water was added to make up 11µl. 1µl of random hexamer primer was added to the tubes and centrifuged shortly to mix the solution. The samples were incubated at 65°C for 5 min followed by the addition of 8µl of the reaction mixture, which consisted of 5x reaction buffer, RiboLock RNase Inhibitor, 10mM dNTP mix and reverse transcriptase. The tubes were incubated at 25°C, 42°C and 70°C for 5, 60 and 5 min, respectively. The incubation stages were conducted in a thermocycler (peqSTAR X Cycler, Peqlab, Erlangen, Germany). The cDNA was diluted down to 1:5 with RNase-free water and stored at -20°C until required for qPCR.

2.2.4.5 Quantitative-Polymerase Chain Reaction (qPCR) All primers used for qPCR were designed using the NCBI database. Gene and species specificity of the amplicon sequences were tested using the BLAST program. All oligonucleotides were obtained from Eurofins Genomics (Ebersberg, Germany) and the working primer solutions were diluted to 10pmol/µl. Stock solutions were stored at -20°C. An assortment of primers was designed, following primer linear amplification validation. The final primers used for gene expressions studies are detailed in Table 2.6.

Table 2.6- DNA Oligonucleotide primers sequences for RT-PCR

Protein Primer Name Sequence 5’ to 3’ Tm (°C) Product Size pACTB F1 GATCAAGATCATCGCGCCTCC 61.8 21 β-Actin [Housekeeping Gene] 1,2 pACTB R1 GGAATGCAACTAACAGTCCGCC 62.1 22 pNOS3 F2 CGGCGCTATGAGGAGTGGAA 61.4 20 Endothelial Nitric Oxide Synthase 1 pNOS3 R2 ATCTCTCCCGGGTAGGTGCT 61.4 20 pCDH5 F2 GCGAGTTCACCTTGTGCGAG 61.4 20 Vascular Endothelial Cadherin 1 pCDH5 R2 CGAGGAGGGAGATCACTGCG 61.4 20 pCD31-ii F1 CACCGAGGTCTGGGAACAAAG 61.8 21 Platelet Endothelial Adhesion Molecule 1 pCD31-ii R1 TCTGCTCTGCGGTCCTAAGT 59.4 20 pCNN1 F1 CCAGCATGGCCAAGACGAAAG 61.8 21 Calponin–1 2 pCNN1 R1 CCCAGCTTGGGGTCGTAGAG 63.5 20

α-Actin 2 pACTA2 F1 CCCCAGAAGAGCATCCGACC 63.5 20

pACTA2 R1 GAGTCCAGCACAATGCCAGT 59.4 20 pSMTN F1 GGGGCTATGAGCATGTGGAC 61.4 20 Smoothelin 2 pSMTN R1 ACGCACTTCCAGTCAGGCTC 61.4 20 pMYH11 F1 AGCCCAGAAGAACGAGAGCG 61.4 20 Myosin Heavy Chain-11 2 pMYH11 R1 TCCAGTGCCGAAATGGTGGA 59.4 20 1 Primers used for PCAECs, 2 Primers used for PCASMCs

A SYBR master mix was created using Absolute QPCR SYBR green Mix (12.5µl), forward primer (1µl), reverse primer (1µl) and RNase-free water (5.5µl). 20µl of the SYBR master mix was added to each well in a RT-PCR 96 well plate, along with 5µl of the diluted cDNA sample. Water and Amplicon based controls were also included. All samples included a technical replicate. The well plates were then centrifuged and loaded into a PCR cycler (peqSTAR 96Q, Peqlab, Erlangen, Germany). The following profile was used for amplification:

Initial Activation - 95°C for 15 minutes

Denaturing Stage - 95°C for 15 seconds

Annealing Stage - 60°C for 30 seconds 40 cycles

Elongation Stage - 72°C for 30 seconds

Denaturing Stage - 95°C for 15 seconds

Annealing Stage - 60°C for 30 seconds

Melting Curve - 95°C for 5 seconds

Held at 4°C after melting stage

2.2.4.6 Data Analysis for Gene Expression Quantification Melting curve analysis was conducted to ensure amplified product specificity. The raw Ct values were analysed using the comparative ∆∆Ct method in Microsoft® Excel. The equation used to calculate the relative fold change using the ∆∆Ct method was:

(()()) 퐹표푙푑 퐶ℎ푎푛푔푒 = 2 where EG was the Ct value of the sample well with the specific gene of interest, AvgEH the mean Ct value of the sample wells with the housekeeping gene, AvgCG the mean Ct value of the control wells with the specific gene of interest and AvgCH the mean Ct value of the control wells with the housekeeping gene. The results were statistically analysed and plotted using GraphPad Prism 6.

2.2.6 Flow Cytometry PCAECs at P5 were seeded on 24 well plates in EGM-2 until approximately 80% confluency. The medium was aspirated and the wells were washed three times with PBS. The drug compounds were added to the wells in triplicates and were incubated for 24h. The medium with the drug was removed and the wells were washed three times with PBS. 400µl of Serum free media (EBM-2) was added to each well for a 24 hour serum starvation phase. 15µg/ml of pHrodo™ Green-labelled LDL (L34355, Invitrogen) were added to the wells and further incubated for 4h.

The medium with the LDL was then discarded and the wells were treated with 100µl of Accutase (A6964, Sigma) for cell detachment and were left in the laminar flow cabinet for 10min. The wells were checked using a light microscope to ensure efficient cell detachment. 300 µl of complete media was then added to the wells to ensure 3 times the volume of Accutase. The cell suspensions were transferred to Eppendorf tubes and centrifuged at 300g for 5min. The supernatant was discarded and the cell pellet was re-suspended in 150µl of FACS running buffer (MACSQuant, 130-092-747, Miltenyi, Bergisch Gladback, Germany). 15µl (1/10) of Anti-CD31 antibody [LCI-4], Allophycocyanin (APC) (GTX44005, GeneTex) was added to all the drug treated samples as well as the serum free control samples. The tubes were left at room temperature for 10min before transferring the contents to a round bottom 96-well plate. The well plate was placed on the Miltenyi chilling platform of the flow cytometer (MACSQuant Analyser 10, Miltenyi Biotec, Bergisch Gladback, Germany). 1µl of propidium iodide (130-093-233, Miltenyi, Bergisch Gladback, Germany) was added to each well to stain for dead cells. The recorded data was analysed using FlowJo (Treestar Inc, Oregon, USA).

2.2.7 Statistical Analysis The data was plotted as mean ± 95% confidence interval. All data were analysed using Graphpad Prism 6. Analysis of variance (ANOVA) testing followed by Bonferroni multiple comparison post hoc testing were performed. A statistical significance was recognised as p < 0.05. The metabolic activity assessment of the drugs were normalised and baseline corrected to the positive control (untreated) as a 100% and negative control (No Cell) as 0% values in order to achieve appropriate comparisons. This was achieved by using the following equation:

100 × (푂퐷 − 푂퐷) 푀푒푡푎푏표푙푖푐 퐴푐푡푖푣푖푡푦 (%) = (푂퐷 − 푂퐷) Where ODSample, Positive and Negative correspond to the optical density of the test sample, the mean of the positive controls and the mean of the negative controls, respectively. OD OD

2.3 Results 2.3.1 Metabolic Assessment of Drugs on Endothelial Cells 2.3.1.1 Human Umbilical Vein Endothelial Cells The metabolic activity of the human umbilical vein endothelial cells (HUVECs), exposed to the different concentrations of paclitaxel and everolimus for 24h, 48h, 72h and 144h is presented in Figure 2.3. After 24h exposure to all three concentrations of paclitaxel, the metabolic activity of the HUVECs dropped to around 14%, 29% and -8% of the untreated group for 1, 10 and 100 µM, respectively. After 48h, these values demonstrated no significant change (P>0.99, P>0.99 and P=0.36 for 1, 10 and 100µM, respectively). For the 1µM (P=0.015) and 10µM (P<0.0001) samples, a significant difference was observed between the 48h and the 72h exposure times. For the 72h and 144 culture plates, there was no significant difference between any of the concentrations and the No Cell negative control, as well as between the groups (P>0.99). In the case of everolimus, a 24h exposure demonstrated a significant loss of metabolic activity amongst all three concentrations (P<0.0001), where the metabolic activity dropped to 28, 42 and -25% for 1, 10 and 100µM respectively. The values for 1 and 10 µM were almost double that of paclitaxel at the same exposure time. After 48, the metabolic activity dropped further to 18% and 29% for the 1 and 10 µM concentrations, but with no significant statistical difference (P=0.40 and 0.09, respectively). After 72h of exposure, there was no significant differences between the treated samples and the No Cell negative control (P>0.99, P>0.99 and P=0.18 for 1, 10 and 100µM, respectively). For the 1 and 10µM sample, a significant reduction of metabolic activity between the 24h and the 72h group was observed dropping to 0.76% and 0.28% respectively (P<0.0001). The same phenomenon was observed for the samples exposed for 144h, with P=0.66, 024 and 0.53 for the 1, 10 and 100µM samples, respectively. Although there were significant differences between the samples treated with 100µM, these values were either negative or displayed no significant difference with the negative control. Both commercially used drugs exhibited a similar rate of metabolic decrease. Both sets of treated samples demonstrated a significant difference (P<0.0001) compared to the untreated group for the respective time points.

Figure 2.3 – Mean metabolic activity of HUVECs treated with different concentrations of Paclitaxel and Everolimus for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

The metabolic activity of the HUVECs exposed to the different concentrations of curcumin for 24h, 48h, 72h and 144h is presented in Figure 2.4. After 24h exposure to curcumin, the HUVECs demonstrated a reduced metabolic activity amongst the 5, 10 and 25µM concentrations, relative to the untreated control (P<0.001, 0.0001, 0.0001, respectively). The highest concentration (100µM) demonstrated no significant difference compared to the positive control (P>0.99). At all exposure times tested, there no significant differences between the 10 and 25 µM treated samples and the negative control (P>0.99). For each concentration, there was no statistical difference between the metabolic activities of any time point, except for the 100µM sample. This shows that for all the concentrations except for the 100 µM, there was no further changes after 24h. The metabolic state of the 100µM treated cells started at 95% at 24h of exposure and dropped to 58%. This change yielded a significant difference with P=0.01. The metabolic activity dropped down to 25% at 72h, with a significant difference with P=0.021. Between the final two exposure groups of the 100µm sample (72h vs 144h), there was no statistical difference (P>0.99). For all time points, the 1 µM sample demonstrated no statistical differences with the corresponding positive controls. The 100 µM sample displayed an anomalous metabolic activity that did not find a place with the observable reduction in metabolic activity upon concentration increase to 25µM.

The metabolic activity of the HUVECs exposed to the different concentrations of magnolol for 24h, 48h, 72h and 144h is presented in Figure 2.5. All magnolol concentrations above and including 37.5µM induced a significant reduction in the metabolic activity of the cells relative to the untreated control (P<0.0001), as well as matching that of the negative control (No-Cells group) with no significant difference (P>0.99) for all exposure times. The 3.75µM dosed sample showed no change between the 24 and 48h exposure times (P>0.99) but exhibited a drop to 76% upon 72h of magnolol exposure. A

statistical significance was observed within the change in metabolic activity between 48h and 72 (P<0.05) but no further change between 72h and 144h (P>0.99). On the other hand, ferulic acid (Figure 2.6) did not induce any significant effects on the HUVECS after 24h, at any of the concentrations tested. Upon 48h of FA exposure, the 5 and 50 µM sample exhibited an increase of metabolic activity to 133% and 136%, respectively, with a significant difference compared to the untreated positive control (P=0.02 and P<0.01, respectively). After 72h, the HUVECs treated with 50µM ferulic acid demonstrated a significantly increased metabolic activity to 147% (P<0.001). After 144 hours, the HUVECs treated with 5, 50 and 100µM ferulic acid exhibited a significant rise in their metabolic activity to 156, 145 and 134%, respectively, with P<0.0001, P<0.001 and P=0.014, compared to the positive control. Between the exposure groups, the only observable statistical significant change was within the 5µM treated cells, between 24h and 72h (P<0.01) and 72h and 144h (P<0.01). None of the treated samples shared any statistical significance with the negative control.

The metabolic activity of the HUVECs exposed to the different concentrations of Exendin-4 for 24h, 48h, 72h and 144h is presented in Figure 2.7. The cells seemed to be unresponsive to Exendin-4 for the first 48h, except for an anomaly observed for the 1nM concentration that reduced the metabolic activity of the cells to 37% (P<0.001 compared to the positive control). Additionally, the 50nM concentration also caused a significant reduction in the metabolic activity of the HUVECs after 48h, where a significant difference was observed between the 50nM metabolic activity and the untreated control as well as the negative control (P<0.01 for both). The reduced-metabolic-activity trend with the 50nM concentration continued through to 72h and 144h and showed significant differences compared to the untreated (P<0.001 and P<0.01 for 72h and 144h, respectively). At 72 hours, the 25nM concentration induced a significant increase in the metabolic activity of the cells to 143% (P=0.026, compared to untreated). After 144h exposure to the 5, 10 and 25nM dosages the cells demonstrated a significant boost in their metabolic activity to 153%, 188% and 212%, respectively (P<0.01, 0.0001 and 0.0001, respectively – compared to untreated). The 5nM sample demonstrated a significant increase between the metabolic rate at 72h and at 144h (P<0.01). The 10nM sample exhibited significant differences between 24h and 144h (P<0.001), 48h and 144h (P<0.001) and 72h and 144h (P<0.01). The same significant differences were observed amongst the cells treated with 25nM (P<0.0001 for 144h vs 24h and 48h). For the metabolic decreasing 50nM concentration, a difference was observed between 24h and 72h (P<0.01) as well as 144h (P=0.019).

Figure 2.4 - Mean metabolic activity of HUVECs treated with different concentrations of curcumin for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.5 - Mean metabolic activity of HUVECs treated with different concentrations of magnolol for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.6 - Mean metabolic activity of HUVECs treated with different concentrations of ferulic acid for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.7 - Mean metabolic activity of HUVECs treated with different concentrations of Exendin-4 for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

2.3.1.2 Porcine Coronary Artery Endothelial Cells Similar to the case of the HUVECs, curcumin induced a similar effect to the metabolic activity of porcine coronary artery endothelial cells (PCAECs) (Figure 2.8). The 1µM treated sample demonstrated a significant increase after 18h (P<0.01) and 144h (P=0.03) of exposure with a metabolic activity of 159% and 151%, respectively. The other time point wells exhibited no significant difference between the untreated control and the 1µM treated samples. The 5 µM sample demonstrated a significant decrease in metabolic activity amongst the 18h (P<0.0001), 36h (P<0.01), 48h (P<0.01) and 72h (P<0.0001) exposure times. The 10 and 25 µM treated samples demonstrated a significant reduction of metabolic activity (P<0.0001 for all time points). These values were not statistically different from the negative control. A similar observation amongst the 100µM sample, compared to that of the curcumin-treated HUVECs. The high dose sample displayed no statistical differences compared to the untreated control at the 18h point (P>0.99). A linear negative trend was observed as exposure time was increased. For each exposure time relative to the subsequent time point, the 1µM treated sample demonstrated no significant difference. For the 5µM treated PCAECs, a significant increase in metabolic activity was observed between the 18h and the 24h time points (P<0.01). The succeeding time points were demonstrated no significant differences apart from the 72h to 144h (P<0.01) comparison. The 10 and 25 µM samples demonstrated a fixed metabolic activity throughout the duration of the experiment where no differences were observed between the time points for these specific concentrations. The 100µM treated sample, however, displayed a slow decay of metabolic activity. After 18h, the PCAECs exhibited a metabolic activity level of 106%, this dropped down to 57% at the 24h point (P=0.010 for 18h vs 24h). There were no further comparable differences between the subsequent time points and the preceding time points. However, the level of significant differences of the 100µM and the untreated control kept increasing (except for the 72h time point).

Exposure to magnolol produced similar metabolic activities in the PCAECs (Figure 2.9), as in the case of the HUVECs. The only exception was the significant increase in metabolic activity of the cells treated with 3.75µM at the 36h time point (P<0.01). The cells treated with 37.5µM and higher shared no significant differences relative to the No Cells negative control.

The effect of ferulic acid on the PCAECs is presented in Figure 2.10. Relative to the No Cells control, all dosed samples at all time points demonstrated a significant differences (P<0.0001). At the 18h point, the only statistically significant increase observed was amongst the low 0.5µM concentration compared to the untreated control. Ferulic acid induced an increase in cellular metabolic activity when used at the concentrations of 5, 50 and 100 µM after 24 hours displaying a boost to 161, 196 and 190%, respectively (P<0.05, 0.0001 and 0.001, respectively). After 144h, 0.5 (P<0.01), 5 (P<0.05) and 50 (P<0.05) µM of ferulic acid induced a significant increase in the metabolic activity of the PCAECs

compared to the untreated control, with values of 182, 165 and 166%, respectively. Between the time groups for the specific concentrations, only 0.5µM treated cells displayed a significant difference, between 18h and 24, 36, 48 and 72h (P<0.01), which could be attributed to the anomalous metabolic activity observed at 18h.

For the Exendin-4 treated samples (Figure 2.11), relative to the untreated control group for the respective exposure times, the significantly different samples were observed amongst the 1nM (P=0.024) and 5nM (P=0.014) samples at 18h, 25nM (P<0.01) at 24h and all the treated samples at 144h. At 144h, the biggest metabolic increase over the control was observed in the 50nM treated cells (P<0.0001) with a metabolic activity of 234%. Values of 175%, 200%, 175% and 194% were observed amongst the PCAECs treated with exendin-4 with concentrations of 1nM (P=0.036), 5nM (P<0.01), 10nM (P=0.034) and 25nm (P<0.01), respectively. Although significant differences were observed between some time points for the individual concentrations tested, a clear time based trend was not observed. For the 5nM sample, there was a significant increase in metabolic activity between the 72h and 144h exposure times (P=0.039). This observation was not observed with the 10nM samples. This 72h vs 144h difference was repeated with the 25nM (P=0.013) and 50nM (P<0.001) treated samples.

Figure 2.8 - Mean metabolic activity of PCAECs treated with different concentrations of curcumin for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.9 - Mean metabolic activity of PCAECs treated with different concentrations of magnolol for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.10 - Mean metabolic activity of PCAECs treated with different concentrations of ferulic acid for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.11 - Mean metabolic activity of PCAECs treated with different concentrations of Exendin-4 for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

2.3.2 Metabolic Assessment of Drugs on Smooth Muscle Cells 2.3.2.1 Human Coronary Artery Smooth Muscle Cells The commercially used DES drugs Paclitaxel and Everolimus induced a slightly different effect on the human coronary artery smooth muscle cells (HCASMCs), compared to the HUVECs (Figure 2.12). The metabolic activity of the cells treated with 1 and 10 µM for both drugs remained high during the duration of the study but still demonstrated a significant difference compared to the untreated control, with the exception of 1µM treated cells at 48h (P>0.99). The 100µM dosed cells dropped their metabolic state to that of the No Cells control group after the first 24 hours (P=0.31 compared to the No Cells control). The metabolic state remained as such for all the assessed time points. The everolimus-treated HCASMCs demonstrated similar effects amongst the low concentrations (1µM and 10µM) on HCASMCs compared to that of paclitaxel. After 144h, the 1µM treated cells were comparable to the untreated control group. The cells treated with 10µM Everolimus demonstrated an initial reduction in their metabolic activity, to almost the minimal value (negative control), but for longer exposure times they exhibited an increase in metabolic activity (from 12% at 24h to 79% at 72h). Between the time points, the only key differences noted were associated with the 1µM treated samples between 72h and 144h (P<0.01) as well as the increase noted before with the 10µM treated cells between 24h and all subsequent time points (P<0.0001).

Figure 2.12 - Mean metabolic activity of HCASMCs treated with different concentrations of Paclitaxel and Everolimus for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Curcumin’s effect on the metabolic activity of HCASMCs is shown in Figure 2.13. Unlike the observation with the ECs, curcumin did not demonstrate significant effects on the HCASMCs amongst the concentrations between 1 and 10µM. An exception to this was observed at 72h amongst the 10µM treated samples, where a significant difference between the tested sample and the untreated control was observed (P=0.034). For the 25µM treated cells, a significant drop to 13% from 89% was observed at 48h of exposure. This effect was retained throughout the duration of the study. The 100µM treated cells demonstrated immediate minimisation of metabolic state after the initial 24h time point. The 25µM (from 48h onwards) and 100µM displayed no significant difference between the No Cell negative control and the tested samples amongst all time points. Analysis between the time points demonstrated a significant difference between the 24h and 72h exposure amongst the 10µM treated cells (P<0.01). Another clear time based difference between the 25µM treated cells at 24h and 48h (P<0.001), 72h (P<0.0001) and 144h (P<0.0001) was observed.

The effects of magnolol on the metabolic activity of HCASMCs are displayed in Figure 2.14. After 24h of magnolol exposure, significant differences between the untreated control and the 93.9µM (P=0.02), 187µM (P<0.0001) and 375µM (P<0.0001) were observed, with a drop of metabolic activity to 50%, 0.5 and 0%, respectively. The 187 and 375µM treated HCASMCs retained this reduction in metabolic activity amongst all time points. At 48h, the metabolic activity of the 93.9µM sample dropped down further to 19%, where an insignificant difference between the negative control and the tested value was observed (P>0.99). This comparison between the 93.9µM and the No Cell negative control was observed for the 72h and 144h exposure times as well. The 3.75µM treated samples started showing a significant loss of metabolic activity from the 48h time point onwards, compared to the untreated control. For the 3.75µM treated cells, the metabolic activity showed significant differences between

the 24h and 48h (P<0.01), 72h (P=0.037) and 144h (0.016) exposure times. The 37.5µM sample demonstrated statistical time-based differences between 48h and 72h (P=0.012) followed by reversion back to the same metabolic activity at 144h (72h vs 144h with P=0.025). The only statistical difference that occurred amongst the 93.9µM treated samples were between 24h exposure and 144h exposure (P<0.001).

The metabolic activity of HCASMCs exposed to different concentrations of ferulic acid for different exposure times is illustrated Figure 2.15. After 24h of exposure, none of the concentrations demonstrated any significant difference compared to the positive control. This was also observed for the 48h exposure set. At 72h, the 50M treated sample demonstrated a great boost in metabolic activity to 240% (P<0.0001 compared to untreated). This boost was not significant at the 144h time point. However, the 100M treated cells did exhibit a significant increase (180%) in metabolic activity compared to the untreated control (P=0.035). Relative to the negative control, all treated samples demonstrated varied levels of significant differences and none of the samples shared similar values to that of the negative control. Between the time groups for the individual concentrations, 50M treated cells at 24h vs 72h (P<0.01), 48h vs 72h (P=0.043) and 72h vs 144h (P=0.030) displayed statistical differences. For the 100M treated HCASMCs, observable differences were seen between 24h and 48h (P=0.035) and 24h and 144h (P=0.011).

The metabolic activity of the HCASMCs under exposure to Exendin-4 is illustrated in Figure 2.16. At the two lowest concentrations, 1nm and 5nm, after 24h of exposure, the metabolic activity of the cells dropped significantly to 30% and 55%, respectively (P<0.0001 and P=0.022, respectively). The values for these concentrations appear to be inconsistent with the metabolic activities for the subsequent time points. The 5nM (P=0.032) and 10nM (P=0.015) treated cells demonstrated a significant increase over the untreated sample at the 48h point. At the 72h point, the cells treated with 10nM and 25nM of exendin-4 demonstrated an increase in metabolic activity to 186% and 141%, respectively (relative to control, P<0.0001 and P=0.049, respectively). The 144h exposed cells only yielded a significant increase over the untreated control amongst the 25nM treated cells (P=0.032). Throughout all the time points, the 50nM treated cells did not exhibit any significant changes in metabolic activity relative to the control group. For the anomalous 1nM treated cells at 24h, significant differences were observed compared to the 48h, 72h and 144h exposure times (P<0.0001). This occurrence was identical amongst the 5nM and the 10nM treated cells (P<0.0001 for 24h compared to all other time points for both concentrations). A significant change was also observed between the 48h and 72h exposure time for the 10nM treated cells (P<0.01), as well as for the 72h vs the 144h time points (P<0.01). The 25nM treated samples demonstrated no key changes between the different time points. Although the 50nM

treated cells displayed no significant difference amongst the 24h set, significant differences between the metabolic activity at 24h and 48h (P=0.036), 72h (P<0.0001) and 144h (P<0.001) were observed.

Figure 2.13 - Mean metabolic activity of HCASMCs treated with different concentrations of curcumin for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.14 - Mean metabolic activity of HCASMCs treated with different concentrations of magnolol for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.15 - Mean metabolic activity of HCASMCs treated with different concentrations of Ferulic Acid for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.16 - Mean metabolic activity of HCASMCs treated with different concentrations of Exendin-4 for 24h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

2.3.2.2 Porcine Coronary Artery Smooth Muscle Cells The metabolic activity results for the porcine coronary artery smooth muscle cells (PCASMCs) exposed to curcumin for 18h, 24h 36h, 48h, 72h and 144h as illustrated in Figure 2.17. The PCASMCs treated with 5µM and 10µM of curcumin displayed a significant increase in metabolic activity at 18h, compared to the untreated control (P<0.0001 and P<0.001, respectively). The 100µM sample at the 18h time point exhibited a diminishing effect on the metabolic state where the activity level dropped to 17% (P<0.0001 vs untreated control). This occurrence transcended across all the time points, where there was no difference between the 100µM treated cells and the No Cell negative control. The 5µM treated sample at 24h and 36h demonstrated the similar boost in metabolic activity as initially observed at 18h. Throughout all the time points, the cells treated with 1µM remained unchanged with no significant increase or decrease in metabolic activity. The 5µM did display in increase in metabolic activity with no major changes between the exposure times after the initial 18h observation. For the 10µM treated cells, changes between the exposure times were observed between 24h, 36h, 48h and 72h compared to the final 144h outcome (P<0.001, P<0.0001, P<0.01 and P<0.01, respectively). Visually, the 25µM treated cells displayed a decreasing trend as exposure time was increased. This was reaffirmed by the increasing level of statistical significance observed between 18 and 48 (P=0.046), 72 (P<0.001) and 144h (P<0.0001), as well as 24h vs 72h (P<0.001) and 144h (P<0.0001). A change between 36h and 144h (P=0.02) was also seen. Throughout the course of the assessment, the cells treated with 25µM of curcumin went from a metabolic activity level of 73%, linearly down to 1.8%.

The metabolic activity results for Magnolol are shown in Figure 2.18. Under Magnolol treatment the PCASMCs demonstrated a similar activity compared to the human cells. The high concentrations of Magnolol (187µM and 375µM) reduced significantly the activity of the cells in all exposure times tested, with the results reaching the readings of the negative (No-Cells) control. On the other hand, the cells treated with 37.5µM of Magnolol, exhibited a massive increase of their metabolic activity for all exposure times, even reaching 340% of the activity of the untreated control at 144h (P<0.0001). The 93.9µM concentration did not induce any significant changes in cell activity for exposure times up to 72h, but it did induce a significant decrease after 144h of exposure down to a minimal metabolic state (P<0.0001). A slight increase to 131% was observed at 36h for the 93.9µM concentration level and displayed a statistical difference relative to the untreated control (P=0.020). The effect of 3.75µM on the improvement of metabolic activity was observed at 72h and remained unchanged. Significant differences between all the preceding time points and the 72h was observed. The 37.5µM treated cells caused the metabolic activity of the cells to statistically decrease from 18h to 24h (P=0.023). An increase between the metabolic activity at 48h and 72h was observed as well as from 72h to 144h (P<0.0001 for both comparisons). The 93.9µM treated cells showed insignificant changes for the major

part of the experiment, until the drop from 72h to 144h (P<0.0001). For 187 and 375 µM, the metabolic activity remained at a minimum and unchanged between all exposure times (P>0.99).

The metabolic activity of PCASMCs, exposed to ferulic acid is presented in Figure 2.19. None of the concentrations demonstrated any changes in metabolic activity compared to the untreated control sample at the respective time point upon treatment with ferulic acid.

The metabolic activity of PCASMCs treated with an increasing concentration of exendin-4 is illustrated in Figure 2.20. Upon 18h of exposure to the drugs, the 50nM treated samples were the only group to display a significant change. The metabolic activity increased to 139% and was statistically significant compared to the untreated sample (P<0.01). For 24h and 36h of exposure, the cells demonstrated no response to exendin-4. At 48h, the 50nM treated cells demonstrated another statistical significant difference (to 140%), however, there was no change between the 18h metabolic activity and the 48h time point. At 72h of exposure, the cells treated with 5nM, 25nM and 50nM all displayed an increase in metabolic activity to 137%, 136% and 174% (P=0.015, P=0.021 and P<0.0001, respectively and compared to untreated control). These values remained unchanged after 144h of exposure, but still demonstrated a significant difference compared to the control group of P<0.01, P<0.01 and P<0.0001 for 5nM, 25nM and 50nM, respectively. The 50nM treated cells demonstrated the greatest change amongst the time points, where significance was observed between 48h and 72h (P=0.030).

Figure 2.17 - Mean metabolic activity of PCASMCs treated with different concentrations of curcumin for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.18 - Mean metabolic activity of PCASMCs treated with different concentrations of magnolol for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.19 - Mean metabolic activity of PCASMCs treated with different concentrations of ferulic acid for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

Figure 2.20 - Mean metabolic activity of PCASMCs treated with different concentrations of Exendin-4 for 18h, 24h, 36h, 48h, 72h and 144h. The results were normalised to the untreated control group (100%) and baseline corrected to the No Cells negative control. The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control (black) or the negative control (red), at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels. Brackets indicate groups with the same level of significant difference to the untreated or negative control.

2.3.3 Wound Healing Assay Cell migration and morphology was assessed using the wound healing assay. The percentage reduction of the inflicted wound area by the PCAECs, treated with the different drug concentrations, is shown in Figure 2.21. The drug concentrations were chosen specifically as a result of the initial metabolic activity screening. Everolimus at a high concentration provided sufficient inhibitive capacity to be used as the negative control. Median concentration values of magnolol and curcumin were assessed for the purpose of additional verification of the metabolic activity study. Ferulic acid and exendin-4 were studied at 50µM/500µM and 10nM/100nM, respectively, as an effective mid-range value as well as an extreme value based on the promising data observed in the metabolic activity study. Overall, the reduction in the wound area achieved by the magnolol-, curcumin- and everolimus-treated groups was minimal and significantly reduced compared to the untreated control (p<0.05 for all time points). On the other hand, the cells treated with both concentrations of ferulic acid demonstrated very similar behaviours to the untreated control (p>0.05 for all time points), with almost 90% of the wound area repaired by the end of the experiment (50h). The cells that were treated with the two exendin-4 concentrations demonstrated a comparable behaviour to the untreated control, but the migratory rate

of these cells was significantly slower compared to both the ferulic- acid-treated and untreated cells (p<0.05).

The calcein staining for the cell viability of the treated and untreated PCAECs that were subjected to the scratch assay after 48h exposure are shown in Figure 2.22. 100 µM everolimus, 10µM curcumin and 93.9µM magnolol displayed no viable cells after 48h. The untreated samples demonstrated a slight increase in cell size over the healed wound area as did the exendin-4-treated samples and the samples treated with 50µM ferulic acid. Both the ferulic-acid- and Exending-4-treated samples demonstrated 100% confluency over the wound region, apart from the cells treated with 100nM of Exendin-4. All viable samples demonstrated a cobblestone-like cell morphology.

The percentage reduction of the inflicted wound area by the PCASMCs, treated with the different drug concentrations, is shown in Figure 2.23. In this case, the PCASMCs did not show the same migratory pattern and ease of visualisation as the PCAECs. The larger, more elongated cell shapes of the PCASMCs made it difficult to get an accurate representation of the wound closure. The only cell group to show no gap closure was the one treated with 100µM of everolimus. At 12h, all treated groups (except everolimus) demonstrated no significant difference compared to the untreated control. At 24h, the PCASMCs treated with 500µM ferulic acid achieved approximately 40% reduction in the wound area, whereas the groups treated with 10nM Exendin-4, curcumin and magnolol demonstrated about 22% reduction in the wound area (p<0.01). After 36h of drug exposure, the untreated group showed a reduction of the wound area of about 60%, whereas the ferulic acid samples achieved below 50% (p<0.01) and the Exendin-4 samples below 40% (p<0.0001). After 48h, the exendin-4- and ferulic-acid- treated groups increased their migratory rate, and no significant difference could be observed between them and the untreated control beyond 24h. On the other hand, the magnolol- and curcumin- treated samples demonstrated significantly reduced wound close capacity beyond 24h compared to the untreated control (p<0.0001 and p≤0.05, respectively).

The calcein staining results for the treated and untreated PCAECs that were subjected to the scratch assay after 48h exposure are shown in Figure 2.24. In contrast to the PCAECs, the PCASMCs did not migrate in a controlled manner and were scattered in the wound gap, except for the case of the everolimus-treated samples that presented no viable cells. The curcumin-treated cells demonstrated a significant change in their morphology, compared to the untreated samples. The untreated samples exhibited a rhomboid-shaped morphology with some regions of cells exhibiting elongation. The magnolol-treated cells also demonstrated this polygonal morphology, with a smaller size and shorter length. The exendin-4- and ferulic-acid-treated PCASMCs showed a very clear morphological shift towards a sharply elongated shape. All viable samples demonstrated migration, but none of them achieved full coverage of the wound area.

Figure 2.21 – Percentage reduction of wound area achieved by PCAECs under different drug exposure over a 48h period. The data was expressed as mean of the reduced wound area for each time point (n=3)

Figure 2.22 - Fluorescence microscopy of calcein-stained PCAEC monolayers, subjected to the scratch assay, after 48h exposure to different drug dosages. Scale bar: 200µm.

Figure 2.23 - Percentage reduction of wound area achieved by PCASMCs under different drug exposures over a 48h period. The data was expressed as mean of the reduced wound area for each time point (n=3)

Figure 2.24 - Fluorescence microscopy of calcein-stained PCSMCs, subjected to the scratch assay, after 48h exposure to different drug dosages. Scale bar: 200µm.

2.3.4 Gene Expression of Cells Treated with Drugs In order to identify any potential phenotypic changes that might be induced by the drug treatments on to the cells, qPCR was conducted. Primers were designed specifically for the porcine endothelial and smooth muscle cells for genes pertaining to the respective cell type.

2.3.4.1 Endothelial Cells The endothelial nitric oxide synthase (NOS3), vascular endothelial cadherin (CDH5) and platelet endothelial adhesion molecule (CD31) expressions of the PCAECs treated with the different drug compounds for 144h were compared to the untreated control. The relative fold changes in the gene expression of the treated cells with respect to the untreated control are shown in Figure 2.25. The expressions of CDH5 and CD31 in the treated cells demonstrated no significant difference compared to the untreated the control (p=0.6426 and p=0.1923, respectively). The PCAECs that were treated with 100µM of Exendin-4 demonstrated approximately 50% downregulation of NOS3, which was significantly reduced compared to the NOS3 expression of the cells treated with 10µM exendin-4 and 500µM ferulic acid.

A B

Figure 2.25 – Fold change in CDH5 (A), CD31(B) and NOS3 (C) expression in PCAECs after 144h exposure to Exendin-4 10nM (Ex10), Exendin-4 100nM (Ex100), Ferulic Acid 50µM (FA50) and Ferulic Acid 500 µM (FA500), compared to the untreated control. Data was expressed as means ± 95% C.I. (n=3). Asterisks indicate significant difference between the originator and end-bracket column at the 0.05 (*) level.

2.3.4.2 Smooth Muscle Cells The myosin heavy chain 11 (MYH11), smoothelin (SMTN), α-Actin (ACTA) and calponin (CNN1) expressions of the PCASMCs treated with the different drug compounds for either 24h, 48h or 72h were compared to the untreated control. The fold change in the expression of the aforementioned genes is shown in Figure 2.26. The PCASMCs treated with 50µM of ferulic acid demonstrated a significant increase in the expression of all target genes, with MYH11 increasing 5-fold, ACTA 4-fold and SMTN and CNN1 2-fold after 48h, compared to the corresponding expressions of the untreated cells. For the higher concentration of ferulic acid (500µM), a 2.7-fold increase in the expression of MYH11 and a 5-fold increase in the expression of SMTN and ACTA was observed after 48h, compared to the untreated sample. SMTN exhibited a higher fold change at 500µM than at 50µM of ferulic acid. ACTA demonstrated a further increase to 10-fold and at the 72h exposure to 500µM of ferulic acid, whereas the expression of CNN1 exhibited no change over the course of 72h. In the case of the PCASMCs treated with exendin-4, the gene expression increased at 48h and decreased after an additional 24h. The CNN1 expression of the PCASMCs treated with 10nM exendin-4 and the SMTN and CNN1 expressions of the PCASMCs treated with 100nM exendin-4, did not exhibit this trend. The expression of SMTN and CNN1 in the PCASMCs treated with 100nM exendin-4 displayed no significant changes over the 72h.

2.3.5 Endothelial Cell Function The flow cytometry results of the functionality assessment of the PCAECs treated with ferulic acid and exendin is shown in Figure 2.27. Both treatments demonstrated an improved viability compared to the serum free control. It was also noted that around 98% of the viable cells were positive for CD31 for all groups. The study also showed that all of LDL positive cells were also positive for CD31.

Figure 2.26 – Fold change in MYH11, SMTN, ACTA and CNN1 expression in PCASMCs after 24h, 48h and 72h exposure to Ferulic Acid 50µM (A), Ferulic Acid 500 µM (B), Exendin 10nM (C) and Exendin 100nM (D), compared to the untreated control. Data was expressed as means ± 95% C.I. (n=3). Asterisks indicate significant difference between the originator and end-bracket column at the 0.05 (*), 0.01 (**), 0.001 (***) and 0.0001 (****) level.

Figure 2.27 - Flow cytometry results of PCAECs cultured in to Exendin 10nM (Ex10), Exendin 100nM (Ex100), Ferulic Acid 50µM (FA50) and Ferulic Acid 500 µM (FA500). Data was expressed as means ± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the serum-free control at the 0.05 (*) and 0.01(**) level.

2.4 Discussion The work presented in this chapter aimed at assessing the cytotoxic limits of a cohort of drug compounds, as well as their effectiveness in inhibiting SMCs and promoting ECs proliferation. Cell morphology, phenotypic alterations and cell migration were assessed in order to ensure that the drug compounds did not negatively alter the functionality and state of the cells. The CCK-8 metabolic activity assay that utilised the WST-8 salt was utilised in order to assess cell viability and population in terms of the metabolic activity. The high susceptibility of the WST-8 salt to reduction made it very sensitive to the metabolic activity of the cells through mitochondrial dehydrogenases within only living cells. Although the MTT assay could also provide such information about the cells, it has been reported to be less sensitive than WST-8, as well as more cytotoxic, since that assay needs to enter the cell membrane. This is problematic for long term assessments of cells or if subsequent analysis needs to be conducted [225]. The assessment of the metabolic activity investigated both human and porcine cells in order to identify potential cross-species observations, with a view to potentially using porcine cells as readily available alternative cell source to model stent endothelialisation. HUVECs were chosen due to their reported similarity to HCAECs as well as being used as a research standard [226]. There are many advantages in using HCAECs in the assessment of the endothelialisation potential of stents, such as an accurate coronary artery endothelial cell susceptibility to inflammatory markers [227]. HUVECs have also been reported to have a higher level of adhesion molecule expression than the coronary endothelial cells [227] and be more sensitive to apoptosis than HCAECs [228].

The commercially available drugs (paclitaxel and everolimus) used in DESs were assessed to identify the effects of these anti-proliferative compounds on both ECs and SMs. The effect of these drugs on the metabolic activity of HUVECS and HCASMCs indicated a clear difference in terms of sensitisation, since the SMCs, treated with both paclitaxel and everolimus, exhibited minimal disruption in their viability and metabolic state, whereas the HUVECs, treated with the same concentrations of these drugs, showed clear signs of apoptosis. These results might explain the poor endothelialisation of DESs in vivo. The 10µM dosage of both drugs on SMCs showed a proliferation decline of around 25%, whereas the ECs exhibited complete cell death with no noticeable differences compared to the WST- 8 No-Cell control. It is important that an optimum drug for DESs promotes proliferation and viability of ECS whilst reducing the proliferation of SMCs.

Curcumin demonstrated an increasingly toxic effect on the HUVECs, except for the inconsistent result observed when the HUVECs were treated with 100µM. The same trend was present in the PCAECs, suggesting that the cell source might not be the critical factor here. It was hypothesized that the distinct yellow colour of curcumin might have interfered with the metabolic activity signal due to the high concentration of 100µM. However, this could not have been the case, since the same

concentrations were applied to SMCs and this inconsistent behaviour was not observed. A study by Koo et al. [229] reported a rapid decline in cell viability when increasing the curcumin concentration above 5µM. They reported that a concentration of 100 µM was highly cytotoxicity on HUVECs. Their study assesses HUVECs that were treated with vascular endothelial growth factors (VEGF) and the apoptotic nature of ECs under the treatment of curcumin was linked to the blocking of the VEGFR-2 mediated ERK1/2 signalling pathway, which is a critical pathway in EC proliferation. Since VEGF exists in EGM-2, curcumin might have induced suppressive effects on the VEGF receptor. It has been reported that curcumin inhibited the VEGF proliferative pathway through the inhibition of COX-2 and MAPK. Fu et al. [230] suggested that HUVEC apoptosis by curcumin might be caused by the activation of caspase-3.

The apoptotic effect of curcumin on SMCs has been linked to the activation of nuclear factor-κB, as well as to the inhibition of protein kinase activity and AP-1. It has also been suggested that serum and growth factor-stimulated vascular SMCs can have their proliferation inhibited by curcumin [231]. The study suggested that the anti-proliferative effect might also be due to the induction of the apoptotic process through the MAPK p38 pathway. The authors also claimed that curcumin did not cause an anti- proliferative effect on HUVECs, which contradicts the findings of the present work, as well as of previous publications [229][230]. This could be due to the low calculated curcumin release per day, of 1.8ng/ml which could be too low to cause any inhibitory effect. The results obtained in the present work demonstrated that treatment with 1µM curcumin had no negative effects on both SMCs and ECs. However, and based on these results, curcumin was discarded from further assessments due to the lack of any tangible beneficial effects on the tested cells. No concentration of curcumin tested in present work suggested an indication of improved endothelial proliferation whilst inhibiting smooth muscle cell proliferation. Curcumin has been extensively studied as a treatment for cardiovascular disease due to its anti-inflammatory and antioxidant effects. Several studies have shown the reduction of intercellular adhesion molecule 1 (ICAM-1) in ECs upon short dosage times (under 24 hours). Increased ICAM-1 expression of ECs within atherosclerotic lesions has been observed, indicating the potential of curcumin as an anti-inflammatory therapy [232].

Studies have shown a dose dependent effect of Magnolol on SMCs, inducing apoptosis [174]. The results obtained in the present study with human SMCs agree with these findings from previous studies. However, porcine SMCs demonstrated an opposite reaction and exhibited a highly significant increase in the metabolic activity of the cells treated with the two lower concentrations (3.75µM and 37.5µM). This could be potentially attributed to cross-species differences. Although previous studies have suggested the porcine model to be an appropriate alternative for vascular pathological studies [233], this was not apparent in the results of the present work. The ECs that were treated with

magnolol demonstrated that any concentration higher than 3.75µM led to cell apoptosis. As such magnolol was not assessed further in the present study.

On the other hand, ferulic acid induced both HUVEC and PCAEC proliferation, indicating its potential to be used as a proliferation promoting substance. These results were in accord with previous reports that reported a maximum EC proliferation upon under a 50µM concentration of ferulic acid. This proliferative capacity has been associated with an increase in cyclin D1 and VEGF expression in ECs [234]. The WST-results in the present study also suggested that ferulic acid increased more the proliferation of ECs than SMCs. Although SMCs were not inhibited by ferulic acid, most of the treated cell groups showed a maintained metabolic state, particularly under the 500µM concentration. This suggested that ferulic acid might represent a plausible strategy for ensuring faster endothelialisation whilst maintaining SMC proliferation.

Exendin-4 demonstrated a rapid increase in cellular metabolic activity in all concentrations used. However, the HUVECs treated with 50nM Exendin-4 exhibited a loss of proliferative capacity compared to the untreated control. The PCAECs, on the other hand, benefited from this high dosage by exhibiting a significant metabolic boost (almost double compared to the untreated control). The SMCs also exhibited an increase in metabolic activity, but not to the same extent as the ECs, thus making Exendin- 4 an appropriate candidate for drug eluting stents. Due to the GLP-1 receptor expressed in ECs and exendin-4 being a GLP-1 receptor agonist, the latter has been reported to cause a proliferative response due to PKA-PI3K/Akt-eNOS activation through the GLP-1 receptor-mediated pathway [195]. Despite the ERK1/2 pathway being reported as a cell proliferative pathway, Erdogdu et al. (2010) [195] suggests that it does not play an important role in the proliferation of ECs upon exendin-4 treatment. Phosphorylation of ERK1/2 leads to the cell proliferation motif and is induced by the addition of angiotension II. Studies have shown the inhibition of phosphorylation by treating angiontensin II- treated-SMC with exendin-4 and thus causing an reduction in SMC proliferation[191,235].

There were no significant difference in cell migration rate of PCAEC when comparing the control untreated samples with exendin-4 and ferulic acid treated cells. However, the post-scratch cell viability assay provided an insight into morphological changes. As expected from the initial metabolic activity assessments, curcumin, magnolol and everolimus treated cells displayed no cell viability. The larger endothelial cells observed within the migration zone of the exendin-4 and control PCAEC were attributed to the process of cell migration. This enlargement of the cell can be associated with the extension, attachment and contraction stages of endothelial cell migration as seen in Figure 2.28.

Figure 2.28 – Model of migration mechanism of endothelial cells. Adopted from Lamalice at al. (2007) [236]

The visual cobblestone-like monolayer morphology has been reported to be a characteristic of statically cultured ECs and provide an indication that the cells have not undergone differentiation upon drug exposure [237]. The status of the ECs was verified by assessing the genetic expression of CD31, CDH5 and NOS3. The lack of any significant differences in the fold change expression relative to the untreated control suggested that the tested drug compound had no effect on EC differentiation. Although these markers are not entirely EC-specific, previous studies have suggested that when assessed in combination were effective markers for assessing EC purity after stem cell differentiation [238]. CDH5, or VE-cadherin, has been reported to be an important protein associated with the maintenance of the restrictive barrier of the endothelium [239]. CD31, also known as platelet endothelial cell adhesion molecule (PECAM-1) has been identified at intercellular junctions, whereas NOS3 (or eNOS ) can be attributed to the production of nitric oxide in endothelial cells which is crucial for vasomotor tone control [240].

The morphology of the untreated PCASMCs demonstrated a rhomboid structure, which has been described to be characteristic of a synthetic SMC phenotype [15]. A synthetic phenotype exhibits higher proliferative capacity with a stronger affinity to migration than the contractile phenotype. The contractile phenotype morphology has been reported to assume a spindle-shape, with a non- proliferative function [15]. The contractile phenotype has been reported to expresses certain markers that the synthetic, undifferentiated phenotype does not, such as α-smooth muscle actin, smooth muscle – myosin heavy chain and calponin [241]. It has also been reported that SMCs might exist in- between both phenotypes as described in Figure 2.29 [242] . It is clear that the SMCs can perform both

functions and that a clear indication of a purely contractile state lies with the expression of MYH11 and SMTN. CNN1 is an actin-binding protein associated with SMC contraction through regulation of myosin ATPase activity [243]. SMTN has been reported to be present in fully differentiated SMCs, key in muscular pulsating blood vessels and directly influence contractile potential [244]. MYH11 has been reported to be a critical protein for cell motility, cytoskeleton and material transport [245]. ACTA has also been described as directly involved in structure, cell mobility and generation of contractile forces.

Figure 2.29 - Expression levels of SMC-associated genes for the two phenotypes. Adopted from Rensen et al. (2007) [242]

In the present study, the exendin-4-treated PCASMCs exhibited a mixed population of spindle-shaped and rhomboid-shaped, indicating cross-functional SMCs, whereas the ferulic-acid-treated PCASMCs appear to have a larger population of contractile SMCs. The upregulation of the smooth muscle cell contractile markers in the PCASMCs, treated with ferulic acid, alongside the spindle-shaped morphology indicated the occurrence of SMC differentiation. The qPCR data for exendin-4 treated PCASMC displayed an upregulation of contractile markers after 24 hours of drug treatment. After 72 hours, these markers demonstrated reduction of genetic expression reverting back to the initial value. This may be due to the occurrence of SMC dedifferentiation back to the synthetic phenotype and was visually verified by the observation of the rhombus-shaped cells. In addition, the WST-8 study showed further validation of this theory as an increase in metabolic activity was seen amongst exendin-4 treated PCASMCs after 72 hours.

Low density lipoproteins have reported to be important intracellular carriers of cholesterol and are absorbed by receptor mediated endocytosis [246]. One of the key functions of ECs is the uptake and catabolism of low-density lipoprotein [247], [244]. As such, LDL uptake has been described as a potent assessment of EC function [249]. In the present study, flow cytometry was used to assess LDL uptake of ECs exposed to the different drug compounds. A study by Gaffney et al. assessed the uptake of native LDL (nLDL) and acetylated (Ac-LDL) by bovine aortic ECs and observed a higher uptake amongst

sub confluent cells compared to a confluent endothelial monolayer [250]. Another study verified this by showing a 5- to 10-fold decrease in receptor-bound native LDL as cells progressed from a proliferating state to a quiescent confluent morphology [251]. On the other hand, modified Ac-LDL showed the opposite effect, with confluent EC monolayer demonstrating an increased amount of bound Ac-LDL. Confluent ECs have been reported to internalise and degrade Ac-LDL at a high capacity via a scavenger-receptor-mediated mechanism [251]. ECs have a higher affinity in up-taking modified LDL (acetylated or oxidised). This LDL uptake process has been reported to be a potent marker for ECs due to the cellular functional specificity amongst vascular ECs [252].

It has been reported that native LDL might become partially oxidised and hence, allow for EC uptake of native LDL amongst confluent drug-treated cells [253]. Moreover, it has been reported that LDL induces cytotoxicity in ECs when cultured in serum-free media [254]. This might be due to the oxidation of the LDL and can be prevented by treating the cells with anti-oxidants [255]. The flow cytometry results produced in this work showed the prevention of reduced cell-viability when cultured with exendin-4 and ferulic acid. Ox-LDL was reported to induce EC apoptosis through a death receptor (Fas) that produced an apoptotic signal once activated. Lysophosphatidylcholine (LPC) is a major constituent of oxLDL and causes EC sensitisation to the Fas-mediated apoptosis. Sata and Walsh [256] have reported the downregulation of the FLIP protein, upon treatment of ECs with oxLDL. The FLIP protein was shown to control the Fas-mediated apoptosis [256]. High LDL concentrations have the potential to damage EC function and might affect the EC-SMC interface in vivo [257]. Ferulic acid has been deemed a powerful antioxidant due to its unsaturated side chains, which readily forms a stabilised phenoxy radical [258]. The scavenging potential of ferulic acid might also alleviate hyperoxia-induced apoptosis [259]. The signalling pathway (Figure 2.30) defining the effect of exendin-4 on the cells indicate that the antioxidant capacity increases upon treatment is due to the cAMP, PI3K and PKC pathway and might be the reason for exendin-4-treated cells demonstrating an improved cell viability compared to the serum-free control.

The assessment on the different drugs and their concentrations provided an early indication as to their possible effect on endothelialisation and SMC inhibition. Ferulic acid and exendin-4 demonstrated a strong endothelialisation capacity compared to the proliferation of SMCs. It was important to identify a drug that poses a minimal effect on the SMC without causing complete cell apoptosis. It is key that the SMCs in the stented region do not proliferative excessively but should maintain a healthy quiescent phenotype. There are many SMC-EC interface molecules that need to be accounted for before proceeding to in vivo assessments. The work done in this chapter provides an early systematic workflow which can be used to identify therapeutic agents that have the

potential to be used within drug eluting stents in order to combat the effects of restenosis and thrombosis.

Figure 2.30 - Signalling pathway demonstrating the antioxidaive potential of exendin-4. Adopted from Oh et al. (2017) [260]

Chapter III: Polymer Assessment

3.1 Introduction The coating of stents and other medical devices is the interface between the device and the surrounding tissue in vivo and, thus, it must not induce any cytotoxic effects. Moreover, and in the case of blood contacting devices, such as stents, the coating should ideally promote EC adhesion and proliferation, which for the case of DESs, are closely related to effectively tailoring the drug- encapsulating and drug-releasing properties of the coating. Poly(D,L-lactic-co-glycolic) acid (PLGA) has been commonly used as a coating of DESs for clinical use. However, PLGA is a hydrophobic polymer with poor cell adhesive properties. In order to improve cell adhesion, the present work investigated the development of a novel stent coating by blending PLGA with a tobacco-plant-derived recombinant human collagen Type I (rHCol1) or rat tail collagen Type I (rtCol1). This chapter describes the work conducted to optimise the process of preparing and assessing the PLGA-collagen blends as coatings for DESs that could support endothelialisation. Rapid adhesion studies were also performed in order to identify whether the PLGA-collagen coatings were suitable for in situ endothelialisation.

3.2 Material and Methods 3.2.1 Coating Preparation PLGA 50:50 (23987, Polyscience, Hirschber, Germany) was weighed using a balance (Extend, Sartorius, Germany) and placed into a 5ml glass vial fitted with a Teflon screw cap. Appropriate volumes of 1,1,1,3,3,3-Hexafluoro-2-propanol (HFIP) (804515, Merck) was pipetted into the glass vial in accordance to 1% (w/v) PLGA-HFIP. The glass vial containing the PLGA solution was tightly sealed and placed onto a shaking platform (Shaking Incubator 3031, GFL, Germany), and was shaken at 150rpm for 1h. For the collagen-PLGA blend samples, the collagen solution, rtCol1 (5mg/ml in 20mM acetic acid, 3440-100 Cultrex) and rHCol1 (3.2mg/ml in 10mM hydrochloric acid, Collplant Ltd, Israel) was placed on ice under a laminar flow cabinet and the respective collagen solution was added to the dissolved PLGA solution (1% (w/v)) in accordance to the sample constituent tables for each described experiment. The calculated collagen volumes were based on the mg/ml concentrations for rtCol1 and rHCol1 listed by the supplier in order to achieve the desired w/w% with PLGA. The glass vials containing the collagen-PLGA blend samples were placed in a beaker with ice and onto the shaking platform and was shaken at 150 rpm for 1h. The ice was replaced after 30min. For sample neutralisation, 100mM of sodium hydroxide (NaOH) was added to the collagen-PLGA blend based on the required volume of rhCol1 solution used (as listed in the respective methods). rHCol1 is supplied at 3.2mg/ml dissolved in 10mM of hydrochloric acid. The samples were further shaken on ice for 15min prior to transferring the calculated volumes (as listed for each subsequent coating experiment) to the respective multiwell culture plates for coating the bottom of the well. The deposited solution volumes were calculated to

an approximated final deposited coated thickness of 10µm by using the well surface areas (9cm2 for 6 well plate, 2cm2 for 24 well plate and 0.9cm2 for 48 well plate). The samples were washed with 70% (v/v) ethanol and subjected to UV sterilisation before any cell seeding.

3.2.2 Assessment of Cell Attachment The affinity of the developed coatings to cell attachment was assessed by seeding PCAECs on culture plates coated with the each of the developed coatings. The bottom of culture wells were coated as previously described using the values given in Table 3.1 for an individual well. A 0-100% (w/w) range of added collagen was prepared and applied to the wells in a 24 well plate in triplicate.

Table 3.1 - Sample and constituent list for a single well coating (in 24 well plate).

Sample PLGA (mg) HFIP (µl) Volume of Collagen (µl) (rtCol1 and rHCol1) rtCol1 N/A N/A 520 rHCol1 N/A N/A 813 PLGA 2.68 268 N/A 20% (w/w) 2.23 223 90/140 40% (w/w) 1.91 191 153/240 60% (w/w) 1.68 168 201/314 80% (w/w) 1.49 149 239/372 100% (w/w) 1.34 134 268/419

For each coating, 3 wells of a 24-well-plate were used, the bottom of which was coated with the respective coating. The coated wells were seeded with PCAECs at P5 and cultured in EGM-2 with 1%

(v/v) P/S under standard culture conditions (37°C, 5% (v/v) CO2). One well from each coating group was cultured for 30min and the second for 1h. Following culture, the medium from the wells was aspirated and the plates were washed with DPBS. Subsequently, 600µl of 10% (v/v) WST-8 in HBSS was added to each well and incubated for 2h before measuring the absorbance. The wells were then washed again with DPBS and supplemented with EGM-2 with 1% (v/v) P/S. The assessed wells were placed in the incubator and cultured for a further 72h at 37°C and 5% (v/v) CO2. Subsequently, calcein staining was performed on the wells, as described previously, in order to assess the viability of the PCAECs seeded onto the coated wells.

3.2.3 Assessment of the Gelation and Preparation of rHCol1 The gelation and preparation method of rHCol1 was assessed in terms of its effect on the metabolic activity of seeded ECs. Three different gelation temperatures were assessed, including 4°C, room temperature (RT) and 37°C, together with the effect of NaOH as neutralisation agent. Each gelation temperature was assessed in triplicate, with and without NaOH as the neutralising agent. For each

sample, 3 wells of a 6 well plate was used, the bottom of which was coated with rHCol1. A volume of 3.656 ml of rhCol1 was used per well in order to achieve an approximated coating thickness of 10µm. For the neutralised samples, rHCol1 solution was supplemented with 365 µl of 100mM NaOH and shaken on ice for 30min before it was transferred to the respective wells for curation under the different temperatures. All samples were allowed to gelate for 18h before they were washed three times with PBS.

Subsequently, PCAECs cultured to P5 were seeded at a density of 16,000 cells/cm2 in EGM-2, with 1% (v/v) P/S, in the coated wells. Triplicates of uncoated wells were also seeded with the same density of

PCAECs and used as control. The well plates were then cultured at 37°C and 5% (v/v) CO2 until the uncoated wells had reached 80% confluence. The WST-8 metabolic activity assay (1.5ml/well) was, subsequently, used on the samples, as described previously. Following addition of the WST-8 solution, the wells were incubated for a further 4h, prior to transferring 200µl (in technical replicates of six) of the WST-8-supplemented medium from each well to a corresponding well of a 96-well plate for measuring the absorbance at 450nm and 630nm (background) using a plate reader.

3.2.4 Assessment of Coating Cytotoxicity The metabolic activity of the cells seeded on the various coatings provides an indication of cell proliferation. PLGA with increasing concentrations (0%, 5%, 10%, 15% and 20% w/w) of rtCol1 and rHCol1 were prepared (as described above) using the constituent volumes listed in Table 3.2 and used to coat the bottom of a 48-well plate in triplicates. rHCol1 samples without the neutralisation agent were also assessed. The wells were solvent cast under a fume hood for 48 hours and then washed with 70% (v/v) ethanol, and subject to UV sterilisation under a laminar flow cabinet.

Table 3.2 - Sample and constituent list for a single well coating (in 48 well plate)

Sample PLGA HFIP (µl) Volume of Collagen (µl) Volume of 100mM NaOH (for (mg) (rtCol1 and rHCol1) rHCol1) (µl) – if applicable rtCol1 N/A N/A 247 N/A rHCol1 N/A N/A 386 38.6 PLGA 1.27 127 N/A N/A 5% 1.21 121 12/19 1.9 (w/w) 10% 1.15 115 22/36 3.6 (w/w) 15% 1.08 108 32/51 5.1 (w/w) 20% 1.02 102 41/64 6.4 (w/w)

PCAECs were cultured to P6 and seeded at 8,000 cells/cm2 onto the coated well plates. Control Uncoated samples were also seeded with the same PCAEC density and used as controls. The seeded well plates were cultured in EGM-2 with 1% (v/v) P/S for 48h under 5% (v/v) CO2 at 37°C. The medium in the wells was aspirated and the wells were washed three times with PBS, before adding 500 µl of 10% (v/v) WST-8 in HBSS solution for assessing the metabolic activity of the PCAECs. The well plates were then incubated for a further 4h before transferring 150 µl (with technical triplicates) of the WST- 8-supplemented medium from each well to a corresponding well of a 96-well plate for measuring the absorbance at 450nm and 630nm (background) using a plate reader. The absorbance values were normalised to the values of the PLGA-only samples in order to identify the benefit of the integration of collagen.

3.2.5 Fourier Transform Infrared (FTIR) Spectroscopy Coated glass cover slips were used for infrared spectroscopy. The cover slips were coated using the procedure described in §3.2.1. All rhCol1 samples were neutralised as described in §3.2.3. The cover slips were coated with the following compositions:  rHCol1  rtCol1  PLGA  PLGA – 20% (w/w) rHCol1  PLGA – 20% (w/w) rtCol1  PLGA – 20% (w/w) rHCol1 – Drug loaded with Exendin-4 (7, 35 and 70 ng/mg)  PLGA – 20% (w/w) rtCol1 – Drug loaded with Exendin-4 (7, 35 and 70 ng/mg)  PLGA – 20% (w/w) rHCol1 – Drug Loaded with Ferulic Acid (0.16, 0.8 and 1.6 % (w/w))  PLGA – 20% (w/w) rtCol1 – Drug loaded with Ferulic Acid (0.16, 0.8 and 1.6 % (w/w)))

A PerkinElmer 100 FTIR spectrophotometer (PerkinElmer) with an attenuated total reflection (ATR) module, using a 1mm2 diamond/ZnSe crystal and a mercury cadmium telluride detector was used for the spectral measurements. The resolution was set at 4cm-1 with 8 accumulations. The spectra were recorded in the scanning range of 650 - 4000cm-1 in absorption mode. Sample analysis was conducted using the OMNIC software. The recorded absorption peaks represented the frequency of vibrations between certain bonds between the atoms of the material. The peak size indicated the quantity of the specific bond within the material.

3.2.2 Statistical Analysis The data was plotted as mean ± 95% confidence interval. All data were analysed using Graphpad Prism 6. ANOVA testing followed by Bonferroni multiple comparison post hoc testing were performed. A statistical significance was recognised as P < 0.05.

3.3 Results 3.3.1 Assessment of Cell Attachment The mean metabolic activity of the PCAECs seeded onto the different coatings after 30min and 1h is illustrated in

Figure 3.1. At 30 minutes, both the PLGA-rHCol1 and PLGA-rtCol1 blended samples demonstrated an increase in metabolic activity compared to the PLGA-only group. Compared to the untreated (uncoated) control, the rHCol1-only (P=0.011) and the 40% (p<0.01), 60% (p<0.001) and 100% (P=0.030) rHCol1 blends demonstrated a significant increase in the metabolic activity. The rtCol1 displayed a visible increasing trend in metabolic activity with increasing collagen concentration. However, only the 60% and 100% rtCol1 blends demonstrated a significant increase (P=0.020 and P<0.0001, respectively) compared to the untreated control. Considering that the 30min and 1h samples had independent culture plates, the results of the 1h samples for rHCol1 showed a high similarity to the 30min ones but with a higher metabolic percentage. The increase in metabolic activity amongst the 80% and 100% (w/w) also demonstrated a higher level of statistical significance (p<0.001 and p<0.0001 respectively). The rtCol1 samples also showed a similar trend to their 30 min counterparts. The 40% and 60% rtCol1 blends went from a 50% increase in 30 min (ns (P=0.59) and P=0.020, respectively) to 100% metabolic increase in 1 h relative to the untreated sample (p<0.0001 for both). The significant increase in metabolic activity of the rtCol1 100% (w/w) sample was observed to be insignificant (P=0.56) after 1 h of culture. Another significant observation was the PLGA coated sample increasing in metabolic activity from 54% in the 30 minute sample group to 140% in the 1 h group. Amongst the 1 h well plate, the significant differences between the PLGA coated sample and the PLGA-collagen blended samples (both, rtCol1 and rHcol1) showed a reduction in statistical significance, where only the 60% and 100 % (w/w) (P<0.01 and P<0.0001, respectively) demonstrated an increase over PLGA. For the rtCol1 blended samples, after 1 hour of culture, only the 40% (w/w) and the 60% (w/w) samples demonstrated a statistical significant increase over the pure PLGA sample P<0.05 and P=0.021, respectively). Furthermore, after 1 hour of culture with the rtCol1 samples, there was no significant increase in metabolic activity of the pure rtCol1 samples compared to the PLGA.

Figure 3.1 – Mean metabolic activity of PCAECs seeded onto the different PLGA-collagen blends after (A) 30min and (B) 1h of culture. The rHCol1 samples were not neutralised for this study. The results were normalised to the untreated (uncoated) control group (100%). The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the untreated control at the 0.05 (*), 0.01 (**), 0.001(***) and 0.0001 (****) levels (ns indicates not significant difference).

The calcein staining of the coated samples after 72h of further culture after the WST-8 study is shown in Figure 3.2. The viability of the PLGA and untreated samples displayed a high level of confluency and viability. All rtCol1 samples demonstrated a similar level of viability to the untreated control. The rtCol1-only sample exhibited a higher calcein signal with increased confluency than the PLGA-collagen blends. Among the rtCol1 group, the lowest viability was achieved for the 80% sample. Apart from the 40% rHCol1 blend sample that demonstrated a strong confluency, the rest of the rHCol1 samples showed poor viability.

3.3.2 Assessment of the Gelation and Preparation of rHCol1 The addition of a neutralising agent to rHCol1 during the gelation step was examined by assessing the cell viability of the PCAECs that were seeded onto rHCol1 coatings, prepared at different temperatures with or without neutralisation, after 48h of culture. The results are displayed in Figure 3.3. The addition

3.3.3 Assessment of Coating Cytotoxicity The mean metabolic activity of the PCAECs cultured on rtCol1-PLGA blend coatings for 48h is displayed in Figure 3.4. Pure rtCol1 (P<0.0001) exhibited the highest metabolic activity, which was similar to the untreated control (P<0.0001) and rose to approximately double the metabolic activity measured for the PLGA group. The 5% and 10% groups demonstrated no improvement in metabolic activity (P>0.99) but the 15% and 20% (P<0.01) groups demonstrated a significant increase in metabolic activity to about 150%.

The mean metabolic activity of the PCAECs cultured on rhCol1-PLGA blend coatings without the addition of a neutralising agent are illustrated Figure 3.5. The PCAECs showed a significant loss in metabolic activity upon cultivation on rhCol1. The addition of rhCol1 to the PLGA showed a linear decline in metabolic activity of the PCAECs for increasing rhCol1 concentrations within the polymer blend. Pure rhCol1 (P<0.0001) demonstrated a 50% reduction of metabolic activity relative to the PLGA control group. Both 5% and 10 % (w/w) rhCol1 samples exhibited a significant improvement over the PLGA sample (130% (P<0.0001) and 120% (P=0.015), respectively). The PLGA blend containing 15% (w/w) exhibited no significant difference (P=0.85) compared to the PLGA only sample.

The addition of a neutralising agent to rhCol1 induced a different trend compared to the non- neutralised samples (Figure 3.6). Most noticeable was the major increase in metabolic activity of the cell seeded onto the rhCol1-only coated sample, at 300% of the PLGA-only coated sample (P<0.0001). Increases in the metabolic activity were also observed for all rhCol1-PLGA blend groups, apart from the case of the 15% sample, which did not demonstrate a significant difference compared to the PLGA coated group (P=0.21). The 5 and 10% (w/w) samples exhibited similar results with an increase of 62% over the normalised PLGA value (P=0.041 and P=0.046, respectively). The 20 % (w/w) exhibited the superior results amongst the blended polymer samples with an increase of 93% (P<0.01) over the PLGA.

101

Figure 3.4 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rtCol1 blends. The results were normalised to the PLGA group (100%). The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the PLGA group at the 0.01(**), 0.001(***) and 0.0001 (****) levels (ns indicates no significant difference).

Figure 3.5 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rHCol1 (non-neutralised) blends. The results were normalised to the PLGA group (100%). The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the PLGA group at the 0.05 (*), 0.001(***) and 0.0001 (****) levels (ns indicates no significant difference).

102

Figure 3.6 - Mean metabolic activity of PCAECs seeded onto the different % w/w PLGA-rHCol1 (neutralised) blends. The results were normalised to the PLGA group (100%). The data was expressed as mean± 95% C.I. (n=3). Asterisks indicate significant difference between the test groups and the PLGA group at the 0.05 (*), 0.01(**), 0.001(***) and 0.0001 (****) levels (ns indicates no significant difference).

3.3.4 Chemical Characterisation The ATR-FTIR spectra of the PLGA, rHCol1, rtCol1, and 20% (w/w) collagen-PLGA blend coatings are shown in Figure 3.7. The PLGA sample showed a sharp characteristic peaks at 1746 cm-1 and 1083 cm1, as well as dominant peaks at 1269 cm-1 and 1130 cm-1, within the fingerprint region. Smaller peaks in the 1500-1300 cm-1 range were also observed. Weaker peaks were identified at 1163 cm-1, 952 cm-1, 891 cm-1, 864 cm-1, 842 cm-1 and 735 cm-1. A very sharp but narrow peak was also observed at 684 cm- 1, together with a shoulder peak at 1048 cm-1.

The rHCol1 sample demonstrated a broad peak within the 3500-3000 cm-1 range, with two broad shoulder peaks at 3073 cm-1 and 2979 cm-1. Two major characteristic peaks were observed at 1629 cm- 1 and 1546 cm-1. The fingerprint region had prominent secondary peaks at 1451 cm-1 and 1280 cm-1. Tertiary peaks within this region were observed at 1400 cm-1, 1338 cm-1, 1238 cm-1, 1203 cm-1, 1161 cm-1, 1080 cm-1, 1031 cm-1, 972 cm-1 and 939 cm-1, after which the spectrum continued to increase in absorbance to the edge of the measured spectrum.

The rtCol1 spectrum was very similar to the rHcol1 one, with a common broad peak within the 3500- 3000 cm-1 range. The primary characteristic peaks were observed at 1653 cm-1 and 1552 cm-1. The 1653 cm-1 peak seemed to have an overlapping secondary peak, whereas the rHcol1 sample had a more prominent 1629 cm-1 peak. The only other obvious difference was the occurrence of a sharp peak at

103

1404 cm-1. The rHCol1 showed a much smaller peak at 1400cm-1. The 20% (w/w) rtCol1-PLGA samples did not lack the strong broad peak in the 3500-3000cm-1 as observed in the pure rhCol1 sample, but had a significantly lower level of absorbance. A noticeable but small peak was observed at 2946 cm-1, which was also observed in the pure rtCol1 sample, with a slight shift. The most prominent peaks in the spectrum was seen at 1747 cm-1 and 1162 cm-1, both of which were present in the PLGA sample. All key peaks observed in the PLGA sample were also seen in the rtCol1-PLGA sample. Two weak peaks (1663 cm-1 and 1551 cm-1) matched to the prominent peaks of the pure rHCol1 sample.

The rtCol1-PLGA sample exhibited the prominent broad 3500-3000cm-1 peak to a strong level of absorbance, compared to the pure rtCol1 sample. Similarly to the rHCol-PLGA sample, all PLGA peaks were observed at a similar level of absorbance prominence. Additional peaks at 1644 cm-1 and 1553 cm-1 mimicked the primary peaks in the pure rtCol1 sample. However, an additional peak was observed, bridging the two aforementioned peaks. The point of bridging, at the 1600cm-1 region was not observed in any of the pure sample spectra.

The FTIR spectra of rHCol1-PLGA blend samples loaded with a low (0.16% w/w), medium (0.8% w/w) and high (1.6% w/w) concentration of ferulic acid are illustrated in Figure 3.8. All samples displayed the key peaks of the pure PLGA samples. The sample with the low dose of ferulic acid demonstrated all the major peaks observed in the rhCol1-PLGA sample with a slightly more prominent 3500-3000 cm- 1 broad peak. The medium dose sample lacked the 3500-3000 cm-1 broad range peak and the major rhCol1 peaks in the 1700-1500 cm-1 range. The high dose sample spectrum demonstrated strong rhCol1 peaks relative to the observed PLGA peaks. This exhibited a stronger spectral resemblance to the pure rhCol1 sample. The fingerprint region of the high ferulic acid concertation sample demonstrated a mixed spectral pattern consisting of both PLGA and rhCol1 peaks that were not observed in the rhCol1- PLGA sample. In addition, the absorbance increase towards to edge of the spectrum at the low wavelengths was also observed.

The FTIR spectra of the rtCol1-PLGA blend samples loaded with a low (0.16 % w/w), medium (0.8 % w/w) and high (1.6 % w/w) concentration of ferulic acid are shown in Figure 3.9. All samples exhibited the broad range peak towards the high end of the spectra at 3300cm-1, similar to that of the pure rtCol1 sample. The low and medium ferulic acid concentration spectra exhibited the prominent PLGA peak at 1746cm-1. The high dose sample also exhibited this peak, but at a much lower relative absorbance. The low and medium dosed samples also displayed the key fingerprint peaks seen in the pure PLGA sample. The absorbance level of the characteristic rTCol1 peaks in the 1750-1600 cm-1 range was similar to that of the rtCol1-PLGA sample. The high dosed sample demonstrated similar patterns to that of the rHCol1- PLGA high dosed sample, where strong collagen peaks were observed overshadowing the prominent PLGA 1760-1740 cm-1 peak. The fingerprint region also shared a similar quality, where both collagen

104

and PLGA peaks were present and overlapping. The absorbance rise towards the 400cm-1 spectral edge was observed in all samples, but was more prominent in the sample with the high ferulic acid concentration.

The FTIR spectra of the rHCol1-PLGA samples loaded with a low (7ng/mg), medium (35ng/mg) and high (70ng/mg) concentration of Exendin-4 are illustrated in Figure 3.10. All three samples exhibited spectral similarities to the rHCol1-PLGA sample, although with slightly more collagen prominence amongst these exendin-4 dosed samples. The high dose sample demonstrated a stronger collagen peak absorbance than that of the PLGA within the 1800-1500 cm-1 region. However, the major spectral peak lied in the fingerprint region and comprised of the key 1080cm-1 peak.

The FTIR spectra of the rtCol1-PLGA samples loaded with a low (7ng/mg), medium (35ng/mg) and high (70ng/mg) concentration of Exendin-4 are shown in Figure 3.11. The low and medium dose samples showed very weak rtCol1 peaks when compared to the pure rtCol1-PLGA sample. The higher dose demonstrated slightly more prominent rtCol1-based peaks. The normally occurring shoulder peak observed in the previous samples at 1046 cm-1 was highly pronounced and with a significantly stronger absorbance than the pure PLGA sample.

Figure 3.7 - FTIR spectra of PLGA, rtCol1, rHCol1, and their respective blends (20% (w/w)) with PLGA

105

Figure 3.8 - FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (brown) and high (blue) concentrations of Ferulic Acid

Figure 3.9 - FTIR spectra of rtCol1-PLGA (20% (w/w)) blend coating with low (blue), medium (violet) and high (red) concentrations of Ferulic Acid

106

Figure 3.10- FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (violet) and high (blue) concentrations of Exendin-4

Figure 3.11- FTIR spectra of rHCol1-PLGA (20% (w/w)) blend coating with low (red), medium (brown) and high (blue) concentrations of Exendin-4

107

3.4 Discussion Owing to the favourable drug releasing properties and its FDA and CE approval, PLGA has been regarded as an ideal polymer for drug release applications. The versatility of synthetic polymers gives them a major advantage over natural polymers. PLGA 50:50 has also been reported to have a relatively fast degradation time of 1-2 months, which gives it a preference for drug release applications. However, the hydrophobic nature of PLGA – inherently due to the methyl side groups within the lactide segment – might pose a challenge for cell adhesion. Thus, integrating a natural polymer, such as collagen could potentially enhance the biological cues that could pertain to improved initial cell adhesion.

1,1,1,3,3-hexafluori-2-propanol (HFIP) has been used for electrospinning purposes of collagen and synthetic polymers and is highly acclaimed over other organic solvent such as chloroform or dichloromethane[261]. There have been limited studies on collagen-PLGA blends for thin film solvent casting techniques. A study by Dulnik et al. [262] demonstrated increased cell proliferation on collagen treated with HFIP than with acetic acid. The study also showed improved cell proliferation on HFIP- treated PCL, compared to acetic acid treated PCL. The study suggestd that HFIP provided an easier access to RGD sequences for integrin receptors, which might explain the capacity for collagen treated samples to exhibit improved cellular activity. Previous studies have suggest that the dissolution of collagen within a solvent could cause conformational changes and, therefore, have the ability to retain some solvent within the dried collagen sample [263]. The study has also shown that synthetic polymers have a lower capacity to retain HFIP [263]. This might explain the cytotoxic behaviour observed in the present study with the non-neutralised rhCol1 samples, in addition to the highly acidic solvent that the rhCol1 sample is prepared with. 10mM of hydrochloric acid equates to a pH 2.04, whereas 20mM of acetic acid equates to a pH 3.2. The lack of neutralisation might have caused the retention of residual hydrochloric acid within the blended collagen and, hence, an observable decrease in cell viability with increasing rHCol1 concentrations. The commercially available rtCol1 dissolved in acetic acid did not require neutralisation due to the weak acidity of acetic acid. Neutralization of the pH has been reported to be critical in collagen gelation. Collagen readily precipitates at a basic pH and can be attributed to the protonation of the amines. The purpose of this study, however, was not to create a gel-like substance, but rather a polymer with embedded dried collagen. Upon contact with the medium (generally maintained at pH 7.4), this collagen would have the opportunity to gelate and undergo fibrillogenesis. RTCol1 coating of cell culture well plates and flasks as stated by the manufacturers’ instructions does not require neutralisation in order to achieve improved cell attachment. Upon neutralisation of the rhCol1 samples, an improvement in cell viability and proliferation of the ECs was observed. The expected cytotoxic behaviour of non-neutralised rhCol1 was not observed for the early adhesion assay and could be attributed to the slow release of residual acid over the course of the

108

subsequent 3 days after the initial WST-8 study. The early cell attachment study provided a solid endorsing indication for the potential of in situ endothelialisation. If the seeded ECs were shown to adhere within 30 min after being subjected to a light washing stage and still demonstrate the ability to create a confluent monolayer after 3 days of culture, it might suggest the possibility of in situ endothelialisation. This assessment provides an initial indication for rapid endothelialisation of the stent as soon as it has been implanted and could benefit from possible in situ cell seeding. A fast development of an endothelial layer over the stent struts would alleviate the issue of thrombus formation.

A study by Jose et al. [264] demonstrated that higher collagen concentrations such as 100% (w/w) (1:1) and 54% (w/w) (7:13) caused collagen dissolution and swelling. This suggested that higher concentrations of collagen would ultimately damage the integrity of the PLGA base polymer. The study suggested 25 w/w% as a strategy for bone tissue engineering. Our data demonstrated that the metabolic activity of the polymer blends with higher % (w/w) of collagen showed improved cell proliferation and adhesion. However, a compromise must be made and since sustained degradation of the coating for drug release is the ultimate goal for a drug eluting stent platform, the utilisation of 20 w/w% would provide an adequate increase in cell adhesion and proliferation without greatly destabilising the physical and degradation properties of the PLGA.

FTIR spectroscopy is a useful tool for elucidating structural composition and bonding of polymers. Along this line, it was important to identify whether the coating production method, the addition of collagen or the drug had any detrimental effect on the chemical/structural integrity of the coating. The major peaks within the PLGA spectrum indicated carbonyl stretching of the esters at 1746 cm-1 and C- O stretching of the aliphatic polyesters at 1083cm-1 and 1163cm-1. The smaller peaks located in the fingerprint region of 1450-805, largely corresponded to C-H bending [265,266]. The C-C stretch occurred at 864cm-1 and the small but prominent shoulder peak at 1048cm-1 was also associative of the C-O stretch [267]. Both collagen samples exhibited key functional peaks, particularly corresponding to the characteristic Amide A and B located at 3318cm-1 and 3073cm-1, respectively. The amide I, II and II bands correspond to 1652, 1551 and 1238 cm-1, as reported in literature [268]. These values were present in the rtCol1 spectrum. The rHcol1 demonstrated a slight shift in the amide II peak to 1546cm-1 but the I and III variant exist. Within the rHcol1 sample, the main observable peak lied at the 1629cm-1 value but the amide I peak appeared as a shoulder peak. The rtCol1 sample also exhibited a shoulder peak, mirroring the phenomenon occurring with rHCol1 (Figure 3.12). This result has been reported before by Lazarev et al. [269] indicating the amide I peak for rat collagen was present at 1656 cm-1 with a shoulder peak occurring at 1630cm-1, supporting the findings of the present study. Previous studies have reported that the amide I peak at 1629cm-1 was indicative of a shift corresponding to

109

structural rearrangement of pre-existing proteins and was characteristic of intermolecular hydrogen bonding. It has been suggested that a peak at 1629cm-1 demonstrated high hydrogen bonding and helical structure organisation [270,271]. Bovine tendon collagen has been reported to express amide I absorption at this wavelength and may suggest that superior relative level of biomechanical forces on the bovine tendon compared to the rat tail corresponds to the higher level of structural bonding suggested previously [271]

The differences in amide I:amide II IR absorbance ratios have been reported to be indicative of collagen orientation [272]. The presence of both key characteristic peaks in the PLGA and respective collagen within the blended polymer, indicated a physical distribution of the collagen within the PLGA, as opposed to chemical bonding and, therefore, would not influence the physical properties of the individual constituents. Studies have indicated that the interaction between collagen and PLGA is limited, since slight shifts in the amide I peak only correspond to conformational changes [273]. The spectral data also suggested that the distribution of collagen might not have been homogenous within the PLGA, since certain spectrum (rtCol1-PLGA with 1.6% (w/w) FA, rHCol1-PLGA with 1.6% (w/w) FA and rhCol1-PLGA with 70ng/mg Exendin-4) demonstrated an overwhelming absorbance of the amide bands associated with collagen relative to the PLGA peaks. A 20% (w/w) concentration of collagen translates to a mass ratio of 1:5 of collagen to PLGA. The majority of peaks corresponds with the fact of a higher PLGA composition by displaying observable but small amide bands, indicating presence of the collagen, albeit at a much lower absorbance than the PLGA peaks. This could be attributed to the gelatinous nature of collagen and the concept of blending hydrophobic materials with hydrophilic materials and the inherent nature of immiscibility [274]. There will never be a truly homogenous mix using a low collagen concentration. The sample could have homogenously dispersed regions of collagen. The study did not conclude that the PLGA was not disrupted by the collagen, but rather that the physical interactions between the collagen and the PLGA did not exhibit a fully homogenous behaviour throughout the coating as certain measured regions indicated higher collagen concentrations.

110

1629.18 1652.69

Figure 3.12 - Observation of a double peak within the rHCol1 (blue) and rtCol1 (red) samples associated with the amide I band.

Following neutralisation, the rHCol1 blends outperformed their rtCol1 counterparts and demonstrated the superiority of recombinant human collagen as an improved collagenous product. To the best of the author’s knowledge, recombinant human collagen Type I has never been hybridised with PLGA to create coatings that demonstrate improved endothelial cell proliferation.

This chapter optimised the preparation and development of a polymeric blend utilising a synthetic (PLGA 50:50) and natural (rHCol1 and rtcol1) polymer to act as an optimal coating for drug eluting stents, due to its cell adhesion and proliferation promoting capabilities whilst retaining the chemical properties of both components. This suggested that the PLGA component would be able to provide a suitable platform for drug delivery whilst the collagen component would promote cell adhesion and proliferation due to its inherent biological cues.

111

Chapter IV: Development of In vitro Simulation Systems for Stent Testing

4.1 Introduction This chapter describes the development and commissioning of two functional simulation systems for the in vitro endothelialisation and assessment of coated whole small-calibre vascular stents and flat stent meshes in future studies. The systems were designed to provide near-physiological conditions in order to expose the seeded cells to physiologically-relevant physical simulation. The chapter describes the development of a multi-station pulsatile flow bioreactor for whole stent implantation and testing within a blood vessel in vitro and a culture chamber that was designed to subject cell-seeded planar stent meshes to static and dynamic culture. Both systems were designed so that to accommodate controlled culture medium conditioning, in terms of temperature, pH and dO2, in a commercially- available chemostat unit.

4.2 Culture Chamber System This system was designed to enable the high throughput screening of stent meshes in terms of their ability to support cell attachment and their potential cytotoxicity. The designs and models that are described in this section were made using Autodesk Inventor 2018 and Blender 2.79.

4.2.1 Static Culture System 4.2.1.1 Design Specification & Conceptual Design The static culture system was designed to provide a simplified version of the stent-artery system, so that to controllably assess short-term EC attachment on stents under static conditions. The relevance of this model was to simulate the system of a stented artery during surgery, with the stented arterial wall isolated from the lumen of the artery. This set up would be necessary in order to apply an EC suspension to the stent in situ and allow the cells to attach to the stent for a short period of time before they were exposed to the arterial flow (

Figure 4.1). In order to simplify the design of the culture system, a planar set-up of the stent-artery system was chosen, which would also ease the real-time evaluation of the stent-artery system. In this set up, the cut-opened artery would be placed flat onto a holder, below a flat stent mesh that would be pressed against the flat artery, in order to mimic the in vivo stent-artery interaction, with the aid of a clamp.

The tissue-mesh mating holder was required to hold a large enough arterial tissue segment that would allow for multiple characterisation techniques to be applied on the sample. As such, and since the focus of this project was small-calibre (<6mm) arteries, the minimum width of arterial tissue that could be accommodated was 15mm. Moreover, the tissue-mesh holder was also required allow for sufficient

112

nutrient and oxygen supply to the cultured arterial tissue segment and seeded ECs. Moreover, it was also required to be autoclavable, easy to clean, non-cytotoxic to cells and operate at 37°C and 5% (v/v)

CO2 in culture medium.

The conceptual design of the holder is illustrated in Figure 4.2. It comprised a 316L stainless steel (SS316) base with a perforated SS316L bed, for allowing culture medium access to the abluminal side of the arterial tissue, and two clamps for securing the arterial tissue and flat stent in place. The design also comprised a laser-cut SS316L flat stent mesh that featured the commercially-used stent strut thickness of 150µm (Figure 4.3). The dimensions of the flat stent mesh were 300mm 400mm

0.15mm. × ×

Figure 4.1 – Concept of in situ endothelialisation.

113

Figure 4.2 – Schematics of the planar tissue-stent mating holder of the static system, featuring the individual components.

Figure 4.3 – Schematic of the flat stent mesh

4.2.1.2 Manufactured device The flat stent mesh was manufactured from 150µm-thick SS316L sheets (300-829-68, Goodfellow, Hamburg, Germany), which were laser-cut at the Laser Zentrum Hannover. Initial tests indicated that struts with a width of 90µm displayed poor cut quality. Thus the strut width was increased to 150µm, which demonstrated consistent results (Figure 4.4). Following laser cutting, the stent meshes were electropolished (Poligrat GmbH (Munich, Germany) in order to remove surface artefacts and residual

114

debris. The complete manufactured tissue-stent mating holder featuring the stented arterial segment is illustrated Figure 4.5. Furthermore, and in order to reduce the amount of culture medium required, as well as the amount of staining solution for staining and visualizing the cellular colonization of the stent struts, the holder was placed in a standard culture petri dish (Ø 90 mm) and an autoclavable silicone construct that surrounded the holder was fabricated. The silicone construct, which is illustrated in Figure 4.6, reduced the required fluid volume for culturing and staining the stented artery to 4ml. The silicone construct was fabricated by mixing a two-part silicone solution (rema®Sil Duplicating Silicone A and B, Dentaraum, Ispringen, Germany) at a 1:1 ratio in a falcon tube, and thoroughly stirring, using a glass rod, at RT. The mesh holder was placed in the middle of the petri dish and the silicone solution was poured into the dish and left at RT for 1h before removing the holder.

Figure 4.4 – Differences in cut quality on different strut widths. a) 160µm and b) 90µm. Observation of surface frequent pitting and corrosion in the 90 µm cut sample.

Figure 4.5 - Manufactured tissue holder with sample tissue and mesh

115

Figure 4.6 - Silicone mould for the mesh holder 4.2.3 Dynamic System Flow and shear are critical aspects that influences tissue culture, cell proliferation, ECM production and pharmacokinetics of drug release. Owing to this, a dynamic system was to enable the high throughput screening of stent meshes in terms of their ability to support cell attachment under physiologically- relevant flow.

4.2.3.1 Design Specification & Conceptual Design Similarly to the case of the static system, the dynamic culture system was designed to provide a simplified version of the stent-artery system. In this case, however, the dynamic system was designed to controllably assess short-term EC attachment and long-term EC proliferation and endothelialisation of model flat stents under dynamic-flow conditions. The design specifications of the dynamic system included:  planar set-up of the stent-artery system;  utilisation of the tissue-stent mating holder of the static system or its principle  use of autoclavable materials and easy to handle and clean geometry;  culture medium perfusion capability through the entirety of the tissue;  exposure of cultured stent-arterial tissue system to controllable flow and shear stress;  exposure of cultured stent-arterial tissue system to luminal laminar flow;  integration with chemostat apparatus for culture medium conditioning.  inexpensive manufacturing for multi-station setup.

Based on the design specification, a parallel plate flow system was devised that could provide physiologically-relevant laminar-flow-induced shear stresses onto the lumen of the stented artery. The system allowed the generation of shear stress (τW) that could be calculated according to:

6휇푄 휏푊 = 2 where µ represents the culture medium viscosity, Q푏 theℎ applied culture medium flow, b the width of the chamber and h the height of the culture medium in the chamber. The system utilised silicone

116

spacers and gaskets to provide flow at a fixed height above the stented artery. This allowed for the accurate control of the shear stress. The schematic 3D model of the devised chamber is illustrated in Figure 4.7. The design also featured independently-controlled separate culture medium circulations on the top (luminal side) and bottom (abluminal side) of the stented artery and viewing window for assessing the stented artery in real time.

Figure 4.7 - Schematics of the flow chamber of the dynamic system for the planar tissue-stent system, featuring the top (red) and bottom (blue) flow unit, the tissue-stent mating holder, the stent mesh and the upper (pink) and lower (grey) silicone gaskets.

4.2.3.2 Prototyping Prior to manufacturing the dynamic chamber, it was deemed necessary to create an inexpensive prototype in order to assess the device in terms of potential points of leakage and any additional design flaws. The prototype was created by composite 3D printing in the Institute of Mechatronic Systems (NIFE, Hannover, Germany). The components of the device were printed using a Stratasys Objet 350 Connex3 3D Printer (Stratasys, Minnesota, USA). The window portions of the device were printed with Med610®, a transparent material, whereas the main body was printed with VeroBlack®. The printed components as shown in Figure 4.8, together with the stainless steel base tissue holder, adopted form the static system. Initial evaluation of the design demonstrated several points of poor sealing at the region indicated in Figure 4.9. This led to modifications in the original design by removing the “extended” edges at the point of the inlet and the outlet, thus, reducing the length of the flow path within the device. In addition, the inlet and outlet ports of the main (luminal) circulation were altered to ease manufacturability, since some regions were too thin and pose as potential crack initiation sites. The final design of the flow system is shown in Figure 4.10.

117

Figure 4.8 – 3D printed components (from left to right: Top flow unit, top silicone gasket, bottom silicone gasket and the bottom flow unit) of the flow chamber (top) and the assembled prototype with the mesh holder adopted from the static system (bottom).

Figure 4.9 - Region of leakage observed in the prototype

Figure 4.10 - Final Design of the Flow Chamber System

118

4.2.3.3 Manufactured Device The main body of the dynamic system was manufacture from polysulfone, which features high transparency for sample observation. The fitted barbed inlets and outlets were manufactured from polyether ether ketone (PEEK). Both materials have been reported to be autoclavable and bioinert

[275,276]. The manufactured components and assembly of the dynamic system are illustrated in Figure 4.11.

Figure 4.11 – Assembled and the manufactured Flow Chamber System

4.3 Multi-Station Pulsatile Flow System Although the developed static and dynamic systems could provide useful platforms for a relatively high-throughput in vitro screening of stent coatings and drug compounds, in terms of their effect on stent endothelialisation, using a clinically-relevant stent-living artery interaction, a more near-clinical simulation model would be necessary prior to in vivo animal testing. Specifically, a functional simulation system able to duplicate relevant pulsatile flow and, thus, more realistically simulate the forces imposed on to the stent-artery system, as well as able to accommodate balloon angioplasty surgery, thus, simulating potential damage to the arterial wall, would be a highly valuable tool prior to testing new stents into an expensive and potentially unethical large animal model. Along these lines, a multi-station pulsatile flow bioreactor system that could subject long intact arterial segments to near- physiologically pulsatile flow and pressure was developed and optimised. The general concept of the system was based on the bioreactor system reported by Riches et al. [277].

4.3.1 Description of System Design The multi-station pulsatile flow system comprised several parts, which were designed to enable modularity, ease of handling and autoclavability. The system was composed of 5 main parts, including the tissue stations, base frame, linear actuator, heating system and culture medium conditioning unit (Figure 4.12).

4.3.1.1 Base Frame The base frame was designed to accommodate 4 identical test stations (tissue stations) and was manufactured from aluminium. The base featured adjustable regulator screws at each test chamber position, which were connected to one of two vascular graft holders (one on either side of the graft)

119

inside the tissue chambers and allowed the adjustment between the two holders so that to be able to accommodate different lengths of vascular segments. The base also included a sled drive, which was attached to metallic bellows of the test stations and was driven by an actuator fixed onto the frame. The translation of the sled drive, induced by the actuator, generated a culture medium volume displacement by the compression of the metallic bellows, which in turn generated pulsatile flow in the test stations,

4.3.1.2 Tissue Stations The stations were designed to house small and medium-sized (up to ø15mm) grafts, which could be mounted using sutures or cable ties between two hollow cylindrical holders (Figure 4.13). The spacing between the holders was adjustable by the regulating screw on the base frame. The holders were designed so that they could facilitate the luminal flow and were connected to the compliance chambers via silicone tubing. The tissue stations also featured transparent polycarbonate windows on either their sides, whereas the rest of their parts were manufactured from SS316L stainless steel, expect the compliance chambers, which were manufactured by polysulfone (Quadrant® PSU 1000). Sealing between the different components was achieved by silicone rubber O-rings.

Figure 4.12 - Schematic of the multi-station pulsatile flow bioreactor

120

Figure 4.13 - Schematic of the tissue chamber with the attached compliance chamber

The pulsatile flow through the lumen of the graft attached into the stations was generated by the compression of a flexible metallic bellow, which was connected to the slide drive of the base frame. The cyclic translation of the slide drive generated a cyclic compression of the metallic bellows, which, in turn generated the volume displacement of the culture medium and the pulsatile flow through the lumen of the attached grafts. The luminal flow was then directed to the compliance chambers, which were designed to be secured on the top of the tissue chambers, in order to enhance the compactness of the stations. The compliance chambers, which during operation of the system would be sealed from the atmosphere and contain approximately 150ml of the air, were necessary in order to dampen the peaks in the flow and pressure waveforms that were generated by the flow in the culture medium within the rigid sections of the stations [278]. Following the compliance chamber, the culture medium flow was directed back into the main chamber of the station, and in the space surrounding the graft, through a ports on the top of the main chamber.

4.3.1.3 Linear Actuator A linear actuator was used to drive the sled guide (cyclic translation) and generate the volume displacement and subsequent pulsatile flow in the metallic bellows of the stations. The actuator comprised a stepper motor (STX-115-07 STX AC Stepnet Panel, Copley Controls, Canton, USA) and an electric ball-screw actuator (N2T31V-10-5B-4-MF3-FT1M, IDC Industrial Devices Ltd, Shropshire, UK).

4.3.1.4 Heating System Control heating at the each station was used in order to compensate for the heat loss in the culture medium between the medium conditioning unit and the bioreactor, and maintain the culture medium temperature at 37°. A bespoke Watlow system (Watlow, Minnesota, USA) heating and control system (Figure 4.14) was designed and manufactured for used for the heating control of the culture stations. The system comprised 4 EZ-Zone PM Express temperature control units (one unit per station) that were fixed into a WatConnect control panel, and 4 associated silicone rubber heating pads, which were glued

121

onto base frame, underneath each of the 4 stations. The system also included a PT100 thermocouple that was secured onto one of the 4 stations, for temperature feedback control.

Figure 4.14 – Heating system components. Left: Control unit; right: Heating pads.

4.3.1.5 Chemostat The culture medium was conditioned in a Sartorius chemostat (Biostat B, Sartorius, Germany) prior its perfusion to the bioreactor stations. The chemostat system consisted of the Biostat B control unit and a 5L autoclavable glass chemostat vessel (Figure 4.15). The Biostat B control unit was an all-purpose system that allowed for complete control of all essential culture medium parameters, including partial gas pressures, pH and temperature. Air, O2, CO2 and N2 supplies were connected to the control unit which permitted for finely tuned aeration of the medium. In order to ensure the sterility of the system, the gas mixture from the control unit passed into the chemostat vessel through a filter (Midisart 2000

PTFE filter, Sartorius, Germany) before it was sparged into the culture medium. The temperature, pO2 and pH of the culture medium was monitored using a PT100 thermocouple, a dissolved oxygen (DO) optical sensor (VisiFerm DO 325, Hamilton, USA) and a pH sensor (EasyFerm Plus VP 325, Hamilton, USA), respectively. A heating jacket fitted around the chemostat vessel was used for heating up the culture medium, whereas O2 and N2 supply was used to maintain the pO2 at the required level (21%).

Moreover, CO2 and basic solution supply into the chemostat vessel were used for maintaining the pH value at 7.4. A water cooled condenser with an attached filter was used to allow for the release of gas from the chemostat without the escape of water vapour. Mixing of the medium was achieved by a stirrer located in the middle of the chemostat.

122

Figure 4.15 - Chemostat system including the control unit and the chemostat vessel.

4.3.1.6 Culture Medium Circulation & Bioreactor Setup The bioreactor system was set up as illustrated in Figure 4.16 and Figure 4.17. The modular design, of the bioreactor allowed for the stations to be removed from the main unit and transferred to a laminar flow cabinet for mounting the grafts under sterile conditions. Following sample mounting, the stations were either sealed or isolated from the atmosphere using medical quick connectors or filters, respectively. The stations were then transferred and mounted to the base frame, which was placed on a bench top. The culture medium was perfused from the chemostat vessel to the bioreactor stations using a peristaltic pump (Figure 4.17). The culture medium entered each of the bioreactor stations through an inlet port, which was located distal to the metallic bellow. The volume displacement in the bellows generated pulsatility in the culture medium, which was pushed through the lumen of the mounted vascular graft (Figure 4.18). Subsequently, the medium was directed to the compliance chamber before it re-entered the main chamber, filling up the outside of the vessel. The fluid was then allowed to flow out of the main chamber and back into the chemostat vessel.

123

Figure 4.16 - Diagram illustrating the bioreactor system set up.

Figure 4.17 – Full setup of the bioreactor system with two stations

124

Figure 4.18 - Flow path within the tissue chamber (Blue lines indicate fluid path and green lines indicate pressure sensor positions)

4.3.2 Parameter Control and Data Acquisition

Apart from the culture medium temperature, pH and pO2, which were set and monitored in the chemostat vessel, and the temperature at the station level, which was monitored and controlled by the bespoke Watlow system, described above, the system design also allowed for the controlling of the flow rate, generated by both the linear actuator and peristaltic feeder pump, as well as the controlling and monitoring of the pressure.

The minimum and maximum flow rates of the peristaltic pump (MCP Standard, Ismatec, Wertheim, Germany) with a 3 channel pump head (SB 3V, Ismatec, Wertheim, Germany) 3 channel pump head (SB 3V, Ismatec, Wertheim, Germany) for specific tubing sizes, provided by the manufacturer (Table 4.1), were used to derive approximated pump speed setting based on silicone tubing with 4.8mm ID and 1.6mm wall thickness. The calculated speed settings provided by the manufacturer corresponded well to the experimentally-measured flow rate of the pump (Figure 4.19). However, all further flow calculations were based on the experimentally-obtained pump setting-flow rate profile.

In order to measure the inlet and outlet pressure profiles generated in the bioreactor stations, two types of autoclavable pressure transducers were utilised including a PRESS-S-000 and PRESS-N-025 transducer (PendoTech, Princeton, USA). A 1/4” – 28 UNF hole at the inlet (Figure 4.20) was used to accommodate a threaded barbed fitting onto which a PRESS-N-025 transducer was attached via a 5cm long silicone tubing. The PRESS-S-000 flow through transducer was placed distally to the artery after the culture media left the primary chamber towards the compliance chamber (Figure 4.21).

Table 4.1 - Ismatec data sheet values

Tubing Inner Diameter Min. Value (3.8rpm) [Speed: 1] Max. Value (240rpm) [Speed:999] 4.8mm 2.2ml/min 530ml/min

125

Figure 4.19 - Experimental and factory values (dotted) of the flow rate generated by the peristaltic pump.

Figure 4.20 - Marked 1/4” – 28 UNF hole for PRESS-S-000 pressure transducer fitting (right).

Figure 4.21 – Positioning (black arrow) of the distal pressure transducer (PRESS-N-025; right).

126

The pressure waveforms were recorded using an Analog to Digital Converter (ADC) (PowerLab 4/35, ADInstruments, Dundedin, New Zealand) and associated LabCharts software. The pressure transducers were calibrated using a two-point calibration. Calibration tests were performed using a manometer (C9557 Dry Use Pressure Meter, Comark, Norfolk, UK). Compressed air and a clamp was used to adjust the pressure value. The PRESS-N-025 (Sensor 1 and 4) was connected at the end point of the silicone tubing with a section accommodated by a 3-way connector, where the perpendicular point had an output connection to the manometer. The S-000 (Sensor 2 and 3) was connected in a similar fashion, where a luer seal is used to seal the end of the pressure transducer. The calibration setup is displayed in Figure 4.22. The compressed air or clamping was increased in increments and the generated pressure (in mmHg) was recorded by the manometer, together with the corresponding voltage (in mV) recorded by the LabCharts software for the pressure transducers. A standard curve was created and the zero and slope values were determined and entered into the LabCharts program as shown in Figure 4.23. The transducer data sheet stated a sensor output value of 0.2584mV/psi per volt of excitation, which corresponds to 0.2584 mV per 51.71484 mmHg at an excitation voltage of 10V. These values were verified using the test setup.

Figure 4.22 - Schematic of the pressure transducer calibration set up

127

Figure 4.23 – Standard curve of raw voltage (mV) vs pressure (mmHg) for 2 sensors of each type.

4.3.3 Heating System Calibration In order to calibrate the heating system, the chemostat was filled with distilled water (DW) and maintained at 37°C using the integrated PT-100 temperature sensor and the heating jacket. Once the temperature was stabilised within the chemostat, the connections to the bioreactor chambers were fitted DW was perfused to the bioreactor using the peristaltic pump at a flow rate of 200ml/min. The heating control system of the bioreactor was set at 37°C. Without the use of the heating pads underneath the stations, a drop of 5°C was observed between the chemostat and the bioreactor stations. Since the PT100 sensor used for measuring the temperature was not sterilisable, it was placed within a thin-walled stainless steel gasket that was fitted on the top of a station and protruded into the tissue chamber. Owing to this, it was necessary to measure the real liquid temperature within the station to see whether it corresponded to the temperature measured by the PT100 sensor placed in the gasket. The temperature measured by the PT100 placed in the gasket was 1°C lower than the actual temperature of the fluid temperature in the station. As such, the temperature at the Watlow heating control unit required to be set to 36°C in order to achieve a real fluid temperature of 37°C.

4.3.4 Motor Profile Calculations The bioreactor was designed to be programmable for a range of different pressure of flow waveforms. The main design aim was to simulate physiological arterial pressures. Referring to the physiological pressure profile of the aorta present in Figure 4.24 [279], and as the left ventricle contracts (systole, the pressure increases rapidly, causing the blood to flow rapidly through the aorta. As the cardiac muscles relaxes (diastole), the pressure slowly decreases causing the ventricle to fill up. The slight

128

increase in pressure that occurs after the systolic phase is attributed to the backflow of blood that rebounds off of the semilunar valves and is known as the dicrotic notch.

Figure 4.24 - Representative aortic pressure waveform. Adopted from Klabunde (2016) [279]

In order to simulate such a physiological pressure profile, the programmable Stepnet stepper controller and associated CME2 software were used to define the motion waveform of the linear actuator. The Stepnet controller was designed to work on the principle of using three programmable current settings, including hold, run and boost, which corresponded to stop, constant velocity and acceleration/deceleration, respectively. The required parameters for the actuator distance (µsteps), acceleration (rps2), deceleration (rps2) and constant velocity (rpm) were inputted in the CME2 software. A simple mathematical program was developed to calculate these values for a given flow profile required to be generated by the actuator. Since the stepper controller output was a serial RS- 232, a USB adapter was used to connect the stepper to the PC.

In order to determine the required stroke distance, acceleration, deceleration, and velocity needed from the linear actuator, a series of calculations were required based on initial input parameters and constants. The actuator motion could be simplified to a trapezoidal profile. An example of such a profile is shown in Figure 4.25. This profile determined the base shape of the speed-time graph for the systole/diastole cycle. The percentage of time that the actuator spend accelerating, decelerating and at constant velocity within a single stroke during the simulated systole and diastole were the key input parameters that determined the shape of the trapezoidal profile. The frequency of the linear movement represented the heart rate and governed the total time for a single stroke. The key actuator values provided by the manufacturer and the rest of input values to the motor function equation are presented in Table 4.2.

129

Table 4.2 - Motor profile parameters and critical values

Input Parameters Units Flow rate per station ml/min Frequency Hz (1 Hz = 60 beats/min) Systolic and diastolic timeframe % (totalling to 100%) Systolic/diastolic accel. decel. and constant velocity timeframe % (totalling to 100%) Constants Value Metallic bellow cross-sectional Area 1362.45 mm2 Revolutions per axial inch 5 revolutions No. of steps per 360° 200 Step Angle 1.8°

The distance moved by the motor was calculated according to the following formulation:

3 퐹푙표푤 푣표푙푢푚푒 (푚푚 ) 푚푖푛 퐵푒푙푙표푤 퐴푟푒푎 퐷푖푠푡푎푛푐푒 푚표푣푒푑 (푚푚) = 퐹푟푒푞푢푒푛푐푦 × 60

퐷푖푠푡푎푛푐푒 푚표푣푒푑 푅표푡푎푡푖표푛 표푓 푀표푡표푟 (푑푒푔) = 360° × 5.08

푅표푡푎푡푖표푛 표푓 푚표푡표푟 푚푖푐푟표푠푡푒푝푠 푫풊풔풕풂풏풄풆 풊풏 풎풊풄풓풐풔풕풆풑풔 = × 푆푡푒푝 퐴푛푔푙푒 (1.8°) 푤ℎ표푙푒 푠푡푒푝푠

1 푃푒푟푐푒푛푡푎푔푒 푖푛 푆푦푠푡표푙푒 푇푖푚푒 푠푝푒푛푡 푎푐푐푒푙푒푟푎푡푖푛푔, 푇푎 (푚푠) = × 퐹푟푒푞푢푒푛푐푦 100

푃푒푟푐푒푛푡푎푔푒 푖푛 퐴푐푐푒푙푒푟푎푡푖표푛 × × 1000푚푠 100

퐷푖푠푡푎푛푐푒 푚표푣푒푑 (푚푚) 푀푎푥 푎푥푖푎푙 푣푒푙표푐푖푡푦 = × 1000푚푠 푇푎 푇푑 + 푇푣 + 2 2

(푀푎푥 푎푥푖푎푙 푣푒푙표푐푖푡푦 (푚푚/푠) × 60 × 5 푟표푡푎푡푖표푛푠) 푴풐풕풐풓 푽풆풍풐풄풊풕풚 (풓풑풎) = 25.4푚푚

푀표푡표푟 푉푒푙표푐푖푡푦 (푟푝푚) ퟐ 푴풐풕풐풓 푨풄풄풆풍풆풓풂풕풊풐풏 (풓풑풔 ) = 60 푠 푇푎 (푚푠) 1000 푚푠

130

Figure 4.25 - Exemplary calculated actuator motion profile during simulated systole and diastole.

4.3.5 Pressure Profile Optimisation Optimising the pressure profile was a critical part for maintaining the physiological haemodynamic environment during the simulated systole and diastole. The shape of the actuator motion profile had a direct impact on the simulated pressure waveform. During optimisation of the pressure profile, the compliance chamber was filled with approximately 150ml of air in order to adequately dampen the high frequency peaks in the pressure profile generated by the flow in the rigid compartments of the bioreactor. Since the key input parameters were defined as percentages of one stroke of the linear actuator, a split into forward motion and backward motion mode was required. Each of these modes required inputs in terms of the time spent accelerating, at constant velocity and decelerating. These values were estimated by deconstructing the physiological pressure profile graph into an acceleration time-graph.

In order to experimentally optimise the pressure profile, a silicone tube with a length of 5cm, inner diameter of 0.5mm and outer diameter of 0.8mm was mounted into one of the bioreactor stations to simulate an artery. Water was used as the test fluid due to its similar viscosity to culture medium supplemented with 10% (v/v) FBS (0.88cP vs 0.94cP) [280]. Moreover, the testing was conducted by inducing a flow rate of 100ml/min through the mounted silicone tube. In order to simulate a peripheral resistance (equivalent to the peripheral resistance encountered by the blood in the systemic circulation) into the system, an adjustable clamp was placed distally to the artery, at the point after the media left the tissue chamber and prior to entering the compliance chamber. The application of the peripheral resistance increased the pressure within the artery, as demonstrated in Figure 4.26. Adjustment of the peripheral resistance was required in order to set the pressure within the physiological range (120/80 mmHg – systolic/diastolic). This pressure profile was produced using the input parameters listed in the Figure 4.26.

Due to the smooth curvature of the pressure peak, the parameters were adjusted in order to better represent the physiological profile. Reducing the duration of the constant forward velocity, resulted in

131

a sharper peak. Although the simulation of the physiological dicrotic notch might pose a difficult challenge, the previous pressure profile demonstrated a slight gradient shift during the pressure drop. By increasing the duration of the backward acceleration and decreasing the duration of the constant velocity, it was possible to create the pressure notch. The optimised pressure profile achieved in illustrated in Figure 4.27. The flow rate (100ml/min/chamber) and pulse rate (1Hz) were kept standard. Compared to the physiological pressure waveform, a very strong resemblance was observed (Figure 4.28). This clearly demonstrated the capacity of this pulsatile flow bioreactor to develop and mimic physiological pressure profiles. It could also be used to mimic pathological conditions, such as arrhythmia and high blood pressure.

Figure 4.26 - Pressure waveform at 1 Hz with and without peripheral resistance. Inset: Profile input parameters

Figure 4.27 - Pressure profile using altered input parameters with peripheral resistance.

132

Figure 4.28 - Superimposed physiological aortic pressure (red) and bioreactor pressure profiles.

4.3.6 Optimisation of Tissue Handling and Mounting During the design, development and commissioning, maintenance of sterility and minimum handling was of primary importance. Initially, barbed fittings were placed directly onto the graft holders inside the tissue chamber, and the graft was mounted with sutures in situ inside the chamber. This approach was proven. The mounting method was modified by firstly fitting and securing with sutures luer-ended barbed fittings to the graft, before securing the luer-ended fittings in the tissue chamber (Figure 4.29). This method greatly reduce the handling time. However, this method was also inconsistent, since the sutures were getting loose after a few days of cycling (Figure 4.30). In order to rectify this issue whilst also improving handling, autoclavable biocompatible cable-ties were used to first secure the grafts onto the luer-ended barbed fittings prior to securing the fittings to the tissue chamber (Figure 4.31).

Figure 4.29 - Sutured carotid artery onto luer barbed fittings

133

Figure 4.30 - Loose Suture causing slippage of artery.

Figure 4.31 - Autoclavable zip-ties keeping the tissue in place

4.3.6 Bioreactor System Set Up The complexity and size of the system requires an optimised workflow in order to maintain sterility, reduce unnecessary handling and increase speed of assembly. The workflow of setting up the system is described in Appendix I.

134

Chapter V: Tissue Viability & Stent Endothelialisation under Static & Dynamic Conditions

5.1 Introduction The work presented in this chapter was focused on the assessment of factors affecting tissue viability in order to optimise the sourcing and transportation of arterial tissue for use in the stent testing systems developed. The work presented in this chapter also focused on the validation of the static and dynamic culture systems, in terms of their ability to maintain tissue and cell viability under prolonged cultures. Finally, the chapter describes initial proof of concept studies on stent endothelialisation.

5.2 Materials and Methods 5.2.1 Tissue Procurement and Dissection Porcine aortic arches with the carotid artery still attached were obtained from German Landrace pigs (<9 months old) either from the local abattoir [Source 1] or from BioMedImplant GmbH (Hannover, Germany) following termination of large animal medical device experiments [Source 2]. The aortic arches were dissected in order to obtain the carotid arteries (Figure 5.1). The average length of the carotid artery obtained was 8cm.

Figure 5.1 - Segment of excised porcine aortic arch.

Following excision, the aortas were transported directly to the laboratory in transport medium (described below). The tissues were transferred to DMEM, supplemented with 10% (v/v) FCS and 1%

(v/v) P/S and placed in an incubator at 37°C and 5% (v/v) CO2 for 30min in order to equilibrate the temperature. Subsequently, the tissues were trimmed of excess connective tissue and the carotid arteries were isolated under sterile conditions whilst semi-submerged in DMEM. The carotid arteries were then flushed with pre-warmed DPBS and then tested immediately.

135

5.2.2 Transport Solutions The tissues were transported in the laboratory in either DPBS or Tiprotec (Dr. F. Kohler Chemie Gmbh, Bensheim, Germany), which were supplemented with 0.5mg/ml Gentamycin, 0.2mg/ml Polymixin B and 0.05mg/ml Vancomycin hydrochloride. The solutions were stored at 4°C until required.

5.2.3 Effect of Antibiotic Treatment on Tissue Viability Carotid artery ring samples (n=3) of 3 mm in length obtained from the local slaughter house [Source 1] were flushed with PBS after excision and were incubated in the antibiotic cocktail in DPBS for 1h at

37°C and 5% (v/v) CO2. Samples (n=3) incubated in DMSO under the same conditions and duration served as the negative control. Additional samples (n=3) were stored in DPBS without antibiotics for 1h. All samples were assessed for their metabolic activity using the WST-8 assay. The samples were weighed and the WST-8 metabolic activity results (optical densities) were normalised with respect to the weight of the samples.

5.2.4 Effect of Long-Term Static Culture on Tissue Viability Porcine carotid arteries were obtained from the slaughterhouse and transported in Tiprotec supplemented with antibiotics. Ring segments of 3-5 mm in width were isolated and placed in 24 well plates filled with pre-warmed DMEM with 10% (v/v) FBS and 1% (v/v) P/S and cultured for 7 days at

37°C and 5% (v/v) CO2 in wells within a 24-well plate. Culture media was exchanged every 2 days. Additional ring segments were isolated and directly subjected to the WST-8 metabolic activity assay, representing the Day 0 control. Following assessment, the Day 0 samples were stained with Calcein AM and EThD-1, as described previously. The samples were then washed with PBS and visualised under fluorescence microscopy. The 7 day samples were subject to the same assessment, except for the use of Hoechst 3342 to stain the nuclei instead of EThD-1. In order to further elucidate the effect of long- term culture on tissue viability, the experiment was repeated with triplicate samples assessed at Day 0, Day 2, Day 4 and Day 7 of static culture. In this case, however, the samples were weighed and their WST-8 metabolic activity results (optical densities) were normalised with respect to their weight.

5.2.5 Effect of Cold Storage in Tiprotec and Antibiotics on Tissue Viability Following dissection, a whole carotid artery was placed in Tiprotec, supplemented with the aforementioned antibiotic cocktail, and stored at 4°C for 10h. The samples were segmented to approximately 5mm x 5mm (n=3 per time point) and cultured in different wells of a 24 well plate at

37°C and 5% (v/v) CO2, for periods up to 7 days. Samples (n=3 for each time point)were subjected to WST-8 metabolic activity testing on Day 1, Day 2, Day 3 and Day 7. Live/Dead staining was performed on Day 7 sample after WST-8 incubation to verify validity of the metabolic activity results.

136

5.2.6 Static Culture System Experiments Samples measuring 10mm 20mm were isolated from cut-open carotid arteries obtained from Source

2 were transported in DPBS× (n=1) or Tiprotec (n=1), supplemented with the antibiotic cocktail, were placed on the autoclaved static culture holder, together with an autoclaved uncoated stent mesh that was gently place atop the tissue, and clamped. The holders were placed in a petri-dish (Ø10cm) and filled with DMEM, supplemented with 10% (v/v) FBS and 1% (v/v) P/S until the tissue-stent system was completely covered. The holders with the samples were incubated under standard conditions for 10 days. The culture medium was changed every 2 days. The samples from the arteries that were transported in DPBS were culture for 7 days whilst the samples from the arteries that were transported in Tiprotec were culture for either 3 (n=3), 7 (n=3) or 10 (n=3) days. The viability of the samples was assessed by Calcein AM red and Hoechst nuclei staining.

5.2.7 Dynamic Culture System Experiments Two types of flow experiments were run using the developed dynamic culture system, using either fresh carotid artery samples or PCAEC-seeded silicone membranes. The setup of the apparatus with the samples is illustrated in Figure 5.2. All system components were autoclaved before to mounting the sterile tissues or cell-seeded constructs. The dynamic flow system was run in conjunction with the chemostat, which was used to condition the culture medium to 37°C, pH 7.4 and 21% dO2 prior to perfusion of the medium to the dynamic flow chamber.

Figure 5.2 - Exploded Flow Chamber System Render with endothelialised sheet (left) and excised carotid artery (right).

137

5.2.7.1 Arterial Tissue Cultures The first flow experiment involved porcine carotid arteries from Source 2, which were transported in Tiprotec, supplemented with the antibiotic cocktail. An arterial sample (n=1) measuring 10mm

20mm was laid on the tissue holder bed with the uncoated stent mesh carefully placed over it. A 3-× pronged flow silicone gasket was added on top of the mesh. The bottom gasket was also carefully placed to endure an effective seal. The bottom section of the flow chamber was tightened and set into place. The top part of the chamber was carefully attached. It was critical that the gasket prongs coincide with the fluid slit on the flow chamber. The system was then connected to the chemostat using quick connects. A calculated shear stress of 15 dynes/cm2 was applied onto the samples. The tissue was subject to dynamic conditioning for 2 days followed by 3 days of static culture. Same-size arterial samples (n=3) were maintained in static culture in a 6 well plate for the same duration as the dynamically-stimulated ones. In parallel, the remaining 2 tissue holders were mounted with arterial segments of the same size and stented as an additional control. The flow chamber and tissue holder static mesh controls were unclamped and had the mesh carefully removed as to not cause damage to the tissue. Following culture, the samples were dissected into 4 smaller segments and assessed in terms of their metabolic activity using the WST-8 assay. The samples were weighed and their WST-8 metabolic activity results (optical densities) were normalised with respect to their weight.

5.2.7.2 Endothelialised Silicone Sheet Cultures An alternative approach to the arterial segments, which involved cell-seeded silicone membranes was also investigated for testing stent endothelialisation. Specifically, 3 samples measuring 30mm 40mm were isolated from 200µm-thickness medical grade silicone sheets (SILPURAN, Wacker). The ×samples were then autoclaved and placed in a custom silicone mould (as described in §4.2.1.2) in order to reduce required volumes. The custom silicone mould was prepared in two parts, where the silicone fluid mixture was first poured into a Ø60mm culture dish and allowed to set. Once cured, this layer was removed and a 30mm 40 mm region was marked and cut using a scalpel. The second part involved pouring the silicone× mixture into the culture dish and allowed to cure before pouring another layer (just until the base layer is covered). Without allowing the second layer to set, the cut silicone layer was added to the gently added to the pre-cured layer. This allows fusion of cut segment and the base layer. Subsequently, the samples were coated with fibronectin (354008, Corning). Briefly, 20µl of 200µM fibronectin was added to 1ml sterile distilled water and used to coat each of the silicone sheet samples. The samples were left overnight in order for the fibronectin to effectively bind to the silicone sheets, and the remaining fibronectin solution was aspirated. The sheets were then rinsed with PBS and carefully moved to a culture dish. PCAECs at P5 were seeded onto the fibronectin coated sheets and cultured at 37°C and 5% (v/v) CO2, until confluency. The transparent silicone samples allowed for visualisation under light microscopy to check for confluency. A thick silicone spacer

138

(1mmx30mmx40mm) was autoclaved along with the stent mesh samples and tissue holder. The tissue holder was placed in a custom mould to reduce the culture medium volume. The spacer was placed on the bed of the tissue holder and the endothelialised sheet was carefully placed atop. The uncoated mesh samples were carefully applied to the holder and were clamped down. Sufficient volumes of EGM-2 were added to the tissue holder to endure complete coverage of the endothelialised silicone sheet and mesh. Triplicates were used for visual assessment under calcein staining.

The experiment was repeated with stent meshes that were coated with neutralised rHCol1 at a concentration of 20% (w/w) to PLGA and 1.8% (w/w) of ferulic acid. Prior to the application of the coating, the polymeric blend was shaken on ice at 150 rpm for 15min, before the 3ml of the solution was applied to autoclaved stent mesh samples that were placed in the aforementioned custom silicone moulds in order to reduce the volume of the coating solution required. Following the application of the coating solution, the samples were covered and left in a laminar flow cabinet in order to allow for coating solidification overnight. The coated samples were then washed with 70% ethanol and subjected to UV sterilisation. The coated samples were placed on top of the endothelialised silicone sheet as described for the case of the uncoated meshes. The samples were cultured for 3 days in EGM- 2 with 1% (v/v) P/S before staining them with Calcein AM Green in order to assess cell viability and migration. Samples were coated and stained in triplicates.

5.2.8 Viability Assessment 5.2.8.1 Metabolic Activity Tissue viability was assessed 10% WST-8 solution in Hanks Buffered Saline Solution (HBSS), as previously described was added to each well containing tissue samples. The negative controls for the tissue viability study used tissue samples treated in DMSO for 30 minutes prior to the addition of WST- 8. The plates were then incubated for 2-4 hours and the WST-8 solutions were transferred to a fresh 96-well plate for assessment (with technical triplicates). The plates were loaded into a microplate reader and the absorbance was measured at 450nm and 630nm. For assessment, the background 630nm values were subtracted from the experimental 450nm values. For weight normalisation of the tissues, the optical density values were divided by the weight of the tissue.

5.2.8.1 Fluorescence Imaging Either Calcein AM Green or Calcein AM Red was employed as a staining for cell viability. Ethidium Homodimer-1 (EthD-1) was used to detect dead or dying cells. EthD-1 can only enter cells with a disrupted plasma membrane. Viable cells with a non-compromised membrane do not allow EthD-1 from entering the cell. Calcein and EthD-1 was added to DPBS at a concentration of 1µl/ml each and was mixed thoroughly before being added to the tissue samples. The samples were incubated for 30 minutes under standard conditions and then washed with DPBS three times before fluorescence

139

onto the stent struts, forming a patchy coverage of the struts. At both time points, a strong signal for both the nuclei and calcein staining was also apparent in the arterial tissue itself.

The staining results for the arterial samples, stented with mesh within the tissue holder are presented in Figure 5.12. The struts showed a low endothelialisation with a relatively weak calcein AM red signal. The tissue itself demonstrated strong signs of viability. The corresponding results after 7 days of static culture are presented in Figure 5.13. In this case, there was a significantly higher coverage of the struts with viable ECs. By Day 10, it was apparent that the majority of the struts were covered by viable ECs (Figure 5.14). Moreover, a strong EC viability was observed in the arterial tissue itself.

Figure 5.10 – Nuclei (blue) and viable cells (green) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 2 days. Scale Bars indicate 150 µm.

144

Figure 5.11 - Nuclei (blue) and viable cells (green) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 4 days. Scale Bars indicate 150 µm

145

Figure 5.12 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 3 days. Scale Bars indicate 150 µm

Figure 5.13 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 7 days. Scale Bars indicate 150 µm

146

Figure 5.14 - Nuclei (blue) and viable cells (red) staining of porcine carotid arteries cultured with the uncoated SS stent mesh 10 days. Scale Bars indicate 150 µm

The metabolic activity of the arterial samples subject to dynamic conditioning for 2 days followed by 3 days of static culture are presented in Figure 5.15. The flow (dynamic), tissue holder (TH) (mounted artery segments in TH1 and TH2 under static conditions) and static (tissue cultured in 24-well plate) groups demonstrated far higher metabolic activities than the negative control (WST-8). The data on the left graph was provided to indicate that a clear metabolic activity was observed relative to the blank WST-8 sample. This data was not compared as the weights of the tissue samples were difference and thus was subject to weight normalisation. The normalised data (on the right) showed that the flow group exhibited the lowest metabolic activity when normalised with respect to the weight. Both the cultured samples (including tissue holder) exhibited a significant increase compared to the flow group. The static ring segment shows a significantly higher OD/ng than the statically-cultured tissue holder samples.

The calcein staining results for the PCAEC-seeded fibronectin-coated silicone sheets that were stented with the uncoated mesh and subjected to flow condition are shown in Figure 5.16. The samples demonstrated a poor cell viability and morphology. The large fluoresced artefact in the central image was due to the autofluorescence of the silicone flow gasket. However, small regions of viable PCAECs were observed along the edges of the struts. Figure 5.17 demonstrates the staining of PCAECs seeded

147

on a fibronectin-coated silicone sheet with uncoated stainless steel mesh struts mounted on the tissue holder under static conditions. A confluent layer of cells can be observed with slight indication of viable cells long the edges.

The staining results of PCAEC-seeded fibronectin-coated silicone sheet “stented” with the coated mesh are displayed in Figure 5.18 . After the 3-day culture, a fully confluent viable endothelial layer can be observed. More notably, viable cells were observed on the edges of the mesh strut and some cells were also observed atop the struts as well. More strut edge orientated cells could be observed on the coated mesh than the uncoated mesh.

Figure 5.15 - Metabolic Activity of Porcine Carotid Arteries Cultured under Static and Flow Conditions (Left) and the Weight-Normalised Version (Right). TH1 and TH2 were individually mounted carotid arteries within the tissue holder. This data was combined for the weight normalised version. The static sample indicates carotid segments cultured in a well plate. Data expressed as means ± 95% C.I. (n=3 for Flow and Static and n=6 for Tissue Holder). Asterisks indicate significant difference between the start and end connectors columns at the 0.05 (*) and the 0.0001 (****) levels.

148

Figure 5.16 - Viable cell staining of PCAEC-seeded silicone sheet cultured with the uncoated SS stent for 3 days under flow. Scale bars indicate 200µM

Figure 5.17 – Viable cell staining of PCAEC-seeded silicone sheet cultured with the uncoated SS stent mesh for 3 days under static conditions. Scale bars indicate 200µM.

Figure 5.18 - Viable cell staining of PCAEC-seeded silicone sheet cultured with the SS mesh, coated with PLGA 20% (w/w) rhCol1 and 1.8% (w/w) Ferulic Acid for 3 days. Scale bars indicate 200µM.

149

5.4 Discussion The importance of understanding tissue culture in terms of expected outcomes and typical methods of assessment and visualisation is highlighted in this chapter. The initial study on the metabolic activity did not seem to provide conclusive results as to how the samples were to be compared. It was, therefore, necessary to normalise the optical density values to the weight of the tissue samples. The WST-8 assay provided a linear indication of the cell count and, thus, the assumption that weight normalisation would provide a comparable form of assessing the different tissue groups.

Due to the inconsistency in obtaining tissues from the experimental surgery department, which would be procured under aseptic conditions, it was necessary to find an optimal alternative source of fresh tissue. The key concern in obtaining tissue from slaughterhouses is the lack of sterility. In order to efficiently decontaminate the procured tissues, an optimal disinfection agent needs to be employed. The Cambridge antibiotic solution, which is routinely used in tissue banks, consisting of Gentamycin, Vancomycin and Polymixin B was used in this study to disinfect the procured tissue. The assessment of the effect of the antibiotic cocktail to tissue viability indicated that treatment of the tissue in this antibiotic cocktail for 1h did not produce any observable effects in terms of viability compared to in term of the untreated counterpart.

Due to experimental time constraints it was crucial to be able to preserve unused tissue for future experiments. Tiprotec was used for this purpose both as a transport medium and as a storage medium. Tiprotec is a preservation solution that has been reported to conserve arterial tissue with greater efficacy compared to cold storage solutions such as saline or Custodiol [281]. The key advantage is the functional retention capabilities of the solution as it demonstrated endothelial protection [282]. Calcein staining and metabolic activity assessments have been performed on samples transported in Tiprotec and the antibiotic cocktail and visual assessment suggested a strong fluorescence staining, which was indicative of a healthy endothelium. EC-specific assessments would need to be conducted to verify this claim.

All statically cultured tissue exhibited a similar trend with regards to an initial reduction in cell viability, followed by a reparative status where the tissue regains and develops its metabolic state. This may be attributed to an initial cell shock, where the exposure to the temperature or to the culture media may be an influencing factor. The phase between the transport medium and the culture medium might have been traumatic to the cells. A study by Neutelings et al. [283] reported that mild cold shock (25°C) followed my rewarming activated a pro-apoptotic pathway. Their data also stated that a large proportion of cells died one day after rewarming to 37°C. This supports the trend observed with the reduction in metabolic activity after the first day after cold storage and cold transfer. That study also suggested that rewarming causes an oxidative burst as well as apoptosis.

150

The long term static culture of carotid arteries, after 10h of Tiprotec treatment, also showed viability of the tissue through live/dead staining. Amongst the calcein stained images, an array of elongated cells were observed suggesting possible smooth muscle cells from the tunica media of the artery. There was also a patch of round endothelial-like cells visible, indicating that tissue could be cultured statically for a significant amount of time before the need for transferring it to a dynamic system.

The tissue holder system demonstrated promising results for both the carotid artery and the endothelialised silicone sheet setup. The proof of concept study showed clear visualisation of cell migration and possible proliferation onto the stent struts on the tissue samples. The morphology of the migrated/proliferated cells on the struts resembled that of ECs. The improved cell migration and proliferation of the cells from the tissue sample, compared to the endothelialised sheet, could be potentially attributed to the mechanical properties of the tissue. The compression of the mesh sample onto the tissue caused bulging and thus shortened the distance between the endothelial layer and the top of the struts. The silicone sheets did not have the capacity for a similar level of compressibility and elasticity as native arterial tissue and, therefore, there was a longer distance present between the cells and the top of the struts.

Although, the flow system showed poor cell viability upon calcein staining, this could have been due to the deviation of shear stress calculations based on the nominal parallel plate flow system. Prediction of applied shear stress on an irregular surface may pose some difficulty. Computation fluid dynamics could provide a better approximation of the flow rate required to achieve the appropriate shear stress level.

151

Chapter VI: Conclusion and Future Work

6.1 Conclusion The increasing risk of coronary heart disease throughout the global population has been of critical concern. The increasing population with a requirement for stent-based interventions will ultimately lead to a large population experiencing stent-related pathologies such as restenosis and thrombosis. The main aim of stent manufacturers is to design a stent that eradicates these complications. This can only be achieved through a therapeutic option that promotes rapid endothelialisation whilst simultaneously supressing smooth muscle cell proliferation. This study developed and used systematic assessment techniques to identify potential therapies and coating platforms for use in drug eluting stents. A static and dynamic platform was designed and developed in order to facilitate these coatings and therapies to assess ex vivo performance.

Although initially hypothesised, not all drugs performed as expected for the purpose of dual-functional therapies for combating restenosis and stent thrombosis through in vitro assessments. Within the concentration ranges tested in vitro, magnolol and curcumin demonstrated no optimal effect on the endothelial and smooth muscle cells. Ferulic acid and exendin-4 did provide a level of suitability for this application due to an increase in endothelial cell activity relative to the maintenance of smooth muscle cells. The end goal was to have a drug that promotes endothelialisation faster than the rate of smooth muscle cell proliferation and thus inhibiting the possibility of stent thrombosis associated with current drug eluting stents. It was also observed that at the same concentration, the current commercially available drugs used in drug eluting stents demonstrate that endothelial cells show a higher sensitivity to the anti-proliferatives than the smooth muscle cells, which may contribute to the poor endothelialisation observed in a physiological setting upon insertion of a drug eluting stent. The retention of endothelial functionality and phenotype was observed upon dosage with exendin-4 and ferulic acid.

The assessment of the polymer coatings yielded several insights. It was observed that the rat tail collagen blended with PLGA provided an appropriate surface for cell seeding. The fast adhesion time of 30 minutes upon seeding shows that the hybrid blends of polymer are capable of rapid endothelial cell adhesion and can be incorporated into an aspect of in situ endothelialisation. The tobacco derived recombinant human type I collagen demonstrated a cytotoxic behaviour upon increasing concentrations of the collagen with PLGA. Upon neutralisation of the collagen within the blend, the coatings exhibited improved endothelial cell proliferation. The recombinant human collagen based blends exhibited a higher increase in cell activity as compared to the rat tail collagen and thus indicating improved cell-material behaviour. The addition of drugs to the coatings did not present any IR spectral

152

information of the drug which can be attributed to the low concentrations used. The improved outcome of using recombinant human type I collagen over animal-sourced type I collagen attests that the former type of collagen should be implemented into more biomaterials and coatings.

The commissioning and optimisation of the custom multi-station pulsatile flow bioreactor was achieved during this study. The bioreactor was set up in conjunction with the chemostat for optimising culture media properties whilst maintaining temperature within the tissue chamber. The user handling experience from tissue source to bioreactor was also improved by reducing the handling time of affixing the tissue to the tissue chamber. A thermal control system was employed within the device to compensate for temperature loss throughout the series of tubing from the chemostat to the bioreactor. Most importantly, the pressure waveform generated by the linear actuator based on a series of calculations did achieve near-physiological behaviour.

The newly developed custom static tissue holder and specialised electropolished stainless steel meshes allowed for a simplified planar stented artery in order to improve tissue and cell visualisation upon static culture. The observation of cell migration and proliferation from the tissue and endothelialised sheets onto the struts demonstrated proof of concept of the static system in order to identify endothelialisation rates of different coated mesh samples. The flow system demonstrated cell viability; although poor when using the endothelialised sheet and shows that optimisation of the system needs to be carried out. The preliminary tissue study phase in static indicated possibility of cold storage in tiprotec with the antibiotic solution as cell viability was observed of up to 1 week of culture. The exemplary static assessment of the coated mesh structure affixed atop the endothelialised sheet established cell viability and the early process of cell migration/proliferation to the struts. Once the flow system is optimised, it is hypothesised that the endothelialisation procedure will occur at an improved rate.

This thesis covers the assessment of potential therapeutic drugs dosed on endothelial and smooth muscle cells in order to identify the optimal safe concentration that provides dual functionality on the respective cell types for the purpose of drug eluting stents. Different coating options were optimised and evaluated for endothelial cell interactions and the feasibility of in situ endothelialisation. The optimisation and setup of the multi-station pulsatile flow bioreactor as well as the design and setup of the static and dynamic stent testing platform was achieved. Final porcine tissue studies indicated proof of principle of these systems indicative of feasibility. This project encompasses an assortment of various critical stent-based parameters on an individual basis and developed a system to evaluate these aspects once they are to be assimilated as a final coated mesh structure.

153

6.2 Future Recommendations The work reported in this thesis has paved the way as a development tool for a multitude of future research. Apart from continuing the stent focused research, the bioreactor aspect can dwell into further tissue engineering principles with regards to vascular tissue regeneration and graft synthesis.

Using the systematic assessment of drugs and polymers provides an initial indication towards its potential for use in drug eluting stents. Encompassing these two parameters requires a high degree of optimisation in order to match the drug release/polymer degradation with that of the effective drug concentration. The heterogeneity of the collagen dispersion within the PLGA might be an issue and would require verification. The use of an atomic force microscopy (AFM)-Raman system could provide a visualisation of the dispersion of collagen throughout the coating. The optimisation of the coating procedure is also a required point of study as homogenous thickness of the polymer coating would govern well dispersed drug release.

The static and flow system demonstrated a solid proof of concept and needs validation to ensure that the observable cells on the struts are in fact endothelial cells. The flow system requires further optimisation and would benefit from computation fluid dynamic assessments to identify the optimal flow required to generate physiological shear stresses. One important aspect with regards to stent development is the inhibition of smooth muscle cells and this study focused predominantly on the endothelial aspect. The next stage in the project would be to demonstrate the ability to induce observable smooth muscle cell proliferation over the struts in order to assess the efficacy of SMC inhibition of the coated meshes. It would also be of interest to utilise cobalt chromium mesh samples with thinner struts as this would increase endothelialisation time and provide a more accurate physiological representation. The observable differences of endothelialisation of different coated struts are the primary target for the future of this study.

The bioreactor, with its capability to produce near-physiological pulsatile flow conditions, can be used to house a stented artery in order to observe stent endothelialisation in an improved ex vivo format. The bespoke multistation pulsatile flow bioreactor has the potential to assess the effects of atypical conditions on ex vivo arteries as a biophysical drug model. Conditions like arrhythmia and abnormal blood pressures can be simulated by the highly programmable system. The work presented in this thesis, indicating the possibility of rapid cell adhesion on the polymeric surface, showed that in situ endothelialisation might be a viable option.

154

155

References

[1] Van Buuren F, Dahm JB, Horskotte D. Stent restenosis and thrombosis: etiology, treatment, and outcomes. Minerva Med 2012;103:503–11.

[2] Lee JH, Kim E Do, Jun EJ, Yoo HS, Lee JW. Analysis of trends and prospects regarding stents for human blood vessels. Biomater Res 2018;22:8. doi:10.1186/s40824-018-0114-1.

[3] Van Beusekom HMM, Serruys PW. Drug-Eluting Stent Endothelium: Presence or Dysfunction. JACC Cardiovasc Interv 2010;3:76–7. doi:10.1016/J.JCIN.2009.10.016.

[4] Reejhsinghani R, Lotfi AS. Prevention of stent thrombosis: challenges and solutions. Vasc Health Risk Manag 2015;11:93–106. doi:10.2147/VHRM.S43357.

[5] Saladin K. Human Anatomy. Hum. Anat. 3rd ed., New York: McGraw-Hill Higher Education; 2011, p. 564–71.

[6] OpenStax. Structure and Function of Blood Vessels 2013:1–21. http://cnx.org/contents/58db2cce-b3d9-4904-9049- 80a6cd89264b@4/Structure_and_Function_of_Bloo#tbl-ch21_02 (accessed June 1, 2015).

[7] Secomb TW. Hemodynamics. Compr Physiol 2016;6:975–1003. doi:10.1002/cphy.c150038.

[8] Elad D, Einav S. PHYSICAL AND FLOW PROPERTIES OF BLOOD David Elad and Shmuel Einav. Stand Handb Biomed Eng Des 2004:1–25.

[9] Crystal GJ, Pagel PS. Right Ventricular Perfusion. Anesthesiology 2018;128:202–18. doi:10.1097/ALN.0000000000001891.

[10] Sumpio BE, Riley JT, Dardik A. Cells in focus: endothelial cell. Int J Biochem Cell Biol 2002;34:1508–12.

[11] Liu H-B, Gong Y-F, Yu C-J, Sun Y-Y, Li X-Y, Zhao D, et al. Endothelial progenitor cells in cardiovascular diseases: from biomarker to therapeutic agent. Regen Med Res 2013;1:9. doi:10.1186/2050-490X-1-9.

[12] Walenta K, Bohm M, Schiegl T, Link a., Werner N, Ahlers P, et al. Circulating endothelial progenitor cells and cardiovascular outcomes. N Engl J Med 2005;353:999–1007. doi:353/10/999 [pii] 10.1056/NEJMoa043814.

[13] Ghani U, Shuaib A, Salam A, Nasir A, Shuaib U, Jeerakathil T, et al. Endothelial progenitor cells during cerebrovascular disease. Stroke 2005;36:151–3. doi:10.1161/01.STR.0000149944.15406.16.

[14] Lacolley P, Regnault V, Nicoletti A, Li Z, Michel JB. The vascular smooth muscle cell in arterial pathology: A cell that can take on multiple roles. Cardiovasc Res 2012;95:194–204. doi:10.1093/cvr/cvs135.

[15] Saleh Al-Shehabi T, Iratni R, Eid AH. Anti-atherosclerotic plants which modulate the phenotype of vascular smooth muscle cells. Phytomedicine 2016;23:1068–81. doi:10.1016/J.PHYMED.2015.10.016.

156

[16] Mozaffarian D, Benjamin EJ, Go AS, Arnett DK, Blaha MJ, Cushman M, et al. Heart Disease and Stroke Statistics-2015 Update: A Report From the American Heart Association. Circulation 2014;131:e29–322. doi:10.1161/CIR.0000000000000152.

[17] Frostegård J. Immunity, atherosclerosis and cardiovascular disease. BMC Med 2013;11:117. doi:10.1186/1741-7015-11-117.

[18] Libby P. Current Concepts of the Pathogenesis of the Acute Coronary Syndromes. Circulation 2001;104:365–72. doi:10.1161/01.CIR.104.3.365.

[19] Tesfamariam B. Drug release kinetics from stent device-based delivery systems. J Cardiovasc Pharmacol 2008;51:118–25. doi:10.1097/FJC.0b013e318158540f.

[20] Iakovou I. Thrombosis after stent implantation: how much of a problem is there? Future Cardiol 2008;4:261–7. doi:10.2217/14796678.4.3.261.

[21] Lemesle G, Delhaye C, Bonello L, de Labriolle A, Waksman R, Pichard A. Stent thrombosis in 2008: definition, predictors, prognosis and treatment. Arch Cardiovasc Dis 2008;101:769–77. doi:10.1016/j.acvd.2008.10.006.

[22] Serruys PW, de Jaegere P, Kiemeneij F, Macaya C, Rutsch W, Heyndrickx G, et al. A comparison of balloon-expandable-stent implantation with balloon angioplasty in patients with coronary artery disease. Benestent Study Group. N Engl J Med 1994;331:489–95. doi:10.1056/NEJM199408253310801.

[23] Fischman DL, Leon MB, Baim DS, Schatz R a, Savage MP, Penn I, et al. A randomized comparison of coronary-stent placement and balloon angioplasty in the treatment of coronary artery disease. Stent Restenosis Study Investigators. N Engl J Med 1994;331:496– 501. doi:10.1056/NEJM199408253310802.

[24] Dotter CT, Buschmann RW, McKinney MK, Rösch J. Transluminal expandable nitinol coil stent grafting: preliminary report. Radiology 1983;147:259–60.

[25] Roguin A. Stent: The man and word behind the coronary metal prosthesis. Circ Cardiovasc Interv 2011;4:206–9. doi:10.1161/CIRCINTERVENTIONS.110.960872.

[26] Ako J, Bonneau HN, Honda Y, Fitzgerald PJ. Design Criteria for the Ideal Drug-Eluting Stent. Am J Cardiol 2007;100:3M – 9M. doi:10.1016/j.amjcard.2007.08.016.

[27] Regar E, Sianos G, Serruys PW. Stent development and local drug delivery. [Review] [78 refs]. Br Med Bull 2001;59:227–48.

[28] Wilson Nicholas LM Cruden WM. Advances in coronary stent technology: current expectations and new developments. Res Reports Clin Cardiol 2013:4–85. doi:10.2147/RRCC.S34408.

[29] Whittaker DR, Fillinger MF. The engineering of endovascular stent technology: a review. Vasc Endovascular Surg 2006;40:85–94. doi:10.1177/153857440604000201.

[30] Taylor A, Sigwart U. Endoluminal stenting. London Co Ltd 1996.

157

[31] Yang K, Ren Y, Wan P. High nitrogen nickel-free austenitic stainless steel: A promising coronary stent material. Sci China Technol Sci 2012;55:329–40. doi:10.1007/s11431-011- 4679-3.

[32] Stoeckel D, Bonsignore C, Duda S, Dunitz M. A survey of stent designs. Minim Invasive Ther Allied Technol 2002;11:137–47. doi:10.1080/136457002760273340.

[33] Williams DF. Tissue-biomaterial interactions. J Mater Sci 1987;22:3421–45.

[34] Halwani DO, Anderson PG, Lemons JE, Jordan WD, Anayiotos AS, Brott BC. In-vivo corrosion and local release of metallic ions from vascular stents into surrounding tissue. J Invasive Cardiol 2010;22:528–35.

[35] Köster R, Vieluf D, Kiehn M, Sommerauer M, Kähler J, Baldus S, et al. Nickel and molybdenum contact allergies in patients with coronary in-stent restenosis. Lancet 2000;356:1895–7. doi:10.1016/S1062-1458(01)00267-7.

[36] Saito T, Hokimoto S, Oshima S, Noda K, Kojyo Y, Matsunaga K. Metal allergic reaction in chronic refractory in-stent restenosis. Cardiovasc Revascularization Med 2009;10:17–22. doi:10.1016/j.carrev.2008.01.004.

[37] Poncin P, Proft J. Stent tubing: understanding the desired attributes. In: Poncin P, Proft J, editors., ASM Internationa; 2003, p. 253–9.

[38] Fujiu K, Manabe I, Sasaki M, Inoue M, Iwata H, Hasumi E, et al. Nickel-free stainless steel avoids neointima formation following coronary stent implantation. Sci Technol Adv Mater 2012;13:064218. doi:10.1088/1468-6996/13/6/064218.

[39] Inoue M, Sasaki M, Katada Y, Taguchi T. Quantitative biocompatibility evaluation of nickel-free high-nitrogen stainless steel in vitro/in vivo. J Biomed Mater Res - Part B Appl Biomater 2014;102:68–72. doi:10.1002/jbm.b.32982.

[40] Medtronic. Advantages of Cobalt Alloy for Coronary Stents. Medtronic Vascular; 2003.

[41] Tahir H, Hoekstra a. G, Lorenz E, Lawford P V., Hose DR, Gunn J, et al. Multi-scale simulations of the dynamics of in-stent restenosis: impact of stent deployment and design. Interface Focus 2011;1:365–73. doi:10.1098/rsfs.2010.0024.

[42] Kauffman GB, Mayo I. The Story of Nitinol: The Serendipitous Discovery of the Memory Metal and Its Applications. Chem Educ 1997;2:1–21. doi:10.1007/s00897970111a.

[43] Duerig T, Pelton a, Stöckel D. An overview of nitinol medical applications. Mater Sci Eng A 1999;273-275:149–60. doi:10.1016/S0921-5093(99)00294-4.

[44] Cohen-Inbar O. Nitinol Stenting in Post-Traumatic Pseudo-Aneurysm of Internal Carotid Artery. Open J Mod Neurosurg 2012;02:45–9. doi:10.4236/ojmn.2012.23009.

[45] Stöckel D. The Shape Memory Effect - Phenomenon, Alloys, Applications. EUROflex; 2000.

[46] Barrett RD, Bishara SE, Quinn JK. Biodegradation of orthodontic appliances. Part I. Biodegradation of nickel and chromium in vitro. Am J Orthod Dentofacial Orthop 1993;103:8– 14. doi:10.1016/0889-5406(93)70098-9.

158

[47] Mani G, Feldman MD, Patel D, Agrawal CM. Coronary stents: a materials perspective. Biomaterials 2007;28:1689–710.

[48] Virmani R, Farb A. Pathology of in-stent restenosis. Curr Opin Lipidol 1999;10:499–506.

[49] Martin DM, Boyle FJ. Drug-eluting stents for coronary artery disease: a review. Med Eng Phys 2011;33:148–63. doi:10.1016/j.medengphy.2010.10.009.

[50] Stone GW, Ellis SG, Colombo A, Grube E, Popma JJ, Uchida T, et al. Long-term safety and efficacy of paclitaxel-eluting stents: Final 5-year analysis from the TAXUS clinical trial program. JACC Cardiovasc Interv 2011;4:530–42. doi:10.1016/j.jcin.2011.03.005.

[51] Iqbal J, Gunn J, Serruys PW. Coronary stents: historical development, current status and future directions. Br Med Bull 2013;106:193–211.

[52] Silber S. Cypher versus taxus: Are there differences? J Interv Cardiol 2005;18:441–6. doi:10.1111/j.1540-8183.2005.00084.x.

[53] Shafiq N, Malhotra S, Pandhi P, Behl A, Sharma YP. Cypher versus Taxus: the stent war. Expert Opin Investig Drugs 2006;15:1537–44. doi:10.1517/13543784.15.12.1537.

[54] Chen JL, Gao RL, Yang YJ, Qiao SB, Qin XW, Yao M, et al. Comparison of short- and mid-term outcomes between CYPHER and TAXUS stents in patients with complex lesions of the coronary arteries. Chin Med J (Engl) 2006;119:21–5.

[55] Mehilli J, Dibra A, Kastrati A, Pache J, Dirschinger J, Schömig A. Randomized trial of paclitaxel- and sirolimus-eluting stents in small coronary vessels. Eur Heart J 2006;27:260–6. doi:10.1093/eurheartj/ehi721.

[56] Kastrati A, Mehilli J, von Beckerath N, Dibra A, Hausleiter J, Pache J, et al. Sirolimus-eluting stent or paclitaxel-eluting stent vs balloon angioplasty for prevention of recurrences in patients with coronary in-stent restenosis: a randomized controlled trial. JAMA 2005;293:165–71. doi:10.1016/j.accreview.2005.02.086.

[57] Dibra A, Kastrati A, Mehilli J, Pache J, Schühlen H, von Beckerath N, et al. Paclitaxel-eluting or sirolimus-eluting stents to prevent restenosis in diabetic patients. N Engl J Med 2005;353:663–70. doi:10.1056/NEJMoa044372.

[58] Hossainy S, Prabhu S. A mathematical model for predicting drug release from a biodurable drug-eluting stent coating. J Biomed Mater Res Part A 2008;87A:487–93. doi:10.1002/jbm.a.31787.

[59] Simon C, Palmaz JC, Sprague EA. Influence of topography on endothelialization of stents: clues for new designs. J Long Term Eff Med Implants 2000;10:143–51.

[60] Tsyganov I, Maitz MF, Wieser E. Blood compatibility of titanium-based coatings prepared by metal plasma immersion ion implantation and deposition. Appl Surf Sci 2004;235:156–63. doi:10.1016/j.apsusc.2004.05.134.

[61] Meredith IT, Ormiston J, Whitbourn R, Kay IP, Muller D, Bonan R, et al. First-in-human study of the Endeavor ABT-578-eluting phosphorylcholine-encapsulated stent system in de novo

159

native coronary artery lesions: Endeavor I Trial. EuroIntervention 2005;1:157–64. doi:EIJV1I2A26 [pii].

[62] Singh K, Sun S, Vézina C. Rapamycin (AY-22,989), a new antifungal antibiotic. IV. Mechanism of action. J Antibiot (Tokyo) 1979;32:630–45.

[63] Chang JY, Sehgal SN, Bansbach CC. FK506 and rapamycin: Novel pharmacological probes of the immune response. Trends Pharmacol Sci 1991;12:218–23. doi:10.1016/0165- 6147(91)90555-7.

[64] Gallo R, Padurean a, Jayaraman T, Marx S, Roque M, Adelman S, et al. Inhibition of intimal thickening after balloon angioplasty in porcine coronary arteries by targeting regulators of the cell cycle. Circulation 1999;99:2164–70. doi:10.1161/01.CIR.99.16.2164.

[65] Duda SH, Pusich B, Richter G, Landwehr P, Oliva VL, Tielbeek A, et al. Sirolimus-Eluting Stents for the Treatment of Obstructive Superficial Femoral Artery Disease: Six-Month Results. Circulation 2002;106:1505–9. doi:10.1161/01.CIR.0000029746.10018.36.

[66] Zohlnhöfer D, Klein C a, Richter T, Brandl R, Murr a, Nührenberg T, et al. Gene expression profiling of human stent-induced neointima by cDNA array analysis of microscopic specimens retrieved by helix cutter atherectomy: Detection of FK506-binding protein 12 upregulation. Circulation 2001;103:1396–402. doi:10.1161/01.CIR.103.10.1396.

[67] Suzuki T, Kopia G, Hayashi S, Bailey LR, Llanos G, Wilensky R, et al. Stent-based delivery of sirolimus reduces neointimal formation in a porcine coronary model. Circulation 2001;104:1188–93. doi:10.1161/hc3601.093987.

[68] Hooks MA. Tacrolimus, a new immunosuppressant--a review of the literature. Ann Pharmacother n.d.;28:501–11.

[69] Rowinsky EK, Donehower RC. Paclitaxel (taxol). N Engl J Med 1995;332:1004–14.

[70] Dobesh PP, Stacy Z a, Ansara AJ, Enders JM. Drug-eluting stents: a mechanical and pharmacologic approach to coronary artery disease. Pharmacotherapy 2004;24:1554–77. doi:10.1592/phco.24.16.1554.50955.

[71] Windecker S, Remondino A, Eberli FR, Jüni P, Räber L, Wenaweser P, et al. Sirolimus-eluting and paclitaxel-eluting stents for coronary revascularization. N Engl J Med 2005;353:653–62.

[72] Fattori R, Piva T. Drug-eluting stents in vascular interventions. Lancet 2003;361:247–9.

[73] Nakazawa G, Finn A V., John MC, Kolodgie FD, Virmani R. The Significance of Preclinical Evaluation of Sirolimus-, Paclitaxel-, and Zotarolimus-Eluting Stents. Am J Cardiol 2007;100:S36–44. doi:10.1016/j.amjcard.2007.08.020.

[74] Camenzind E, Wijns W, Mauri L, Kurowski V, Parikh K, Gao R, et al. Stent thrombosis and major clinical events at 3 years after zotarolimus-eluting or sirolimus-eluting coronary stent implantation: A randomised, multicentre, open-label, controlled trial. Lancet 2012;380:1396– 405. doi:10.1016/S0140-6736(12)61336-1.

[75] Pillans P. Immunosuppressants-mechanisms of action and Monitoring. Aust Prescr 2006;29:99–101.

160

[76] Mehilli J, Massberg S. Revisiting the BIOSCIENCE of drug-eluting stent technology. Lancet 2014;384:2086–8.

[77] Natsuaki M, Kozuma K, Morimoto T, Kadota K, Muramatsu T, Nakagawa Y, et al. Biodegradable polymer biolimus-eluting stent versus durable polymer everolimus-eluting stent: A randomized, controlled, noninferiority trial. J Am Coll Cardiol 2013;62:181–90. doi:10.1016/j.jacc.2013.04.045.

[78] Serruys PW, Ruygrok P, Neuzner J, Piek JJ, Seth A, Schofer JJ, et al. A randomised comparison of an everolimus-eluting coronary stent with a paclitaxel-eluting coronary stent:the SPIRIT II trial. EuroIntervention 2006;2:286–94.

[79] Serruys PW, Silber S, Garg S, van Geuns RJ, Richardt G, Buszman PE, et al. Comparison of zotarolimus-eluting and everolimus-eluting coronary stents. N Engl J Med 2010;363:136–46.

[80] Kandzari DE, Leon MB, Popma JJ, Fitzgerald PJ, O’Shaughnessy C, Ball MW, et al. Comparison of Zotarolimus-Eluting and Sirolimus-Eluting Stents in Patients With Native Coronary Artery Disease. J Am Coll Cardiol 2006;48:2440–7. doi:10.1016/j.jacc.2006.08.035.

[81] Park DW, Kim YH, Yun SC, Kang SJ, Lee SW, Lee CW, et al. Comparison of zotarolimus-eluting stents with sirolimus- and paclitaxel-eluting stents for coronary revascularization: The ZEST (Comparison of the Efficacy and Safety of Zotarolimus-Eluting Stent with Sirolimus-Eluting and PacliTaxel-Eluting Stent for Cor. J Am Coll Cardiol 2010;56:1187–95. doi:10.1016/j.jacc.2010.03.086.

[82] Sousa JE, Serruys PW, Costa MA. New frontiers in cardiology: drug-eluting stents: Part I. Circulation 2003;107:2274–9.

[83] Kamath KR, Barry JJ, Miller KM. The Taxus drug-eluting stent: A new paradigm in controlled drug delivery. Adv Drug Deliv Rev 2006;58:412–36. doi:10.1016/j.addr.2006.01.023.

[84] Moses JW, Leon MB, Popma JJ, Fitzgerald PJ, Holmes DR, O’Shaughnessy C, et al. Sirolimus- Eluting Stents versus Standard Stents in Patients with Stenosis in a Native Coronary Artery. N Engl J Med 2003;349:1315–23. doi:10.1056/NEJMoa035071.

[85] Leon MB, Mauri L, Popma JJ, Cutlip DE, Nikolsky E, O’Shaughnessy C, et al. A Randomized Comparison of the Endeavor Zotarolimus-Eluting Stent Versus the TAXUS Paclitaxel-Eluting Stent in De Novo Native Coronary Lesions. 12-Month Outcomes From the ENDEAVOR IV Trial. J Am Coll Cardiol 2010;55:543–54. doi:10.1016/j.jacc.2009.08.067.

[86] Ding N, Pacetti SD, Tang FW, Gada M, Roorda W. XIENCE V (TM) Stent Design and Rationale. J Interv Cardiol 2009;22:S18–27. doi:DOI 10.1111/j.1540-8183.2009.00450.x.

[87] García-García HM, Vaina S, Tsuchida K, Serruys PW. Drug-eluting stents. Arch Cardiol Mex 2006;76:297–319.

[88] Raval A, Parikh J, Engineer C. Mechanism of controlled release kinetics from medical devices. Brazilian J Chem Eng 2010;27:211–25.

[89] Moyer P, Ornato JP, Brady WJ, Davis LL, Ghaemmaghami C a., Gibler WB, et al. Development of Systems of Care for ST-Elevation Myocardial Infarction Patients: The Emergency Medical

161

Services and Emergency Department Perspective. Circulation 2007;116:e43–8. doi:10.1161/CIRCULATIONAHA.107.184047.

[90] Govindarajan T, Shandas R. A survey of Surface Modification Techniques for Next-Generation Shape Memory Polymer Stent Devices. Polymers (Basel) 2014;6:2309–31.

[91] Khang D, Lu J, Yao C, Haberstroh KM, Webster TJ. The role of nanometer and sub-micron surface features on vascular and bone cell adhesion on titanium. Biomaterials 2008;29:970– 83. doi:10.1016/j.biomaterials.2007.11.009.

[92] Samaroo HD, Lu J, Webster TJ. Enhanced endothelial cell density on NiTi surfaces with sub- micron to nanometer roughness. Int J Nanomedicine 2008;3:75–82.

[93] Yeh HI, Lu SK, Tian TY, Hong RC, Lee WH, Tsai CH. Comparison of endothelial cells grown on different stent materials. J Biomed Mater Res - Part A 2006;76:835–41. doi:10.1002/jbm.a.30595.

[94] Pypen CMJM, Plenk H, Ebel MF, Svagera R, Wernisch J. Characterization of microblasted and reactive ion etched surfaces on the commercially pure metals niobium, tantalum and titanium. J Mater Sci Mater Med 1997;8:781–4. doi:10.1023/A:1018568830442.

[95] Chung TW, Liu DZ, Wang SY, Wang SS. Enhancement of the growth of human endothelial cells by surface roughness at nanometer scale. Biomaterials 2003;24:4655–61. doi:10.1016/S0142- 9612(03)00361-2.

[96] Lu J, Rao MP, MacDonald NC, Khang D. Improved endothelial cell adhesion and proliferation on patterned titanium surfaces with …. Acta Biomater 2008;4:192–201.

[97] Nie Z, Kumacheva E. Patterning surfaces with functional polymers. Nat Mater 2008;7:277–90. doi:10.1038/nmat2109.

[98] Bilek MMM, Bax D V, Kondyurin A, Yin Y, Nosworthy NJ, Fisher K, et al. Free radical functionalization of surfaces to prevent adverse responses to biomedical devices. Proc Natl Acad Sci U S A 2011;108:14405–10. doi:10.1073/pnas.1103277108.

[99] Tepe G, Schmehl J, Wendel HP, Schaffner S, Heller S, Gianotti M, et al. Reduced thrombogenicity of nitinol stents--in vitro evaluation of different surface modifications and coatings. Biomaterials 2006;27:643–50. doi:10.1016/j.biomaterials.2005.06.004.

[100] Shabalovskaya S a, Anderegg J, Laab F, Thiel P a, Rondelli G. Surface conditions of Nitinol wires, tubing, and as-cast alloys. The effect of chemical etching, aging in boiling water, and heat treatment. J Biomed Mater Res B Appl Biomater 2003;65:193–203. doi:10.1002/jbm.b.10001.

[101] Lu a, Lu a, Sipehia R, Sipehia R. Antithrombotic and “brinolytic system of human endothelial cells seeded on PTFE: the e!ects of surface modi”cation of PTFE by ammonia plasma treatment and ECM protein coatings. Technology 2001;22:1439–46.

[102] Feng Y, Zhao H, Behl M, Lendlein A, Guo J, Yang D. Grafting of poly(ethylene glycol) monoacrylates on polycarbonateurethane by UV initiated polymerization for improving hemocompatibility. J Mater Sci Mater Med 2013;24:61–70. doi:10.1007/s10856-012-4685-4.

162

[103] Camici GG, Steffel J, Akhmedov a., Schafer N, Baldinger J, Schulz U, et al. Dimethyl Sulfoxide Inhibits Tissue Factor Expression, Thrombus Formation, and Vascular Smooth Muscle Cell Activation: A Potential Treatment Strategy for Drug-Eluting Stents. Circulation 2006;114:1512–21. doi:10.1161/CIRCULATIONAHA.106.638460.

[104] Goddard JM, Hotchkiss JH. Polymer surface modification for the attachment of bioactive compounds. Prog Polym Sci 2007;32:698–725. doi:10.1016/j.progpolymsci.2007.04.002.

[105] Duarte JH. Coronary artery disease: durable-polymer drug-eluting stents might not lead to very late stent thrombosis. Nat Rev Cardiol n.d.;12:3.

[106] Zhu L, Lv YN, Wang LY. Stent thrombosis with biodegradable polymer drug-eluting stents versus durable polymer sirolimus-eluting stents: an update meta-analysis. Cardiology 2015;130:96–105.

[107] ASM International. Poly n-butyl Methacrylate (PBMA)/ Polyethylene Vinyl Acetate (PEVA). In: A. International, ed. Materials and Coatings for Medical Devices: Cardiovascular. Mater. Coatings Med. Devices, Materials Park: ASM International; n.d., p. 432–4.

[108] Tang JZ, Rhodes NP. Release-Controlled Coatings. In: Zhang S, editor. Biol. Biomed. Coatings Handb. Appl., CRC Press; 2012, p. 260–95 SRC – GoogleScholar FG – 0.

[109] Fischell TA, Dishmon D, Elhaddi A, Alsafwah S, Mannem S. The perfect drug eluting stent. Card Interv Today 2009;7:29–36.

[110] Fittipaldi M, Rodriguez LA, Grace LR. The effect of water absorption on the viscoelastic properties of poly(styrene-block-isobutylene-block-styrene) for use in biomedical applications 2015;1664:50003. doi:10.1063/1.4918393.

[111] Lewis a. L, Vick T a., Collias a. CM, Hughes LG, Palmer RR, Leppard SW, et al. Phosphorylcholine-based polymer coatings for stent drug delivery. J Mater Sci Mater Med 2001;12:865–70. doi:10.1023/A:1012803503667.

[112] Gutiérrez-Chico JL, Van Geuns RJ, Regar E, Van Der Giessen WJ, Kelbæk H, Saunamäki K, et al. Tissue coverage of a hydrophilic polymer-coated zotarolimus-eluting stent vs. a fluoropolymer-coated everolimus-eluting stent at 13-month follow-up: An optical coherence tomography substudy from the RESOLUTE All Comers trial. Eur Heart J 2011;32:2454–63. doi:10.1093/eurheartj/ehr182.

[113] Hwang KS. Cathode active material, cathode material having the cathode active material, secondary aqueous lithium-ion battery having the cathode material, and method for preparing the cathode active materia. United States Am Pat No 659032 2010;7 SRC - G.

[114] Hezi-Yamit A, Sullivan C, Wong J, David L, Chen M, Cheng P, et al. Impact of polymer hydrophilicity on biocompatibility: Implication for DES polymer design. J Biomed Mater Res - Part A 2009;90:133–41. doi:10.1002/jbm.a.32057.

[115] Cheng P, Chen M, Udipi K. Biocompatible polymer system for extended drug release., 2009.

[116] Pillai O, Panchagnula R. Polymers in drug delivery. Curr Opin Chem Biol 2001;5:447–51. doi:10.1016/S1367-5931(00)00227-1.

163

[117] Busch R, Strohbach A, Rethfeldt S, Walz S, Busch M, Petersen S, et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater 2014;10:688–700. doi:10.1016/j.actbio.2013.10.015.

[118] Taluja A, Youn YS, Bae YH. Novel approaches in microparticulate PLGA delivery systems encapsulating proteins. J Mater Chem 2007;17:4002. doi:10.1039/b706939a.

[119] Gentile P, Chiono V, Carmagnola I, Hatton P V. An Overview of Poly(lactic-co-glycolic) Acid (PLGA)-Based Biomaterials for Bone Tissue Engineering. OPEN ACCESS Int J Mol Sci 2014;15:15. doi:10.3390/ijms15033640.

[120] Fé Lix Lanao RP, Jonker AM, Wolke JGC, Jansen JA, Van Hest JCM, Leeuwenburgh SCG. Physicochemical Properties and Applications of Poly(lactic-co-glycolic acid) for Use in Bone Regeneration n.d. doi:10.1089/ten.teb.2012.0443.

[121] Wu XS, Wang N. Synthesis, characterization, biodegradation, and drug delivery application of biodegradable lactic/glycolic acid polymers. Part II: biodegradation. J Biomater Sci Polym Ed 2001;12:21–34.

[122] Yoshioka T, Kawazoe N, Tateishi T, Chen G. In vitro evaluation of biodegradation of poly(lactic- co-glycolic acid) sponges. Biomaterials 2008;29:3438–43. doi:10.1016/j.biomaterials.2008.04.011.

[123] Zhu X, Braatz RD. A mechanistic model for drug release in PLGA biodegradable stent coatings coupled with polymer degradation and erosion. J Biomed Mater Res - Part A 2015;103:2269– 79. doi:10.1002/jbm.a.35357.

[124] Qi R, Tian X, Guo R, Luo Y, Shen M, Yu J, et al. Controlled release of doxorubicin from electrospun MWCNTs/PLGA hybrid nanofibers. Chinese J Polym Sci 2016;34:1047–59. doi:10.1007/s10118-016-1827-z.

[125] Göpferich A. Mechanisms of polymer degradation and stabilisation. Polym Degrad Stab 1992;35:95–6. doi:10.1016/0141-3910(92)90141-Q.

[126] Oliviero G, Bergese P, Canavese G, Chiari M, Colombi P, Cretich M, et al. A biofunctional polymeric coating for microcantilever molecular recognition. Anal Chim Acta 2008;630:161–7. doi:10.1016/j.aca.2008.09.069.

[127] Keller S, Blagoi G, Lillemose M, Haefliger D, Boisen A. Processing of thin SU-8 films. J Micromechanics Microengineering 2008;18:125020. doi:10.1088/0960-1317/18/12/125020.

[128] Douroumis D, Onyesom I. Novel Coating Technologies of Drug Eluting Stents Laboratory scale spray drying View project Continuous Cocrystallization for Dissolution Rate Optimization of a Poorly Water-Soluble Drug View project Novel Coating Technologies of Drug Eluting Stents 2011. doi:10.1007/8415_2010_54.

[129] Siemann U. Solvent cast technology – a versatile tool for thin film production. Scatt. Methods Prop. Polym. Mater., Berlin, Heidelberg: Springer Berlin Heidelberg; 2005, p. 1–14. doi:10.1007/b107336.

164

[130] Asahara T, Murohara T, Sullivan A, Silver M, van der Zee R, Li T, et al. Isolation of putative progenitor endothelial cells for angiogenesis. Science 1997;275:964–7.

[131] Werner N, Junk S, Laufs U, Link A, Walenta K, Bohm M, et al. Intravenous transfusion of endothelial progenitor cells reduces neointima formation after vascular injury. Circ Res 2003;93:e17–24. doi:10.1161/01.RES.0000083812.30141.74.

[132] Nakazawa G, Granada JF, Alviar CL, Tellez A, Kaluza GL, Guilhermier MY, et al. Anti-CD34 antibodies immobilized on the surface of sirolimus-eluting stents enhance stent endothelialization. JACC Cardiovasc Interv 2010;3:68–75. doi:10.1016/j.jcin.2009.09.015.

[133] Petersen S, Strohbach A, Busch R, Felix SB, Schmitz K-P, Sternberg K. Site-selective immobilization of anti-CD34 antibodies to poly(l-lactide) for endovascular implant surfaces. J Biomed Mater Res B Appl Biomater 2014;102:345–55. doi:10.1002/jbm.b.33012.

[134] Tan A, Goh D, Farhatnia Y, G N, Lim J, Teoh S-H, et al. An anti-CD34 antibody-functionalized clinical-grade POSS-PCU nanocomposite polymer for cardiovascular stent coating applications: a preliminary assessment of endothelial progenitor cell capture and hemocompatibility. PLoS One 2013;8:e77112. doi:10.1371/journal.pone.0077112.

[135] Vasa M, Fichtlscherer S, Aicher A, Adler K, Urbich C, Martin H, et al. Number and Migratory Activity of Circulating Endothelial Progenitor Cells Inversely Correlate With Risk Factors for Coronary Artery Disease. Circ Res 2001;89:e1–7. doi:10.1161/hh1301.093953.

[136] Peichev M, Naiyer AJ, Pereira D, Zhu Z, Lane WJ, Williams M, et al. Expression of VEGFR-2 and AC133 by circulating human CD34(+) cells identifies a population of functional endothelial precursors. Blood 2000;95:952–8.

[137] Wendel HP, Avci-Adali M, Ziemer G. Endothelial progenitor cell capture stents--hype or hope? Int J Cardiol 2010;145:115–7; author reply 117–8. doi:10.1016/j.ijcard.2009.06.020.

[138] Li J, Li D, Gong F, Jiang S, Yu H, An Y. Anti-CD133 antibody immobilized on the surface of stents enhances endothelialization. Biomed Res Int 2014;2014. doi:10.1155/2014/902782.

[139] Sedaghat A, Sinning JM, Paul K, Kirfel G, Nickenig G, Werner N. First in vitro and in vivo results of an anti-human CD133-antibody coated coronary stent in the porcine model. Clin Res Cardiol 2013;102:413–25. doi:10.1007/s00392-013-0547-4.

[140] Kim TG, Park TG. Biomimicking extracellular matrix: cell adhesive RGD peptide modified electrospun poly(D,L-lactic-co-glycolic acid) nanofiber mesh. Tissue Eng 2006;12:221–33. doi:10.1089/ten.2006.12.ft-31.

[141] Joner M, Cheng Q, Schönhofer-Merl S, Lopez M, Neubauer S, Mas-Moruno C, et al. Polymer- free immobilization of a cyclic RGD peptide on a nitinol stent promotes integrin-dependent endothelial coverage of strut surfaces. J Biomed Mater Res B Appl Biomater 2012;100:637–45. doi:10.1002/jbm.b.31988.

[142] Choi WS, Bae JW, Lim HR, Joung YK, Park J-C, Kwon IK, et al. RGD peptide-immobilized electrospun matrix of polyurethane for enhanced endothelial cell affinity. Biomed Mater 2008;3:044104. doi:10.1088/1748-6041/3/4/044104.

165

[143] Zheng W, Wang Z, Song L, Zhao Q, Zhang J, Li D, et al. Endothelialization and patency of RGD- functionalized vascular grafts in a rabbit carotid artery model. Biomaterials 2012;33:2880–91. doi:10.1016/j.biomaterials.2011.12.047.

[144] Ceylan H, Tekinay AB, Guler MO. Selective adhesion and growth of vascular endothelial cells on bioactive peptide nanofiber functionalized stainless steel surface. Biomaterials 2011;32:8797–805. doi:10.1016/j.biomaterials.2011.08.018.

[145] Hubbell JA, Massia SP, Desai NP, Drumheller PD. Endothelial Cell-Selective Materials for Tissue Engineering in the Vascular Graft Via a New Receptor. Bio/Technology 1991;9:568–72. doi:10.1038/nbt0691-568.

[146] Liu Y, Yang Tan TT, Yuan S, Choong C. Multifunctional P(PEGMA)–REDV conjugated titanium surfaces for improved endothelial cell selectivity and hemocompatibility. J Mater Chem B 2013;1:157–67. doi:10.1039/C2TB00014H.

[147] Ji Y, Wei Y, Liu X, Wang J, Ren K, Ji J. Zwitterionic polycarboxybetaine coating functionalized with REDV peptide to improve selectivity for endothelial cells. J Biomed Mater Res A 2012;100:1387–97. doi:10.1002/jbm.a.34077.

[148] Zhang K, Li J, Wang J, Liu T. Combined REDV Polypeptide and Heparin onto Titanium Surface for the Hemocompatibility and Selectively Endothelialization. J Cell … 2015.

[149] Noel S, Hachem A, Merhi Y, De Crescenzo G. Development of a Polyester Coating Combining Antithrombogenic and Cell Adhesive Properties: Influence of Sequence and Surface Density of Adhesion Peptides. Biomacromolecules 2015;16:1682–94. doi:10.1021/acs.biomac.5b00219.

[150] Graf J, Ogle RC, Robey FA, Sasaki M, Martin GR, Yamada Y, et al. A pentapeptide from the laminin B1 chain mediates cell adhesion and binds to 67000 laminin receptor. Biochemistry 1987;26:6896–900. doi:10.1021/bi00396a004.

[151] Taite LJ, Yang P, Jun H-W, West JL. Nitric oxide-releasing polyurethane-PEG copolymer containing the YIGSR peptide promotes endothelialization with decreased platelet adhesion. J Biomed Mater Res B Appl Biomater 2008;84:108–16. doi:10.1002/jbm.b.30850.

[152] Jun H-W, West JL. Endothelialization of microporous YIGSR/PEG-modified polyurethaneurea. Tissue Eng 2005;11:1133–40. doi:10.1089/ten.2005.11.1133.

[153] De Torre IG, Wolf F, Santos M, Rongen L, Alonso M, Jockenhoevel S, et al. Elastin-like recombinamer-covered stents: Towards a fully biocompatible and non-thrombogenic device for cardiovascular diseases. Acta Biomater 2015;12:146–55. doi:10.1016/j.actbio.2014.10.029.

[154] Kazama K, Okada M, Yamawaki H. A novel adipocytokine, omentin, inhibits platelet-derived growth factor-BB-induced vascular smooth muscle cell migration through antioxidative mechanism. Am J Physiol Heart Circ Physiol 2014;306:H1714–9. doi:10.1152/ajpheart.00048.2014.

[155] Uemura Y, Shibata R, Kanemura N, Ohashi K, Kambara T, Hiramatsu-Ito M, et al. Adipose- derived protein omentin prevents neointimal formation after arterial injury. FASEB J 2015;29:141–51. doi:10.1096/fj.14-258129.

166

[156] Lee JM, Choe W, Kim BK, Seo WW, Lim WH, Kang CK, et al. Comparison of endothelialization and neointimal formation with stents coated with antibodies against CD34 and vascular endothelial-cadherin. Biomaterials 2012;33:8917–27. doi:10.1016/j.biomaterials.2012.08.066.

[157] CHAFFEY N. Alberts, B., Johnson, A., Lewis, J., Raff, M., Roberts, K. and Walter, P. Molecular biology of the cell. 4th edn. Ann Bot 2003;91:401–401. doi:10.1093/aob/mcg023.

[158] Charriere G, Bejot M, Schnitzler L, Ville G, Hartmann DJ. Reactions to a bovine collagen implant. Clinical and immunologic study in 705 patients. J Am Acad Dermatol 1989;21:1203–8.

[159] Parenteau-Bareil R, Gauvin R, Berthod F, Parenteau-Bareil R, Gauvin R, Berthod F. Collagen- Based Biomaterials for Tissue Engineering Applications. Materials (Basel) 2010;3:1863–87. doi:10.3390/ma3031863.

[160] Song J, Rolfe BE, Hayward IP, Campbell GR, Campbell JH. Effects of collagen gel configuration on behavior of vascular smooth muscle cells in vitro: association with vascular morphogenesis. In Vitro Cell Dev Biol Anim 2000;36:600–10. doi:10.1290/1071- 2690(2000)036<0600:EOCGCO>2.0.CO;2.

[161] Khan R, Khan MH. Use of collagen as a biomaterial: An update. J Indian Soc Periodontol 2013;17:539–42. doi:10.4103/0972-124X.118333.

[162] Chen M-C, Liang H-F, Chiu Y-L, Chang Y, Wei H-J, Sung H-W. A novel drug-eluting stent spray- coated with multi-layers of collagen and sirolimus. J Control Release 2005;108:178–89. doi:10.1016/j.jconrel.2005.07.022.

[163] Ramshaw JA, Werkmeister JA, Glattauer V. Collagen-based biomaterials. Biotechnol Genet Eng Rev 1996;13:335–82.

[164] Relou IAM, Damen CA, van der Schaft DWJ, Groenewegen G, Griffioen AW. Effect of culture conditions on endothelial cell growth and responsiveness. Tissue Cell 1998;30:525–30. doi:10.1016/S0040-8166(98)80032-3.

[165] Ige OO, Umoru LE, Aribo S. Natural Products: A Minefield of Biomaterials. ISRN Mater Sci 2012;2012:1–20. doi:10.5402/2012/983062.

[166] Yang C, Hillas PJ, B??ez JA, Nokelainen M, Balan J, Tang J, et al. The Application of Recombinant Human Collagen in Tissue Engineering. BioDrugs 2004;18:103–19. doi:10.2165/00063030-200418020-00004.

[167] Shilo S, Roth S, Dgany O, Mizrahi O, Sheyn D, Pelled G, et al. INSPIRE* RECOMBINANT HUMAN TM TYPE-I COLLAGEN WOUND DRESSING, A NOVEL SCAFFOLD FOR CUTANEOUS WOUND HEALING. n.d.

[168] Fitzgerald GA. Re-engineering Drug Discovery and Development. vol. 17. 2011.

[169] Swanson DR. Medical literature as a potential source of new knowledge. Bull Med Libr Assoc 1990;78:29–37.

[170] Hristovski D, Rindflesch T, Peterlin B. Using literature-based discovery to identify novel therapeutic approaches. Cardiovasc Hematol Agents Med Chem 2013;11:14–24.

167

[171] Wu Y-T, Lin L-C, Tsai T-H. Simultaneous determination of honokiol and magnolol inMagnolia officinalis by liquid chromatography with tandem mass spectrometric detection. Biomed Chromatogr 2006;20:1076–81. doi:10.1002/bmc.644.

[172] Chen J-H, Wu C-C, Hsiao G, Yen M-H. Magnolol induces apoptosis in vascular smooth muscle. Naunyn Schmiedebergs Arch Pharmacol 2003;368:127–33. doi:10.1007/s00210-003-0779-3.

[173] Allahverdian S, Chaabane C, Boukais K, Francis GA, Bochaton-Piallat M-L. Smooth muscle cell fate and plasticity in atherosclerosis. Cardiovasc Res 2018;114:540–50. doi:10.1093/cvr/cvy022.

[174] Wu C-H, Chen C-W, Chen H-C, Chang W-C, Shu M-J, Hung J-S. Elucidating the inhibitory mechanisms of magnolol on rat smooth muscle cell proliferation. J Pharmacol Sci 2005;99:392–9. doi:10.1254/jphs.FP0050473.

[175] Kim H-M, Bae S-J, Kim D-W, Kim B-K, Lee S-B, Lee U-S, et al. Inhibitory role of magnolol on proliferative capacity and matrix metalloproteinase-9 expression in TNF-α-induced vascular smooth muscle cells. Int Immunopharmacol 2007;7:1083–91. doi:10.1016/j.intimp.2007.04.004.

[176] Chen Y-H, Lin S-J, Chen J-W, Ku H-H, Chen Y-L. Magnolol attenuates VCAM-1 expression in vitro in TNF-alpha-treated human aortic endothelial cells and in vivo in the aorta of cholesterol-fed rabbits. Br J Pharmacol 2002;135:37–47. doi:10.1038/sj.bjp.0704458.

[177] Ou HC, Chou FP, Sheu WHH, Hsu SL, Lee WJ. Protective effects of magnolol against oxidized LDL-induced apoptosis in endothelial cells. Arch Toxicol 2007;81:421–32. doi:10.1007/s00204- 006-0172-3.

[178] Hatcher H, Planalp R, Cho J, Torti FM, Torti S V. Curcumin: From ancient medicine to current clinical trials. Cell Mol Life Sci 2008;65:1631–52. doi:10.1007/s00018-008-7452-4.

[179] Chen H-W, Huang H-C. Effect of curcumin on cell cycle progression and apoptosis in vascular smooth muscle cells. Br J Pharmacol 1998;124:1029–40. doi:10.1038/sj.bjp.0701914.

[180] Han J, Pan X-Y, Xu Y, Xiao Y, An Y, Tie L, et al. Curcumin induces autophagy to protect vascular endothelial cell survival from oxidative stress damage 2012. doi:10.4161/auto.19471.

[181] Guo N, Chen F, Zhou J, Fang Y, Li H, Luo Y, et al. Curcumin attenuates rapamycin-induced cell injury of vascular endothelial cells. J Cardiovasc Pharmacol 2015. doi:10.1097/FJC.0000000000000285.

[182] Cheng C-Y, Ho T-Y, Lee E-J, Su S-Y, Tang N-Y, Hsieh C-L. Ferulic Acid Reduces Cerebral Infarct Through Its Antioxidative and Anti-Inflammatory Effects Following Transient Focal Cerebral Ischemia in Rats. Am J Chin Med 2008;36:1105–19. doi:10.1142/S0192415X08006570.

[183] WANG L-Y, TANG Y-P, LIU X, ZHU M, TAO W-W, LI W-X, et al. Effects of ferulic acid on antioxidant activity in Angelicae Sinensis Radix, Chuanxiong Rhizoma, and their combination. Chin J Nat Med 2015;13:401–8. doi:10.1016/S1875-5364(15)30032-7.

[184] Mori T, Koyama N, Guillot-Sestier M-V, Tan J, Town T. Ferulic Acid Is a Nutraceutical β- Secretase Modulator That Improves Behavioral Impairment and Alzheimer-like Pathology in Transgenic Mice. PLoS One 2013;8:e55774. doi:10.1371/journal.pone.0055774.

168

[185] Wang J, Yuan Z, Zhao H, Ju D, Chen Y, Chen X, et al. Ferulic acid promotes endothelial cells proliferation through up-regulating cyclin D1 and VEGF. J Ethnopharmacol 2011;137:992–7. doi:10.1016/j.jep.2011.07.019.

[186] Nicholson SK, Tucker GA, Brameld JM. Effects of dietary polyphenols on gene expression in human vascular endothelial cells. Proc Nutr Soc 2008;67:42–7. doi:10.1017/S0029665108006009.

[187] Wallerath T, Li H, Gödtel-Ambrust U, Schwarz PM, Förstermann U. A blend of polyphenolic compounds explains the stimulatory effect of red wine on human endothelial NO synthase. Nitric Oxide 2005;12:97–104. doi:10.1016/j.niox.2004.12.004.

[188] Kobayashi S, Mimura Y, Notoya K, Kimura I, Kimura M. Antiproliferative effects of the traditional Chinese medicine shimotsu-to, its component cnidium rhizome and derived compounds on primary cultures of mouse aorta smooth muscle cells. Japan J Pharmacol 1992;60:397–401. doi:10.1254/jjp.60.397.

[189] Hou Y-Z, Yang J, Zhao G-R, Yuan Y-J. Ferulic acid inhibits vascular smooth muscle cell proliferation induced by angiotensin II. Eur J Pharmacol 2004;499:85–90. doi:10.1016/j.ejphar.2004.07.107.

[190] Meier JJ. GLP-1 receptor agonists for individualized treatment of type 2 diabetes mellitus. Nat Rev Endocrinol 2012;8:728–42. doi:10.1038/nrendo.2012.140.

[191] Nagayama K, Kyotani Y, Zhao J, Ito S, Ozawa K, Bolstad FA, et al. Exendin-4 Prevents Vascular Smooth Muscle Cell Proliferation and Migration by Angiotensin II via the Inhibition of ERK1/2 and JNK Signaling Pathways 2015. doi:10.1371/journal.pone.0137960.

[192] Hirata Y, Kurobe H, Nishio C, Tanaka K, Fukuda D, Uematsu E, et al. Exendin-4, a glucagon-like peptide-1 receptor agonist, attenuates neointimal hyperplasia after vascular injury. Eur J Pharmacol 2013;699:106–11. doi:10.1016/j.ejphar.2012.11.057.

[193] Nagayama K, Kyotani Y, Zhao J, Ito S, Ozawa K, Bolstad FA, et al. Exendin-4 Prevents Vascular Smooth Muscle Cell Proliferation and Migration by Angiotensin II via the Inhibition of ERK1/2 and JNK Signaling Pathways 2015. doi:10.1371/journal.pone.0137960.

[194] Ding L, Zhang J. Glucagon-like peptide-1 activates endothelial nitric oxide synthase in human umbilical vein endothelial cells. Acta Pharmacol Sin 2012;33:75–81. doi:10.1038/aps.2011.149.

[195] Erdogdu Ö, Nathanson D, Sjöholm Å, Nyström T, Zhang Q. Exendin-4 stimulates proliferation of human coronary artery endothelial cells through eNOS-, PKA- and PI3K/Akt-dependent pathways and requires GLP-1 receptor. Mol Cell Endocrinol 2010;325:26–35. doi:10.1016/j.mce.2010.04.022.

[196] Kawasaki K, Smith RS, Hsieh C-M, Sun J, Chao J, Liao JK. Activation of the Phosphatidylinositol 3-Kinase/Protein Kinase Akt Pathway Mediates Nitric Oxide-Induced Endothelial Cell Migration and Angiogenesis Downloaded from. Mol Cell Biol 2003;23:5726–37. doi:10.1128/MCB.23.16.5726-5737.2003.

169

[197] Abid MR, Guo S, Minami T, Spokes KC, Ueki K, Skurk C, et al. Vascular Endothelial Growth Factor Activates PI3K/Akt/Forkhead Signaling in Endothelial Cells. Arterioscler Thromb Vasc Biol 2004;24:294–300. doi:10.1161/01.ATV.0000110502.10593.06.

[198] Barron V, Lyons E, Stenson-Cox C, Mchugh PE, Pandit A. Bioreactors for Cardiovascular Cell and Tissue Growth: A Review 2003. doi:10.1114/1.1603260.

[199] Watanabe I, Okada S. Effects Of Temperature On Growth Rate Of Cultured Mammalian Cells. J Cell Biol 1967;32:309–23.

[200] Yeatts AB, Fisher JP. Bone tissue engineering bioreactors: Dynamic culture and the influence of shear stress. Bone 2011;48:171–81. doi:10.1016/J.BONE.2010.09.138.

[201] Chen H-C, Hu Y-C. Bioreactors for tissue engineering. Biotechnol Lett 2006;28:1415–23. doi:10.1007/s10529-006-9111-x.

[202] Darling EM, Athanasiou KA. Articular Cartilage Bioreactors and Bioprocesses. Tissue Eng 2003;9:9–26. doi:10.1089/107632703762687492.

[203] Darling EM, Athanasiou KA. Articular Cartilage Bioreactors and Bioprocesses. Tissue Eng 2003;9:9–26. doi:10.1089/107632703762687492.

[204] Partap S, Plunkett NA, O’Brien FJ. Bioreactors in Tissue Engineering. In: Eberli D, editor. Tissue Eng., InTech; 2010, p. 324–36.

[205] Selden C, Fuller B. Role of Bioreactor Technology in Tissue Engineering for Clinical Use and Therapeutic Target Design. Bioengineering 2018;5:32. doi:10.3390/bioengineering5020032.

[206] Barron V, Lyons E, Stenson-Cox C, Mchugh PE, Pandit A. Bioreactors for Cardiovascular Cell and Tissue Growth: A Review 2003. doi:10.1114/1.1603260.

[207] Lyons E, Pandit A. Design of Bioreactors for Cardiovascular Applications. In: Ashammakhi EN, Reis RL, editors. Top. Tissue Eng. 2nd ed., 2004, p. 2–32.

[208] Uno K, Merges CA, Grebe R, Lutty GA, Prow TW. Hyperoxia inhibits several critical aspects of vascular development. Dev Dyn 2007;236:981–90. doi:10.1002/dvdy.21122.

[209] Wootton DM, Ku DN. Fluid Mechanics of Vascular Systems, Diseases, and Thrombosis. Annu Rev Biomed Eng 1999;1:299–329. doi:10.1146/annurev.bioeng.1.1.299.

[210] Dumont K, Yperman J, Verbeken E, Segers P, Meuris B, Vandenberghe S, et al. Design of a new pulsatile bioreactor for tissue engineered aortic heart valve formation. Artif Organs 2002;26:710–4.

[211] Niklason LE, Gao J, Abbott WM, Hirschi KK, Houser S, Marini R, et al. Functional arteries grown in vitro. Science (80- ) 1999;284:489–93. doi:10.1126/science.284.5413.489.

[212] Mundargi R, Venkataraman D, Kumar S, Mogal V, Ortiz R, Loo J, et al. Novel Sensor-Enabled Ex-Vivo Bioreactor—A New Approach towards Physiological Parameters and Porcine Artery Viability. Biomed Res Int 2015.

170

[213] Wang J, Liu L, Wu Y, Maitz MF, Wang Z, Koo Y, et al. Ex vivo blood vessel bioreactor for analysis of the biodegradation of magnesium stent models with and without vessel wall integration n.d. doi:10.1016/j.actbio.2016.12.039.

[214] Yazdani SK, Berry JL. Development of an In Vitro System to Assess Stent-induced Smooth Muscle Cell Proliferation: A Feasibility Study. JVIR 2009;20:101–6. doi:10.1016/j.jvir.2008.09.025.

[215] Punchard MA, O’Cearbhaill ED, MacKle JN, McHugh PE, Smith TJ, Stenson-Cox C, et al. Evaluation of human endothelial cells post stent deployment in a cardiovascular simulator in vitro. Ann Biomed Eng 2009;37:1322–30. doi:10.1007/s10439-009-9701-6.

[216] Ives CL, Eskin SG, McIntire L V. Mechanical effects on endothelial cell morphology: In vitro assessment. Vitr Cell Dev Biol 1986;22:500–7. doi:10.1007/BF02621134.

[217] Kaiser D, Freyberg M-A, Friedl P. Lack of Hemodynamic Forces Triggers Apoptosis in Vascular Endothelial Cells. Biochem Biophys Res Commun 1997;231:586–90. doi:10.1006/BBRC.1997.6146.

[218] Mcfetridge PS, Chaudhuri JB. Design of Vascular Graft Bioreactors. Bioreact. Tissue Eng., Berlin/Heidelberg: Springer-Verlag; n.d., p. 269–83. doi:10.1007/1-4020-3741-4_12.

[219] Usami S, Chen H-H, Zhao Y, Chien S, Skalak~ R. Design and Construction of a Linear Shear Stress Flow Chamber. vol. 21. 1993.

[220] Murayama Y, Suzuki Y, Viñuela F, Kaibara M, Kurotobi K, Iwaki M, et al. Development of a biologically active Guglielmi detachable coil for the treatment of cerebral aneurysms. Part I: in vitro study. AJNR Am J Neuroradiol n.d.;20:1986–91.

[221] Chung B., Robertson A., Peters D. The numerical design of a parallel plate flow chamber for investigation of endothelial cell response to shear stress. Comput Struct 2003;81:535–46. doi:10.1016/S0045-7949(02)00416-9.

[222] Goel SA, Guo L-W, Wang B, Guo S, Roenneburg D, Ananiev GE, et al. High-Throughput Screening Identifies Idarubicin as a Preferential Inhibitor of Smooth Muscle versus Endothelial Cell Proliferation. PLoS One 2014;9:e89349. doi:10.1371/journal.pone.0089349.

[223] Held (BioTek Instruments Inc) P. An Absorbance-based Cytotoxicity Assay using High Absorptivity, Water-soluble Tetrazolium Salts - Cell Quantitation Using WST-8 and the SynergyTM Mx. 2009.

[224] Bratosin D, Mitrofan L, Palii C, Estaquier J, Montreuil J. Novel fluorescence assay using calcein- AM for the determination of human erythrocyte viability and aging. Cytom Part A 2005;66A:78–84. doi:10.1002/cyto.a.20152.

[225] Riss TL, Moravec RA, Niles AL, Duellman S, Benink HA, Worzella TJ, et al. Cell Viability Assays. Eli Lilly & Company and the National Center for Advancing Translational Sciences; 2004.

[226] Moss SC, Lightell DJ, Marx SO, Marks AR, Woods TC. Rapamycin Regulates Endothelial Cell Migration through Regulation of the Cyclin-dependent Kinase Inhibitor p27 Kip1. J Biol Chem 2010;285:11991–7. doi:10.1074/jbc.M109.066621.

171

[227] Lakota K, Mrak-Poljšak K, Rozman B, Kveder T, Tomšič M, Sodin-Semrl S. Serum amyloid A activation of inflammatory and adhesion molecules in human coronary artery and umbilical vein endothelial cells. Eur J Inflamm 2007;5:73–81. doi:10.1177/1721727X0700500203.

[228] De Staercke C, Phillips DJ, Hooper WC. Differential Responses of Human Umbilical and Coronary Artery Endothelial Cells to Apoptosis. Endothelium 2003;10:71–8. doi:10.1080/10623320303369.

[229] Koo H-J, Shin S, Choi JY, Lee K-H, Kim B-T, Choe YS. Introduction of Methyl Groups at C2 and C6 Positions Enhances the Antiangiogenesis Activity of Curcumin. Sci Rep 2015;5:14205. doi:10.1038/srep14205.

[230] Fu Z, Chen X, Guan S, Yan Y, Lin H, Hua Z-C. Curcumin inhibits angiogenesis and improves defective hematopoiesis induced by tumor-derived VEGF in tumor model through modulating VEGF-VEGFR2 signaling pathway. Oncotarget 2015;6:19469–82. doi:10.18632/oncotarget.3625.

[231] Nguyen KT, Shaikh N, Shukla KP, Su S-H, Eberhart RC, Tang L. Molecular responses of vascular smooth muscle cells and phagocytes to curcumin-eluting bioresorbable stent materials. Biomaterials 2004;25:5333–46. doi:10.1016/j.biomaterials.2003.12.033.

[232] Karimian MS, Pirro M, Johnston TP, Majeed M, Sahebkar A. Curcumin and Endothelial Function: Evidence and Mechanisms of Protective Effects. Curr Pharm Des 2017;23. doi:10.2174/1381612823666170222122822.

[233] Martínez-González J, Badimon L. Human and porcine smooth muscle cells share similar proliferation dependence on the mevalonate pathway: implication for in vivo interventions in the porcine model. Eur J Clin Invest 1996;26:1023–32.

[234] Wang J, Yuan Z, Zhao H, Ju D, Chen Y, Chen X, et al. Ferulic acid promotes endothelial cells proliferation through up-regulating cyclin D1 and VEGF. J Ethnopharmacol 2011;137:992–7. doi:10.1016/j.jep.2011.07.019.

[235] Goto H, Nomiyama T, Mita T, Yasunari E, Azuma K, Komiya K, et al. Exendin-4, a glucagon-like peptide-1 receptor agonist, reduces intimal thickening after vascular injury. Biochem Biophys Res Commun 2011;405:79–84. doi:10.1016/j.bbrc.2010.12.131.

[236] Lamalice L, Le Boeuf F, Huot J. Endothelial Cell Migration During Angiogenesis. Circ Res 2007;100:782–94. doi:10.1161/01.RES.0000259593.07661.1e.

[237] Haudenschild CC. Morphology of vascular endothelial cells in culture, Springer, Boston, MA; 1984, p. 129–40. doi:10.1007/978-1-4613-2825-4_13.

[238] Harding A, Cortez-Toledo E, Magner NL, Beegle JR, Coleal-Bergum DP, Hao D, et al. Highly Efficient Differentiation of Endothelial Cells from Pluripotent Stem Cells Requires the MAPK and the PI3K Pathways. Stem Cells 2017;35:909–19. doi:10.1002/stem.2577.

[239] Corada M, Liao F, Lindgren M, Lampugnani MG, Breviario F, Frank R, et al. Monoclonal antibodies directed to different regions of vascular endothelial cadherin extracellular domain affect adhesion and clustering of the protein and modulate endothelial permeability. Blood 2001;97:1679–84.

172

[240] Comhair SAA, Xu W, Mavrakis L, Aldred MA, Asosingh K, Erzurum SC. Human primary lung endothelial cells in culture. Am J Respir Cell Mol Biol 2012;46:723–30. doi:10.1165/rcmb.2011-0416TE.

[241] Chen L, DeWispelaere A, Dastvan F, Osborne WRA, Blechner C, Windhorst S, et al. Smooth Muscle-Alpha Actin Inhibits Vascular Smooth Muscle Cell Proliferation and Migration by Inhibiting Rac1 Activity. PLoS One 2016;11:e0155726. doi:10.1371/journal.pone.0155726.

[242] Rensen SSM, Doevendans PAFM, van Eys GJJM. Regulation and characteristics of vascular smooth muscle cell phenotypic diversity. Neth Heart J 2007;15:100–8.

[243] SMALL J V, GIMONA M. The cytoskeleton of the vertebrate smooth muscle cell. Acta Physiol Scand 1998;164:341–8. doi:10.1046/j.1365-201X.1998.00441.x.

[244] Van Eys GJ, Völler MC, Timmer ED, Wehrens XH, Small J V, Schalken JA, et al. Smoothelin expression characteristics: development of a smooth muscle cell in vitro system and identification of a vascular variant. Cell Struct Funct 1997;22:65–72.

[245] Chettimada S, Joshi SR, Dhagia V, Aiezza A, Lincoln TM, Gupte R, et al. Vascular smooth muscle cell contractile protein expression is increased through protein kinase G-dependent and - independent pathways by glucose-6-phosphate dehydrogenase inhibition and deficiency. Am J Physiol Heart Circ Physiol 2016;311:H904–12. doi:10.1152/ajpheart.00335.2016.

[246] West CML, West DC, Moore JV. Low-density Lipoprotein Uptake and the Sensitivity to Photodynamic Treatment of Cultured Endothelial Cells. Int J Radiat Biol 1991;60:25–8. doi:10.1080/09553009114551451.

[247] Comhair SAA, Xu W, Mavrakis L, Aldred MA, Asosingh K, Erzurum SC. Human Primary Lung Endothelial Cells in Culture n.d. doi:10.1165/rcmb.2011-0416TE.

[248] Coetzee GA, Stein O, Stein Y. Uptake and degradation of low density lipoproteins (LDL) by confluent, contact-inhibited bovine and human endothelial cells exposed to physiological concentrations of LDL. Atherosclerosis 1979;33:425–31. doi:10.1016/0021-9150(79)90035-2.

[249] Shi Q, Aida K, Vandeberg JL, Wang XL. Passage-dependent changes in baboon endothelial cells--relevance to in vitro aging. DNA Cell Biol 2004;23:502–9. doi:10.1089/1044549041562294.

[250] Gaffney J, West D, Arnold F, Sattar A, Kumar S. Differences in the uptake of modified low density lipoproteins by tissue cultured endothelial cells. J Cell Sci 1985;79:317–25.

[251] Sanan DA, Strümpfer AE, van der Westhuyzen DR, Coetzee GA. Native and acetylated low density lipoprotein metabolism in proliferating and quiescent bovine endothelial cells in culture. Eur J Cell Biol 1985;36:81–90.

[252] Voyta JC, Via DP, Butterfield CE, Zetter BR. Identification and isolation of endothelial cells based on their increased uptake of acetylated-low density lipoprotein. J Cell Biol 1984;99:2034–40.

[253] Vicinanza R, Coppotelli G, Malacrino C, Nardo T, Buchetti B, Lenti L, et al. Oxidized low-density lipoproteins impair endothelial function by inhibiting non-genomic action of thyroid hormone-

173

mediated nitric oxide production in human endothelial cells. Thyroid 2013;23:231–8. doi:10.1089/thy.2011.0524.

[254] Henriksen T, Evensen SA, Carlander B. Injury to human endothelial cells in culture induced by low density lipoproteins. Scand J Clin Lab Invest 1979;39:361–8.

[255] Hessler JR, Morel DW, Lewis LJ, Chisolm GM. Lipoprotein oxidation and lipoprotein-induced cytotoxicity. Arteriosclerosis 1983;3:215–22. doi:10.1161/01.ATV.3.3.215.

[256] Sata M, Walsh K. Endothelial cell apoptosis induced by oxidized LDL is associated with the down-regulation of the cellular caspase inhibitor FLIP. J Biol Chem 1998;273:33103–6. doi:10.1074/JBC.273.50.33103.

[257] Xie X, Tan J, Wei D, Lei D, Yin T, Huang J, et al. In vitro and in vivo investigations on the effects of low-density lipoprotein concentration polarization and haemodynamics on atherosclerotic localization in rabbit and zebrafish. J R Soc Interface 2013;10:20121053. doi:10.1098/rsif.2012.1053.

[258] Srinivasan M, Sudheer AR, Menon VP. Ferulic Acid: therapeutic potential through its antioxidant property. J Clin Biochem Nutr 2007;40:92–100. doi:10.3164/jcbn.40.92.

[259] Srinivasan M, Sudheer AR, Menon VP. Serial Review Ferulic Acid: Therapeutic Potential Through Its Antioxidant Property. vol. 40. 2007.

[260] Oh YS, Jun H-S. Effects of Glucagon-Like Peptide-1 on Oxidative Stress and Nrf2 Signaling. Int J Mol Sci 2017;19. doi:10.3390/ijms19010026.

[261] Sadeghi-Avalshahr A, Nokhasteh S, Molavi AM, Khorsand-Ghayeni M, Mahdavi-Shahri M. Synthesis and characterization of collagen/PLGA biodegradable skin scaffold fibers. Regen Biomater 2017;4:309–14. doi:10.1093/rb/rbx026.

[262] Dulnik J, Kołbuk D, Denis P, Sajkiewicz P. The effect of a solvent on cellular response to PCL/gelatin and PCL/collagen electrospun nanofibres. Eur Polym J 2018;104:147–56. doi:10.1016/j.eurpolymj.2018.05.010.

[263] Nam J, Huang Y, Agarwal S, Lannutti J. Materials selection and residual solvent retention in biodegradable electrospun fibers. J Appl Polym Sci 2008;107:1547–54. doi:10.1002/app.27063.

[264] Jose M V., Thomas V, Dean DR, Nyairo E. Fabrication and characterization of aligned nanofibrous PLGA/Collagen blends as bone tissue scaffolds. Polymer (Guildf) 2009;50:3778– 85. doi:10.1016/j.polymer.2009.05.035.

[265] Shibata A, Yada S, Terakawa M. Biodegradability of poly(lactic-co-glycolic acid) after femtosecond laser irradiation. Sci Rep 2016;6:27884. doi:10.1038/srep27884.

[266] Ramos Marques D, Alberto dos Santos L, Ferraz Schopf L, Carlos Soares de Fraga J. Analysis of Poly(Lactic-co-Glycolic Acid)/Poly(Isoprene) Polymeric Blend for Application as Biomaterial n.d. doi:10.4322/polimeros.2013.099.

[267] Blasi P, Schoubben A, Giovagnoli S, Perioli L, Ricci M, Rossi C. Ketoprofen poly(lactide-co- glycolide) physical interaction. AAPS PharmSciTech 2007;8:E78–85. doi:10.1208/pt0802037.

174

[268] Metreveli NO, Jariashvili KK, Namicheishvili LO, Svintradze D V., Chikvaidze EN, Sionkowska A, et al. UV–vis and FT-IR spectra of ultraviolet irradiated collagen in the presence of antioxidant ascorbic acid. Ecotoxicol Environ Saf 2010;73:448–55. doi:10.1016/j.ecoenv.2009.12.005.

[269] Lazarev YA, Grishkovsky BA, Khromova TB. Amide I band of IR spectrum and structure of collagen and related polypeptides. Biopolymers 1985;24:1449–78. doi:10.1002/bip.360240804.

[270] Jackson M, Mantsch HH. The Use and Misuse of FTIR Spectroscopy in the Determination of Protein Structure. Crit Rev Biochem Mol Biol 1995;30:95–120. doi:10.3109/10409239509085140.

[271] De Campos Vidal B, Mello MLS. Collagen type I amide I band infrared spectroscopy. Micron 2011;42:283–9. doi:10.1016/j.micron.2010.09.010.

[272] Bi X, Li G, Doty SB, Camacho NP. A novel method for determination of collagen orientation in cartilage by Fourier transform infrared imaging spectroscopy (FT-IRIS) n.d. doi:10.1016/j.joca.2005.07.008.

[273] Sionkowska A, Skopinska-Wisniewska J, Wisniewski M. Collagen–synthetic polymer interactions in solution and in thin films. J Mol Liq 2009;145:135–8. doi:10.1016/j.molliq.2008.06.005.

[274] Zhang D, Wu X, Chen J, Lin K. The development of collagen based composite scaffolds for bone regeneration. Bioact Mater 2018;3:129–38. doi:10.1016/j.bioactmat.2017.08.004.

[275] Dickinson BL. UDEL® Polysulfone for Medical Applications. J Biomater Appl 1988;3:605–34. doi:10.1177/088532828800300404.

[276] Toth JM. Biocompatibility of Polyaryletheretherketone Polymers. PEEK Biomater Handb 2012:81–92. doi:10.1016/B978-1-4377-4463-7.10007-7.

[277] Riches K, Angelini TG, Mudhar GS, Kaye J, Clark E, Bailey MA, et al. Exploring smooth muscle phenotype and function in a bioreactor model of abdominal aortic aneurysm. J Transl Med 2013;11:208. doi:10.1186/1479-5876-11-208.

[278] Voyvodic PL, Min D, Baker AB. A multichannel dampened flow system for studies on shear stress-mediated mechanotransduction. Lab Chip 2012;12:3322–30. doi:10.1039/c2lc40526a.

[279] Klabunde R. Mean Arterial pressure. Cardiovasc Physiol Concepts 2016. https://www.cvphysiology.com/Blood Pressure/BP006 (accessed September 3, 2018).

[280] Fröhlich E, Bonstingl G, Höfler A, Meindl C, Leitinger G, Pieber TR, et al. Comparison of two in vitro systems to assess cellular effects of nanoparticles-containing aerosols. Toxicol Vitr 2013;27:409–17. doi:10.1016/J.TIV.2012.08.008.

[281] Veres G, Hegedűs P, Barnucz E, Schmidt H, Radovits T, Zöller R, et al. TiProtec preserves endothelial function in a rat model. J Surg Res 2016;200:346–55. doi:10.1016/j.jss.2015.06.062.

175

[282] Veres G, Hegedűs P, Barnucz E, Schmidt H, Radovits T, Zöller R, et al. TiProtec preserves endothelial function in a rat model. J Surg Res 2016;200:346–55. doi:10.1016/j.jss.2015.06.062.

[283] Neutelings T, Lambert CA, Nusgens B V, Colige AC. Effects of mild cold shock (25°C) followed by warming up at 37°C on the cellular stress response. PLoS One 2013;8:e69687. doi:10.1371/journal.pone.0069687.

[284] Quaranta M, Borisov SM, Klimant I. Indicators for optical oxygen sensors. Bioanal Rev 2012;4:115–57. doi:10.1007/s12566-012-0032-y.

176

Appendix I The workflow followed to set up the pulsatile flow duplicator include the following steps:

1. Calibrated the chemostat sensors: dissolved oxygen and pH 2. Cleaned and autoclaved the Univessel and all required tubing. 3. Allowed the Univessel to dry in an oven for 2 days. 4. Allowed for temperature equilibration to room temperature. 5. Connected extra tubing through peristaltic pump (under the safety cabinet) to the chemostat. 6. Added additional supplements (i.e antifoam agent) to the culture media 7. Pumped media from under the safety cabinet directly into the chemostat. 8. Connected the Univessel to the chemostat control unit and conditioned the culture media. 9. Allowed to run for 2-3 days (used as a precaution to check for possible media contaminations) 10. Cleaned and disassembled bioreactor station windows. 11. Placed entire bioreactor station fitted with compliance chamber into a sterilisation bag and autoclaved. Autoclaved additional components in separate sterilisation bags – pressure sensors, zip-ties, barbed fittings. 12. Once dried (placed in oven until so), moved the sterilisation bags to the safety cabinet. 13. Set the computer program for the linear actuator and the pressure sensors. 14. Obtained tissue sample and excised under sterile conditions – removed excess connective tissue. Conducted all excisions with the tissue submerged in pre-warmed DMEM w/ 10% FCS and 1% P/S. 15. Used the zip-ties to connect the tissue to the barbed luer fittings on both ends of the vessel. 16. Carefully connected the luer fittings to their counterpart fitted in the tissue chamber. 17. Closed the bioreactor windows, attached all necessary tubing and pressure sensors and moved the bioreactor system to the frame and locked in place. Connected the quick connect tubes to the corresponding chemostat tubes. 18. Began peristaltic pump and enabled pulsatility – depending on required program. 19. After appropriate dynamic run time, stopped the motor signal by disabling the software. 20. Disconnected the Quick Connects connecting the bioreactor system to the chemostat. 21. Transferred the tissue chambers to the safety cabinet after spraying with 70% ethanol. 22. Unscrewed the top temperature probe port and aspirated the media without touching the tissue. 23. Once completely aspirated, unscrewed one of the windows and carefully removed the luer connects holding the tissue in place. Transferred the luer connects attached to the tissue directly to a petri-dish with pre-warmed DMEM with 10% FBS and 1% P/S.

177

24. Carefully excised the regions connected to the luer fittings and proceeded with tissue assessments as required.

Chemostat Calibration

Prior to any experiments with tissue, it was imperative to calibrate the pH and DO sensors in order to provide highly accurate readings. Any deviation regarding these crucial parameters may cause irreversible damage to the tissue.

Before the system undergoes autoclaving, the Easyferm Plus VP 325 pH sensor was cleaned with ethanol and gently wiped off. The probe was then inserted into a tube containing a pH 7 buffer solution (Calimat, Germany) along with the PT-100 temperature probe and using the Biostat B user interface, waited until the voltage reached equilibrium in order to set up the zero value. The sensor was then rinsed with ethanol and wiped down before submerging the probe into a pH 9 buffer solution (Calimat, Germany) in order to equilibrate for the slope value. Since temperature does influence pH values, the Biostat allows for an automatic system that takes into account the temperature reading during the process of calibration.

The chemostat uses a dissolved oxygen optical sensor which works using a luminescence quenching principle. In the presence of no oxygen, the organic pigment known as luminophore absorbs the excitation light and results in maximum fluorescence. In the presence of oxygen, the O2 molecules quenches the luminophore and stops any emission of fluorescence. The number of O2 molecules colliding with the layer of luminphore results in a measurable fluorescence and thus forms a correlation between dissolved oxygen and voltage [284]. In order to calibrate the Visiferm DO 325 sensor, it was securely fastened into the glass chemostat univessel and ensured sufficient sealing (apart from the gas filter atop the cooling exhaust). To calibrate the zero-value (0% DO), gaseous nitrogen was sparged into the vessel at 300ccm until the voltage reading remained steady for 3 minutes. To calibrate for the slope value, the sparger releasing O2 was ensured to be underwater and the probe was set above the liquid. This was done to account for humidity. The calibration was complete once the voltage value had stabilised.

Chemostat Setup

It is important that all components that have direct contact with the culture media are sterilised efficiently. After thorough cleaning and scrubbing of the internal components, the glass beaker, the tubing, the input and output ports, the sensors, the vessel lid, the exhaust, the sparger and the impeller were all rinsed with distilled water. Around 500ml of water was kept in the beaker to ensure that the pH probe would not dry out during the autoclave procedure. The inlet and outlet tubing, as well as the

178

gas input tube were connected and securely fastened to the top vessel lid using autoclavable zip ties. Quick connectors (Male and Female Connectors – ¼”, Mednet, Munster, Germany) were used in order to connect the inlet and outlet tubing to the desired system. This allowed for a simple and quick connection between the two systems without sacrificing sterility. During the autoclave procedure it was essential for the quick connects to be connected to a dummy (male or female, as per required) connector in order to allow for steam to penetrate through all the tubing for maximum sterilisation efficiency. All air filters and quick connects were wrapped in aluminium foil, ensured that the sealing was fairly loose in order not to overpressure the system (as seen in Figure 0.1). The vessel lid uses knurled knobs intended for hand tightening only and all ports, sensors and fastenings were hand tightened before autoclaving. The system was then sterilised by autoclaving at 121°C and 15 psi for 20 minutes. After the sterilisation run, the system was allowed to cool within the autoclave. All ports were retightened by hand before placing the vessel into a 60°C oven for 24h to ensure the filters were dry before proceeding with media conditioning. The dummy quick connects were removed and aluminium foil was placed securely around the remaining unconnected quick connects in order to minimise contamination. In parallel, a second piece of silicone tubing (1 metre) fixed with a quick connect (and dummy) was also autoclaved. This tubing was used to pump the excess water out of the system and pump media into the system.

Figure 0.1 - Univessel ready for sterilisation

Due to the size of the chemostat vessel, it was not possible to fit under the laminar flow cabinet and therefore had to ensure the output tubing from the vessel was connected to the secondary tubing under the bench. An aspiration pipette connected to the vacuum pump was connected to the output tube of the chemostat vessel and was used to drain the remaining water from the system. Antifoam B

179

was added to the respective culture medium at a concentration of 0.01 % (v/v). The tubing was then fit into the peristaltic pump (BVP Standard 3 channel pump head, Ismatec) and pumped in the respective medium.

The quick connects were then disconnected and rewrapped in fresh aluminium foil under the bench. The chemostat containing the medium was then connected to the Biostat B control unit. The pH and DO sensors were connected along with the gas sparger. The motor was fixed to the primary rotational shaft and the heating blanket was attached. The water inlet and outlet were connected to the exhaust cooler and the water tap was turned on.

It was critical that the culture medium was not contaminated and was pre-conditioned before connecting the chemostat vessel to any bioreactor system. The user interface (Figure 0.2) of the Biostat B control unit allowed for easy parameter control with an automated feedback loop or a manual setting, if required. The settings inputted through the user interface corresponds to Table 0.1

Figure 0.2 - User interface of the chemostat system

Table 0.1 - Input settings for conditioning culture medium using the chemostat

Parameter Set Value Temperature 37°C Stirring 40rpm pH 7.4 pO2 100%

180

Appendix II Curriculum Vitae

Name: Adim De

E-mail: [email protected]

EDUCATION 2014 – 2019 Hannover Medical School PhD Regenerative Medicine N/A 2010-2014 ’14 – ‘19 University of Sheffield MSc Biomaterials and Distinction Regenerative Medicine ’13

2002-2010 Kuwait BEng Biomaterial Sciences Upper Second Class Honours English School and Tissue Engineering ‘10 (2:1) A Levels June 2010 Math A , Chemistry B, Physics B

AS Levels June 2009 Maths A, Chemistry A, Physics B, Design Technology B BEng Dissertation IGCSE June 2008 5 A*’s and 4 A’s

: ‘’Synthesis of Nano-hydroxyapatite for Bone Regeneration using a BEng Extended project continuous hydrothermal technique’’ MSc Dissertation : ‘’Caesium Hydroxide activation of Geopolymer precursors’’ : “Evaluating the integration of Tolfenamic Acid to hydroxyapatite for Doctoral Dissertation In vitro bone regeneration” : “Development of Endothelialisation Methods & Testing

Systems for Vascular Stents” SCHOLARSHIPS BEng Biomaterial Sciences and Tissue Engineering ‘10 PhD Regenerative Medicine ’14 International Engineering Scholarship

ReBirth Scholarship (2014-2015) WORK EXPERIENCE DAAD Scholarship (2015- 2019) Nov 2010, Feb 2011, Nov 2011 – Student Ambassador, Materials Technology and Engineering Open Day ambassador for the Department, The University of department; liason with potential Sheffield students in the Materials department. June 2012- Aug 2012 – Summer Research Experience, Kroto research Institute Summer research position working on producing bilayered porous polycaprolactone scaffold.

181

Academic Contributions

in vitro 09/17 – De A, Pflaum M, Wirth D, Sedding D, Wiegmann B, Haverich A, Korossis S. Systematic evaluation of therapeutic agents for drug eluting stents. Poster presentation at European Society of Biomaterials in Athens, Greece.

12/17 - De A, Pflaum M, Wirth D, Sedding D, Wiegmann B, Haverich A, Korossis S. Metabolic assessment of therapies for the development of drug eluting stents. Poster presentation at TERMISResearch – AM Interests 2017 in Charlotte, NC, USA.

The biomaterials based interest began with bone tissue regeneration, but decided to broaden my horizon and undertake a PhD topic in the cardiovascular field which utilises the skills and knowledge I have obtained at The University of Sheffield. During my PhD, there was a large focus on tissue engineering with the use of various bioreactors and thus, have developed a keen interest in design, development and commissioning of medical devices with respect to tissue engineering. I would like to use my skills in the biotech field – with a strong focus on Research and Development and New TechnicalProductSkills, Development. Awards and Achievements

Microsoft Word, PowerPoint, Excel and Access. Autodesk Inventor, Autodesk CFDFlex, Blender 3D Modelling, Graphpad Prism, OMNIC, Substance Software Fourier Transform Infrared Imaging, X-Ray Diffraction, Raman Spectroscopy, Microwave Synthesising, Scanning Electron Microscopy, cell/tissue culture and biological characterisation techniques. Bioreactor Device Design School Prefect, Student council President/Vice President Gold Duke of Edinburgh Award (Expedition to Czech Republic and Jordan) Model United Nations, Experienced in Public speaking and logical reasoning (conferences in Italy, Kuwait and Singapore) Went for charity summit to Mount Kilimanjaro in August 2012. Raised over £3400 for a charity called DigDeep. Organising committee for the University’s Engineering Department Ball Committee member for the Materials Society (MATSOC) at The University of Sheffield as Social Sec. Entails the organisation of large social gatherings. Accredited by the RBS for entrepreneurship and performance in MATSOC University of Sheffield Squash Team player and committee member, Hannover Boastars Squash Player Languages: English, Hindi, Bengali Interests

   Public Speaking  Bouldering/Rock Climbing  Event Management and Organisation Avid traveller and outdoor fanatic 3D Modelling and DIY projects 182

Author’s Statement of Contribution

I hereby declare that the research presented in this thesis is of original work that I carried out during my candidature at the Hannover Medical School.

The studies were designed by myself and Professor Korossis. All experiments reported in the thesis, unless otherwise stated were executed by myself. All data were collated and analysed by myself. Processing of the raw flow cytometry data was carried out by Dr. Pflaum and analysed by myself. The development of the heating system for the multi-station pulsatile flow bioreactor was carried out by Dr. Kalozoumis. Contributions by Dr. Pflaum in experimental design and concept was in the advisory capacity.

183

Declaration

Herewith, I confirm that I have written the present PhD thesis myself and independently, in compliance with ‘the policy of Hannover Medical School on the safeguarding of good scientific practice and procedural rules for dealing with scientific misconduct’, and that I have not submitted it or parts thereof at any other university worldwide.

Herewith, I agree that MHH can check my thesis by plagiarism detection software as well as randomly check the primary data. I am aware that in case of suspicion, ombudsman proceedings according to § 9 of MHH’s ‘Policy of Hannover Medical School on the safeguarding of good scientific practice and procedural rules for dealing with scientific misconduct’ will be initiated. During such proceedings, the PhD process is paused.

Hannover, 08 February 2019

______

184