EFFECT OF DECELLULARISATION METHODS ON

METHACRYLOYL‐SUBSTITUTED PLACENTAL ECM

HYDROGELS

David Pershouse Bachelor Applied Science/ Bachelor of Mathematics

Submitted in fulfilment of the requirements for the degree of Master of Applied Science (Research)

School of Mechanical, Medical and Process Engineering

Science and Engineering Faculty Queensland University of Technology 2020

Keywords

3D cell culture

Placenta

Cell instructivity

Chorion

Crosslinking

Decellularisation

Extra‐cellular matrix

Functionalisation, functionalization

Hydrogel

Methacrylation

Photopolymerisation, photopolymerization

Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels i

Abstract

Hydrogels based on solubilized extracellular matrices (ECMs) represent a promising source for the creation of cell‐instructive scaffolds for tissue engineering. The diverse range of biochemical cues retained by placental ECM drive cell proliferation, differentiation, and function for all tissue types in utero while providing mechanical support for tissue growth. However, the formation of ECM‐derived hydrogels typically relies on the thermally‐induced self‐assembly of collagenous polypeptides, severely limiting the control over the resulting biochemical and mechanical hydrogel properties. Decellularisation methods of the biomaterial may alter these properties of the final hydrogel. Here the properties of photocrosslinkable hydrogels derived from different decellularisation methods, based on human placental ECM digests, has been explored to overcome aforementioned limitations and allow for precise control over physicochemical properties, while still retaining cell‐ instructive bioactivity. Briefly, human full‐term placenta was obtained following caesarean delivery from consenting donors with ethics approval. Chorionic villi tissue was decellularised using N‐lauroyl sarcosine and Triton X‐100 detergent‐based methods, respectively, and enzymatically digested. Each resulting ECM digest was then functionalized with methacrylic anhydride to result in methacryloyl‐substituted placental ECM (PlacMA). Precursor solutions of PlacMA were liquid at room temperature and formed mechanically stable and transparent hydrogels by visible light‐induced photocrosslinking via cytocompatible free radical chain polymerisation. The mechanical properties of the resulting hydrogels could be tailored through the variation of PlacMA concentration and crosslinking parameters. Methacryloyl‐ functionalisation of native ECM digests from N‐lauroyl sarcosine and Triton X‐100 detergent based methods yield different properties in the final hydrogel. In particular, Triton X‐100 decellularised material provided less batch‐to‐batch variation in hydrogel physicochemical properties and cell viability. Both methods still allowed for the formation of mechanically stable hydrogels with improved control over physicochemical properties. This method can be applied to a large variety of native tissue digests, allowing researchers to mimic native cellular microenvironments in vitro and develop cell‐instructive scaffolds for tissue engineering

ii Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels

Table of Contents

Keywords ...... i

Abstract ...... ii

List of Figures ...... v

List of Tables ...... vi

List of Abbreviations ...... vii

Preface ...... ix

Statement of Original Authorship ...... ix

Acknowledgements ...... x

CHAPTER 1: INTRODUCTION AND LITERATURE REVIEW ...... 1

1.1 Hydrogels ...... 1

1.2 Native Tissue as a Source for Biomaterials ...... 9

1.3 Placenta ...... 11

1.4 Hypothesis and Objectives ...... 16

CHAPTER 2: MATERIALS AND METHODS ...... 17

2.1 Decellularisation ...... 17

2.2 Determination of DNA and sGAG Content ...... 19

2.3 dECM Solubilisation ...... 19

2.4 Functionalisation of sECM with Methacryloyl Groups ...... 20

2.5 1Hydrogen Nuclear Magnetic Resonance ...... 20

2.6 Determination of the Degree of sECM Functionalisation ...... 20

2.7 Preparation of PlacMA hydrogels ...... 21

2.8 Effective Swelling ...... 21

2.9 Mechanical Testing ...... 22

2.10 Cell Expansion ...... 23

2.11 Cell Encapsulation and Culture ...... 24

2.12 Cell Viability ...... 24

2.13 Cell Morphology ...... 25

Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels iii

2.14 Statistical Analysis ...... 25

CHAPTER 3: RESULTS ...... 27

3.1 Decellularisation Efficacy ...... 27

3.2 Solubilisation of dECM and Functionalisation with Methacrylic Anhydride ...... 29

3.3 Mechanical Testing ...... 32

3.4 Effective Swelling ...... 35

3.5 Cell Culture ...... 36

3.6 Cell Morphology ...... 37

CHAPTER 4: DISCUSSION ...... 41

4.1 Decellularisation Assessment ...... 41

4.2 Functionalisation Assessment ...... 43

4.3 Mechanical Assessment ...... 45

4.4 Cell Viability and Morphological Assessment ...... 47

4.5 further work ...... 48

CHAPTER 5: CONCLUSION ...... 51

BIBLIOGRAPHY ...... 53

APPENDICES ...... 67

Appendix A Decellularisation Data ...... 6 7

Appendix B Mechanical Data ...... 68

Appendix C Cellular Viability Data ...... 70

Appendix D ImageJ Code ...... 71

ANNEX 1 ...... 72

iv Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels

List of Figures

Figure 1: Illustration of cells in in 2‐dimensional and 3‐dimensional environments...... 2

Figure 2: Diagram of major ECM components...... 3

Figure 3: Biopolymer based hydrogels by polymer type...... 6

Figure 4: Functionalisation of proteins via methacrylic anhydride reaction...... 7

Figure 5: Molecular comparisons of detergents...... 10

Figure 6: Medical history of first recorded applications of placenta and its derivatives...... 12

Figure 7: Representative cross sectioning of human placenta...... 13

Figure 8: Overview of process to produce PlacMA from human chorionic tissue...... 17

Figure 9: Anatomy of Placenta [108]...... 18

Figure 10: PlacMA hydrogel casting and determination of physical properties...... 23

Figure 11: Decellularisation of placental tissue over time...... 28

Figure 12: Decellularisation efficacy by presence of DNA...... 29

Figure 13: 1H‐NMR Spectra for pre‐ and post‐functionalised sECM...... 30

Figure 14: Results of TNBS Assay...... 31

Figure 15: Representative Stress/ Strain behaviour of PlacMA hydrogels...... 33

Figure 16: Mechanical Data for PlacMA Hydrogels...... 34

Figure 17: Effective swelling of PlacMA hydrogels...... 35

Figure 18: Cell Viability in PlacMA Hydrogels...... 37

Figure 19: Cell Morphology of Cultured PlacMA Hydrogel...... 38

Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels v

List of Tables

Table 1: Growth Factors Identified in Placental Tissue (non‐exhaustive list) ...... 14

Table 2: Timeline of decellularisation ...... 18

Table 3: DNA Content, Before/ After decellularisation. DNA/ Tissue, ng/mg ...... 67

Table 4: sGAG Content, Before/ After decellularisation. sGAG/ Tissue, µg/mg ...... 67

Table 5: Compressive Modulus of PlacMA hydrogels (kPa) ...... 68

Table 6: Strain rate (%) of PlacMA hydrogels at failure ...... 68

Table 7: Stress rate (kPa) of PlacMA hydrogels at failure ...... 69

Table 8: Relative Swelling Change (mg) in Mass in PlacMA Hydrogel ...... 69

Table 9: Viability Data as Live/Total cell count (%) ...... 70

vi Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels

List of Abbreviations

3D Three Dimensional AF Alexa Fluor AM Amniotic Membrane ARC Australian Research Council CARF Central Analytical Research Facility CLSM Confocal Laser Scanning Microscopy CPME Chemistry, Physics, and Mechanical Engineering DAPI 4',6‐diamidino‐2‐phenylindole dECM Decellularised Extracellular Matrix DMEM Dulbecco's Modified Eagle Media DMMB 1,9‐dimethyl‐methylene blue EDTA Ethylenediaminetetraacetic acid ELISA Enzyme‐Linked Immunosorbent Assay EGF Epidermal Growth Factor FBS Fetal Bovine Serum FDA Fluorescein Diacetate G Glycine GAG Glycosaminoglycan GelMA Gelatin Methacryloyl GF Growth Factor HB‐EGF Heparin Binding EGF‐Like Growth Factor H‐NMR Nuclear Magnetic Resonance IC2959 Irgacure 2959; 1‐[4‐(2‐hydroxyethoxy)phenyl]‐2‐hydroxy‐2‐methylpropan‐1‐ one IHBI Institute for Health and Biomedical Innovation IGF Insulin‐Like Growth Factor IGFBP‐1 Insulin‐Like Growth Factor Binding Protein‐1 K Lysine MA Methacryloyl MAAh Methacrylic Anhydride MEM Minimum Essentials Media MS/MS Tandem Mass Spectrometry MSC Mesenchymal Stem Cells MWCO Molecular Weight Cut Off PBE Phosphate Buffered EDTA PBS Phosphate Buffered Saline PC2 Physical Containment Level 2 PFA Paraformaldehyde PI Propidium Iodide

Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels vii

PlacMA Methacryloyl‐Substituted Placental ECM PPE Personal Protective Equipment P Proline QUT Queensland University of Technology RM Regenerative Medicine RT Room Temperature SDS Sodium Dodecyl Sulfate sECM Solubilised Extracellular Matrix SLS N‐Lauroyl Sarcosine, Sarkosyl TE Tissue Engineering TX Triton X‐100 UQCCR University of Queensland Centre for Clinical Research UV Ultraviolet VEGF Vascular Endothelial Growth Factor

viii Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels Preface

This submission first includes a thesis based upon the research performed in 2019 at Institute of Health and Biomedical Innovation, Queensland University of Technology, Australia which has not been previously submitted to any other higher education institution. Annex 1, at the end of this document based upon the research performed in 2018 while at Hubrecht Institute, Universitair Medisch Centrum Utrecht, Nederland which was been previously submitted to fulfil the requirements of the Master of Biofabrication at Universitair Medisch Centrum Utrecht.

Statement of Original Authorship

The work contained in this thesis has not been previously submitted to meet requirements for an award at this or any other higher education institution. To the best of my knowledge and belief, the thesis contains no material previously published or written by another person except where due reference is made.

Signature: QUT Verified Signature

Date: 17-June-2020

Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels ix

Acknowledgements

First, I wish to thank staff and colleagues who have given their time, equipment, and resources for exploring new methods that allowed this project to continue. Dr Jon Harris (QUT), Dr Annalese Semmler (QUT), Yaoqin Hong (QIMR), Dr Peter Katsovic (QUT), Jongryul Park (QUT) and Gelomics (Brisbane, Australia). Thank you to Onur Bas (QUT) and Stephen Pahoff (QUT) for clutch advice and guidance with hydrogels. Thank you to Colin Lloyd (Australia) for alternately configured visual advice. Thank you to Susan Brown and Seen Lim Sing (UQCCR) for handling and provision of tissue.

Thank you to friends and collegues in Regenerative Medicine group (QUT), Nephrology and Orthopaedics (UMC, Utrecht), and the Biofabrication program. Thank you to my Australian family and my adoptive Dutch family for their continued support during this education.

Thank you to Leah Brew who re‐ignited and continued the flame of this project while I left to study electrospinning in the Netherlands. Thank you also for the advice in navigating Utrecht.

Thank you to Distinguished Professor Dietmar Hutmacher (QUT) for the support and guidance throughout this project. Thank you again for setting up the international biofabrication program, which has been a once‐in‐a‐lifetime opportunity. Thank you to Joanne Richardson (QUT), Dr Jacqueline Alblas (UMC, Utrecht), and Roos Nieuwenhuis (Netherlands) for smoothing over the application processes and making it seem easy.

Finally, I wish to thank my supervisors Drs Abbas Shafiee and Christoph Meinert for being a seemingly ever‐present and depthless fonts of knowledge, wisdom, patience, expertise, and passion. They are better supervisors than I am a student.

x Effect of Decellularisation Methods on Methacryloyl‐Substituted Placental ECM Hydrogels

Chapter 1: Introduction and Literature Review

In this literature review hydrogels and placenta are explored. Types of hydrogels will be defined along with the most common methods used to produce them. Placenta is then described anatomically along with the brief history of medical employment of this organ. This section concludes with the respective relevance to this project, the hypothesis, and aims of the project.

1.1 HYDROGELS

Hydrogel Uses and Properties

Peppas et al., (2006) [1] describe hydrogels as hydrophilic polymer chains swollen with water and prevented from dissolving by their crosslinked network. This high‐water content and biocompatibility make them ideal for biological applications [1]. The ability of hydrogels to support cells in a 3D environment has garnered significant interest and promise for tissue engineering and regenerative medicine (TERM). Operating as a cellular scaffold, hydrogels offer improved analogues to in vivo testing than existing flat tissue‐culture‐technologies [2]– [5]. Hydrogel properties are determined in part by the base material. Biological sources typically have good cytocompatibility but poor control over crosslinking and mechanical properties [5]–[13]. Synthetic sources however typically have good crosslinking control and mechanical properties but poor cytocompatibility [12]–[14]. Combinations of synthetic and biologic, so called semi‐synthetic hydrogels, offer a modular mix of control and biocompatibility properties [15]–[18].

Accurate models allow research to be translated into treatment by displaying safety and efficacy [19], [20]. The current track for this translation can be broadly described as ‘benchtop’ science results transcribing into animal models, which then lead into small‐scale human trials before graduating into large human trials and then continued monitoring when released to the public [20], [21]. Failure to evidence efficacy is a common cause of failure when transitioning through these stages [11]. These failures are often expensive in terms of

Chapter 1: Introduction and Literature Review 1

money and investor confidence [19], [22]–[25]. Hydrogels as a research platform therefore offer better accuracy for studying cellular interactions and reactions compared to traditional cell culture methods [7], [15].

Figure 1: Illustration of cells in in 2‐dimensional and 3‐dimensional environments.

Representation of morphological differences of cells cultured on (top) or encapsulated within (bottom) a collagen matrix [26], [27].

Cells are studied in vitro with flat and unphysiologically stiff tissue culture‐ware as a model for in vivo tissue [3], [5]. There are significant phenotypic differences however between these 2D models and true 3D biology [26]–[29] (Figure 1). In 2D culture, cells have relatively equal access to ingredients in media, whether it be gas, nutrient, or drug [28], [29]. Cells in vivo however reside in a more complex arrangements [28], [29]. This results in differences in morphology, proliferation, differentiation, migration, and access to media [30]–[33]. In identifying these contrasts, new in vitro systems are shifting towards utilising 3D cultures for increased accuracy and predictive value [34], [35]. As hydrogels are a 3D platform that can support cell culture, they are being employed to better model diseases and develop new treatments [2], [5], [7], [34]–[38].

Cells in vivo are supported by the ECM which in turn facilitates structural tissue cohesion. This 3D microenvironment consists of proteins, glycoproteins, glycosaminoglycans (GAGs), and polysaccharide molecules, which support structural and bio‐signalling direction. (Figure 2) [39]–[43]. Contained in this milieu are cell adhesion sites, matrix proteases, growth factors and cytokines, that influence the coordination of the heath and development for surrounding cells [39]–[43]. Adhesion sites bind with cells and transfer mechanical stimuli through

2 Chapter 1: Introduction and Literature Review

mechanotransduction, creating biochemical impulses in response [44], [45]. Growth factors and cytokines provide regulatory cues to guide growth [46]. Matrix proteases allow ECM remodelling and repair for continued maintenance and renewal [47]. Understanding these elements allows for creation of hydrogels with improved support for cell studies.

Figure 2: Diagram of major ECM components.

Fibronectin connection to integrin proteins in the phospholipid bilayer and subsequent connection to the cytoskeleton [48].

Hydrogels have a demonstrated use in both commercial and research settings according to their properties [12], [14], [49], [50]. Super porous hydrogels, such as polyacrylamide, are used in the manufacture of hygiene products as they can quickly remove surrounding water through numerous capillaries [12], [51]. Super absorbent hydrogels, such as cashew gum, can hold 1500 times their dry weight in water and are used in agriculture to prevent soil desiccation and allow nutrient dispersion [12], [52]. The flexible, semi‐permeable, and tailorable nature of hydrogels allows them to also function effectively as medical devices; 2‐ hydroxyl ethyl methacrylate (HEMA) is one example that is used as a contact lens; polyethylene glycol (PEG) an example of wound dressing [12], [53], [54]. Here the hydrogel forms a barrier, keeping the tissue underside wet while still allowing gas diffusion, and can prevent microbial migration [12], [55].

Hydrogels are also being explored for drug delivery. The aqueous content that fills the pores can be loaded with drugs for controlled delivery applications [56]. Materials used this way are chosen to work in tandem with the proximate environment and method of delivery, acting as ‘smart hydrogels’ [56], [57]. In this way control of the treatment diffusion is

Chapter 1: Introduction and Literature Review 3

achieved through the resulting degradation, swelling, or other changes to the hydrogel [49], [56], [57]. Oral delivery can use the pH changes through the gut tract, while implant delivery could use local degradative enzymes [56], [57].

Hydrogel Polymer Sources

Hydrogels are softer and more elastic than tissue culture plastic. Mechanical forces are dispersed along the polymer chains and fluid filled pores, allowing recovery after the gel has been deformed [58], [59]. Hydrogels can be characterised by polymer type and composition, physical composition and configuration, electrical charge, or crosslinking type [49], [60], [61]. The networks may be joined by covalent or ionic strong bonds, weak hydrogen bonds, crystal formations, or physical entanglements and junctions [61], [62]. Methods of preparing hydrogels vary based on desired bond type, concentration, and polymer type [61], [62]. Source of material to prepare hydrogels can be broadly classified into the synthetic, biological, and semi‐synthetic [49], [60], [61].

Synthetic hydrogels are most commonly carbon‐based monomers assembled through free radical polymerisation [12], [14], [50], [63]. Briefly, this can be explained in three steps. In the first step, the initiator agent interacts with the monomers and creates an active site [63]. In the second step, the active site bonds and propagates to other monomers [63]. Finally, the polymerisation is terminated either because all the radicals are used and no monomers remain, or the reaction has been elsewise inactivated [63]. Initiation can occur with physical stimuli such as; sonication, pressure, or electromagnetic radiation whether it be from light, heat, or ionising sources [12], [14], [50], [63]. Initiator agents may also be chemical stimuli such as pH, solvent composition, or molecular species [12], [14], [50], [63]. Through selection of monomer and crosslinking type, the physicochemical properties of a hydrogel can be tailored [7], [12], [15].

Biologically‐derived hydrogels (Figure 3) are noted for possessing cyto‐compatible and ‐ instructive properties [6]–[13].Unlike synthetic hydrogels, biological hydrogel materials commonly have no need for further modification to support cell growth [12], [15], [64]. As the material originally contributed to a living system, many of the support mechanisms persist through the processing and preparation stages of decellularisation and solubilisation [6], [7], [12]. ECM bound growth factors and other cell instructive molecules persist along with cellular adhesion sites which facilitate mechanotransduction, such as RGD protein

4 Chapter 1: Introduction and Literature Review

motifs on fibronectin, as well as enzymatic cleavage sites that allow cellular remodelling [6], [7], [65], [66].

Biopolymers, as an organic source for hydrogels, can be ordered into two groups; polysaccharides and proteins [67]. Polysaccharide hydrogels include cellulose, chitosan, and alginate [67]. Polysaccharide hydrogels offer low immunogenicity, high biocompatibility, biodegradability, and are a cost effective biomass with wide availability [12], [67]. Biopolymers typically suffer from poor mechanical properties, however but they can be improved through addition of glycosaminoglycans chondroitin sulphate and hyaluronic acid [67]. The gelation of polysaccharides is sensitive to temperature, pH, and redox reactions; stimuli abundant in biological tissue systems [67]. These gels have therefore been of interest in exploring; drug administration with chitosan for example; carboxymethylcellulose for wound dressing; and hyaluronic acid for bone TE [12], [67]–[69].

Protein hydrogels include the ECM monomers such as fibroin, elastin, collagen, and its derivative gelatin [12], [67]. Gelatin is the most widely studied gel as it is cost effective and broadly available biopolymer with bioactive properties [12], [49], [50], [60], [67]. Physical properties of gelatin, such as amino acid composition, mechanical, and immunogenicity depend on; the method of production, acid or alkali hydrolysis; tissue type, bone or skin; and domesticated animal origin, pig, cow, fish, or poultry [70]–[73]. Polysaccharide and protein groups can be combined, such as gelatin incorporating hyaluronic acid to mimic the native ECM for myo‐, chondro‐, and osteoblast culture [7], [12], [15], [66], [74]. Modification with other materials has also been explored, such as gelatin and hydroxyapatite for bone TE [11], [12], [75].

Protein hydrogels also include tissue that has been decellularised and solubilised using proteases, leaving the structural and instructive ECM microenvironment [6], [8], [10], [13], [18]. Removing host cells can be performed through mechanical, enzymatic or chemical means, such as strong ionic detergents, depending on optimal method for the selected tissue [6], [8], [10], [64], [76], [77]. Decellularisation is vital as the foreign molecules such as antigens, epitopes, and interleukins that promote inflammatory responses effects are removed, which would otherwise hinder cell studies and non‐allogenic transplants [6], [7], [77]. Following the decellularisation steps, an enzymatic digest breaks the ECM superstructure, allowing peptides to reform, thermally gelate, and be cast in moulds or 3D printed [6], [8], [10], [13], [18].

Chapter 1: Introduction and Literature Review 5

Figure 3: Biopolymer based hydrogels by polymer type.

Chart of hydrogels produced using biologically derived polymers [67].

Semisynthetic hydrogels are generated through the covalent addition of a functional group to the biological materials [16], [17], [78]–[80] (Figure 4). This functional group can be crosslinked in the presence of a non‐toxic chemical agent and exposure to short‐wavelength light [16]. This allows biological materials to act as synthetics in their now tuneable ability to control mechanical properties via concentration and exposure while retaining adhesion and cleave sites [16]. These tuneable properties are of great interest to TE research as it allows a hydrogel to fit within a specific biofabrication window [15], [81]. This window is a, typically narrow, band of parameters that mechanically matches a given tissue that also balances requirements for tissue culture [15]. Accurately reproducing the physicochemical

6 Chapter 1: Introduction and Literature Review

characteristics of tissue allows for more accurate modelling of disease and repair therein [5], [7].

Methods of Crosslinking

Physical properties of a hydrogel are vital, whether it be for supporting cell culture as a scaffold, surgical handling during implantation, or lifestyle demand when in use by a patient [12], [16], [82]–[88]. Further modification of organic hydrogels may be required to improve upon degradation resistance, gas and fluid dispersion rates, and mechanical properties such as elasticity, stiffness, tensile and shear strengths [12], [16], [82]–[88]. A common strategy to improve these physical properties is a crosslinking agent[12], [67], [88]. This agent is mixed into the hydrogel precursor solution and promotes additional crosslinking in the hydrogel network [50], [88]. Potential toxicity of the agent must be factored into as any residual agent may need to be washed out [14], [88].

Enzymatic methods link biopolymers according to availability and proximity of other sites [12], [73], [88]. Transglutaminase (TGase) is the most popular enzyme method as isoforms are already present in the human body and has been used clinically for neurodegenerative disorders and fibrotic patients [88]. TGase catalyses amide bonds between the carboxylic and amine groups of glutamic acid and lysine residues [12], [88]. Enzymatic methods are usually performed in the biological window of conditions, allowing biopolymers to not distort or denature when under extreme pH or temperature required for chemical crosslinking [73], [88]. Enzymes are expensive however and require specific substrates to work [88], [89].

Figure 4: Functionalisation of proteins via methacrylic anhydride reaction.

Both hydroxyl and amide groups can react with methacrylic anhydride producing methacrylate and methacrylamide groups, respectively [90].

Chapter 1: Introduction and Literature Review 7

Chemical crosslinking agents catalyse covalent bonding and generally deliver the highest degree of crosslinking with improved mechanical properties [14], [73], [88]. The agents are however, often cytotoxic and can increase the immunoinflammatory response, meaning the unreacted agent must be thoroughly washed out from the scaffold prior to use [88], [91]. Glutaraldehyde reacts with the hydroxyl, thiol, and free amine groups on proteins leading to very high levels of crosslinking from a widely available and cost effective solution [92]. For these reasons, glutaraldehyde is considered the gold standard for chemical crosslinking and has been used in soft tissue films, bone tissue engineering and transplanting cardiac valves [88], [93], [94]. The drawback to glutaraldehyde is the cytotoxicity, which is less of an issue for the recently discovered genipin, derived from gardenia plants [73], [88], [91].

Photopolymerisation uses photo‐chemical crosslinking agents that initiate covalent bond formation when exposed to light in the presence of a photoinitiator [95]–[97]. Highly reactive intermediates (free‐radicals) are the product of this photoactivation that can interact with different groups through chain polymerisation, such as the functionalised moieties of modified peptides [96], [98]. Crosslinking methods based on free‐radicals are used as they produce hydrogels with control over physicochemical properties according to set parameters [70], [98]. Light intensity, crosslinking time, degree of functionalisation, polymer and photoinitiator concentrations can be altered to tailor hydrogel properties [70], [98]. A common photoinitiator is Irgacure 2959 (IC2959; 1‐[4‐(2‐hydroxyethoxy)phenyl]‐2‐hydroxy‐ 2‐methylpropan‐1‐one) that creates an acetyl free radical in the presence of UV light [16], [70], [79]. UV light however is associated with mutagenicity as it can damage intracellular DNA and generate reactive oxygen species [99]. LAP (lithium phenyl‐2,4,6‐ trimethylbenzoylphosphinate) is considered a more promising photocrosslinking agent as it does not require UV light [70], [100].

An example of the possible functionalisation and creation of semisynthetic hydrogels is gelatin methacryloyl (GelMA). Methacrylic anhydride (MAAh) in a proteinaceous solution can react and form methacryloyl groups (a term that includes both resulting methacrylate and methacrylamide groups) functionalised onto proteins [90], [101]. These groups form most favourable with lysine and hydroxylysine amino acids [90], [101]. In the presence of a photoinitiator and an appropriate light source, these groups can be photocrosslinked in a controlled manner [16], [90]. GelMA is of great interest to TE, as it can be cast as a hydrogel, 3D printed, and used to create modular microfluidic blocks [16].

GelMA is the most popular material for studying semi‐synthetic hydrogels due to the availability and ubiquity of gelatin however this is not the only example of a functionalised

8 Chapter 1: Introduction and Literature Review

biopolymer [17], [50], [102]–[105]. Collagen‐methacrylate (CMA), methacrylated‐hyaluronic acid (MeHA), methacrylated‐tropoelastin (MeTro), and chondroitin sulphate‐methacrylate (CSMA) have been created by functionalising methacryoyl groups onto their respective ECM molecules [17], [102]–[105]. These polymers each have their own unique properties which can enhance culture conditions for cell studies, such as gelMA and MeHa being combined for increased stiffness to be more amenable to chondrocyte and osteocyte cells [16], [78], [90].

MAAh is not the only means of functionalising biopolymers. Wang et al., (2012) [106] created a bio‐adhesive by conjugating dextran with 2‐isocyanatoethyl methacrylate to create oxidised urethane dextran (Dex‐U‐AD). This material can be crosslinked with IC2959 and UV light to adhere to gelatin with more force than commercially available fibrin glues. Dex‐U‐AD was confirmed to be biocompatible with murine fibroblasts. Teramoto et al., (2012) [107] created fish gelatin methacylate (FG‐GMA) by functionalising the biopolymer with glycidyl methacrylate. The hydrogels were cast using IC2959 and UV light while the mechanical properties were enhanced through the addition of aluminium silicate nanofibers (imogolite).

In the current literature, there is a deserved increase in the popularity of semi‐synthetic hydrogels. By functionalising biopolymers with crosslinkable chemical groups, many promising 3D cell culture platforms have been created that can be tailored to the needs of a specific tissue [15], [26], [50]. This can allow for in vitro studies that better mimic in vivo systems [15], [26], [50].

1.2 NATIVE TISSUE AS A SOURCE FOR BIOMATERIALS

Tissue can be used as a biological source for cells and ECM [6], [76], [77], [108]. Cells collected from tissue can be expanded and studied or transplanted for therapeutic uses [76], [108]. ECM collected can be likewise studied and transplanted or can be solubilised to produce hydrogels [8], [10], [13], [109]. Isolating either of these materials requires decellularisation and discarding of the undesired material [77]. The desired final product determines which decellularisation process should be employed. Harvesting stem cells for instance requires a method that dismantles ECM but not compromise cellular viability [108], [110]. Collecting ECM, in contrast, requires the cell removal to result in decellularised ECM (dECM) [8], [10], [13], [109]. When using dECM, it is vital to ensure the removal of cellular components such as epitopes, antigens and interleukins, as these can induce an immunoinflammatory response [6], [7], [77]. For purposes of casting hydrogels, dECM is typically solubilised as the

Chapter 1: Introduction and Literature Review 9

small hydrophilic particles can form a fluid‐filled network [50], [88]. This solubilised ECM (sECM) is generally produced through use of proteases [6], [8], [10], [13].

Producing dECM for solubilisation can be accomplished with many techniques and methods; enzymatic, physical, and chemical based [6], [77], [111]. Different tissue types and different procedures can greatly affect the resulting physicochemical properties of resulting sECM hydrogels [70], [77], [111]–[114]. Enzymes such as collagenase lyse the structural proteins of tissue and centrifugation can be used to separate and collect the desired material [113]. Physical methods such as high hydrostatic pressure burst cell membranes and lysed cellular material is later washed out [111]. An example of chemical processing is the production of gelatin, wherein tissue undergoes acid or alkali hydrolysis to decellularise and solubilise cell membranes and structural proteins [11], [70]. Detergent based methods are also an example of chemical decellularisation.

Detergent based decellularisation methods work by disruption the cell membrane layers [64], [77], [115]–[119]. Ionic detergents such as sodium dodecyl sulphate (SDS) and N‐lauroyl sarcosine (sodium lauroyl sarcosine; sarkosyl; SLS) have a long hydrophobic tail and large polar‐head region that solubilises the phospholipid bi‐layer (Figure 5) [115]–[117] Triton X‐ 100 (TX) is non‐ionic and considered a mild detergent that disrupts lipid‐lipid and lipid‐ protein interactions resulting in less damage to ECM proteins [18], [77], [114], [119].

Figure 5: Molecular comparisons of detergents.

A) Sodium dodecyl sulphate B) N‐lauroyl sarcosine C) Triton X‐100

sECM hydrogels are commonly used to culture cells from similar proximity only from a different host, such as osteoblasts on decellularised bone, myocytes onto cardiac ECM, and

10 Chapter 1: Introduction and Literature Review

neuroblastoma cells onto hydrogels derived from central nervous tissue [6]–[13], [83]. Clinically, sECM has been used as patches for soft tissue injury, xenografts, and is currently being explored for injection to an internal injury to promote tissue regeneration [6], [12], [83]. Mechanical stability of sECM hydrogels relies on thermo‐reversible physical protein interactions, which produces weaker gels with high batch‐to‐batch variations [5], [6], [8], [9], [42].

Tissue has been used to create semi‐synthetic hydrogels, functionalising many of the ECM molecules rather than just isolated species [102], [120]. Ali et al (2019) [120] decellularised whole porcine kidney using perfusion before solubilising and functionalising with methacryloyl groups to produce kidney decellularised ECM methacrylate (KdECMMA). Hydrogels were formed using IC2959 and UV light using low concentrations of 1 % ‐ 3 % where the material was able to form controlled structures through extrusion printing. Human primary kidney cells were used in the study where it was found that KdECMMA was able to support cell maturation and tissue formation. This demonstrates that whole tissue is also a viable and promising source for semi‐synthetic hydrogels.

Visser et al., (2015) [102] extracted cartilage, meniscus, and tendon tissues from equine stifle joints, which were separately decellularised, solubilised, and then functionalised with MAAh to produce cartMA, menMA, and tendMA respectively. These polymers were then combined with gelMA and used to study mesenchymal stem cells (MSCs) and chondrocytes. Methacrylation increased the stiffness of the hydrogels, which is an important factor for musculoskeletal repair. ECM production by encapsulated chondrocytes was however, less than those in the gelMA‐only controls. Under cartMA, the MSCs produced more collagen and less glycosaminoglycans, while there was better redifferentiation of chondrocytes in tendMA.

1.3 PLACENTA

The placenta is a temporary organ that in its short existence facilitates the growth of a zygote through to term delivery [6], [7]. The physiological function and cellular content of placenta has been successfully applied in multiple therapeutic settings [121]. Caesarean delivery and decellularisation protocols allow placental tissue to be obtained while minimising allogenic immunological responses [108]. Using this placental tissue for tissue engineering, hydrogels may provide further benefit to the success reported in previous applications [121].

Chapter 1: Introduction and Literature Review 11

•Compendium of 1593 Materia Medica 1910 •Skin Grafts •Genital 1934 Reconstruction •Ophthalmological 1940 Surgery 1946 •Burn Injuries

1947 •Abdominal Surgery •Preparation 1965 Protocols •Endothelial 1997 Progenitor Cells 1999 •FDA Approval

2004 •Stem Cells

Figure 6: Medical history of first recorded applications of placenta and its derivatives.

Silini et al. (2016) [121] documented the history of placenta in medicine (Figure 6). Briefly, the Chinese medical text, Compendium of Materia Medica, prescribes placenta for consumption as a restorative cure‐all [122]–[124]. In 1910, placenta was first used as a skin graft [125]. In 1934, a Parisienne team used placenta for vaginal reconstructive surgery [126]. In 1940 conjunctival defects are repaired using placental tissues [127]. Second degree burns were first treated using placental tissue in 1946 [128]. Moreover in that same year, placenta was first used as part of surgical procedures to close bowel fistulas [129]. In 1965, protocols for processing and storing were published [130]. Endothelial progenitor cells discovered in placental tissue in 2013 [110]. In 1999 the FDA approved acellular placental tissue for wound covering [131], [132]. In 2004, the placental stem cells were identified and extensively characterised in placental tissue [28], [110], [133]–[135].

The human placenta develops from the embryo and connects to the fetus via the umbilical cord. Physiologically, the placenta provides nutrient, gas, and waste exchange, as well as masking the fetus from the maternal immune system [136]. This solute and gas exchange is made possible through the highly vascular chorionic villi (CV) that interface with the vasculature in the maternal decidua [121] (Figure 7). In addition, the placenta functions as an endocrine organ and produces hormones and growth factors [40], [43], [76], [121], [137], [138] (Table 1). The placenta is also a source of structural proteins and GAGs such as heparin‐

12 Chapter 1: Introduction and Literature Review

sulphate, chondroitin‐sulphate, keratin‐sulphate, and dermatan‐sulphate [40], [41]. This broad and rich range of bio‐active components of ECM molecules allows the placenta to support the development of all human tissue types [7], [40]. As such, it has been suggested that placental tissue has potential applications in TERM as a cell instructive biomaterial [121].

Figure 7: Representative cross sectioning of human placenta.

Cross section of human placental chorionic villi [139], A) Cross section of human placenta in utero B) Branches of chorionic villi, b) terminal villi C) Histologic section of human placental tissue, (c1) cytotrophoblast (c2) syncytotrophoblast. D) Layer‐by‐layer of placenta layer [140]. The amnion is tissue that encloses the fetus and amniotic fluid. The chorion interfaces with the maternal side.

The human placenta develops from the embryo and connects to the fetus via the umbilical cord. Physiologically, the placenta provides nutrient, gas, and waste exchange, as well as masking the fetus from the maternal immune system [136]. This solute and gas exchange is made possible through the highly vascular chorionic villi (CV) that interface with the vasculature in the maternal decidua [121] (Figure 7). In addition, the placenta functions as an endocrine organ and produces hormones and growth factors [40], [43], [76], [121], [137], [138] (Table 1). The placenta is also a source of structural proteins and GAGs such as heparin‐ , chondroitin‐, keratin‐, and dermatan‐sulphates [40], [41]. This broad and rich range of bio‐ active components of ECM molecules allows the placenta to support the development of all human tissue types [7], [40]. As such, it has been suggested that placental tissue has potential applications in TERM as a cell instructive biomaterial [121].

Chapter 1: Introduction and Literature Review 13

The placenta weighs on average 500 g at full term birth [108][141]. Australia had 95,894 caesarean births (32%) in 2011 [142]. This represents 47947 kg of vascularised, growth factor laden, stem and progenitor cell providing biomass that can be sourced from a medical by‐ product that is typically discarded.

Table 1: Growth Factors Identified in Placental Tissue (non‐exhaustive list)

Human placenta has also been explored as a source for hydrogel polymers. Zhang et al., (2019) [143] used such hydrogels to study hair follicle regeneration. The placenta was decellularised, solubilised, and then formed into a hydrogel and cultured with dermal papilla cells. When these cells co‐grafted with murine epidermal cells, the hair‐inductive capacity of the follicle was restored. Sosnowska et al., (2019) used ECM biopolymers derived from human placenta to support primary islet cells and found a significantly higher insulin secretion over the controls. Other uses for placenta have been for wound dressing from the amnion, or for stem cells from the chorion and umbilicus (see Figure 6) [121].

Francis et al. (2017) [13] have demonstrated hydrogels made from human placenta in remediating damage from cardiac ischemia. In this paper, chorionic villi were first decellularised with N‐lauroyl sarcosine detergent, then solubilised with pepsin, and it was discovered the DNA content was similar to the commercially available Matrigel product (Corning, USA). The solubilised material was then thermally crosslinked to form a hydrogel in vitro. These gels were studied in vivo by injection into rat models of myocardial infarction

14 Chapter 1: Introduction and Literature Review

and monitored for 4 weeks. It was discovered that the hydrogel would localise to the injection site, present with less scarring after injury, restore myocardial electrical synchronisation, and not have significant decrease in mortality or cardiac output. These hydrogels represent novel, xeno‐ and tumor‐material‐free platforms rich in structural proteins and growth factors with positive associations for regeneration, angiogenesis, stem cell recruitment and reductions in scarring and inflammation [13]. The lack of placental ECM functionalisation however, severely limited the range of mechanical properties achievable with this system.

In this project, the technique of functionalisation will be applied to create semisynthetic hydrogels from solubilised human placental ECM. This biological material will first be decellularised and then enzymatically digested to expose ECM peptides for functionalisation with MAAh. Photocrosslinking will create covalent crosslinking to give control over the mechanical properties. This should produce a cell instructive hydrogel that’s able to overcome the mechanical limitations of other placental‐derived hydrogels and make it an excellent candidate for 3D cell culture, tissue engineering and regenerative medicine applications.

Chapter 1: Introduction and Literature Review 15

1.4 HYPOTHESIS AND OBJECTIVES

We hypothesise that decellularised placental ECM can be functionalised with methacryloyl groups to result in a hydrogel biomaterial with highly tuneable physicochemical and cell‐ instructive properties. We further hypothesise that the application of varying decellularisation method will influence the properties of the resulting hydrogel.

This research aims to investigate the effects of different decellularisation methods on the physicochemical properties of methacryloyl‐functionalised placental ECM (PlacMA). By doing so, we attempt to optimise the choice of decellularisation methods when creating PlacMA. This will allow for a cell instructive, physicochemically tailorable culture platform for 3D cell culture, tissue engineering and regenerative medicine applications.

To address this hypothesis, the project structured into the following three objectives:

1. Creation of a semi‐synthetic hydrogel from human placental tissue using

different decellularisation methods

2. Characterise the physicochemical properties of resulting hydrogels

3. Assess the viability of the hydrogels as a 3D cell culture platform

16 Chapter 1: Introduction and Literature Review

Chapter 2: Materials and Methods

The goal of this project was to characterise and assess the differences from two detergent based methods of decellularisation on photocrosslinkable hydrogels produced from human placenta tissue. The two detergents were N‐lauroyl sarcosine (SLS; Sigma‐Aldrich, USA) and Triton X‐100 (TX; Sigma‐Aldrich, USA). Each part of the process was assessed to determine the effect of different decellularisation methods. Characterisation was carried out through mechanical and cellular viability testing. Briefly, human chorionic villi (CV; hereby referred to as native or raw tissue) were mechanically separated and washed with decellularisation detergents and wash protocols (Figure 8). The resulting decellularised ECM (dECM) was lyophilised and solubilised through enzymatic digestion. The resulting solubilised ECM (sECM) was functionalised by reaction with methacrylic anhydride (MAAh), dialysed, and lyophilised to yield methacryloyl‐substituted placental ECM (PlacMA).

Figure 8: Overview of process to produce PlacMA from human chorionic tissue.

2.1 DECELLULARISATION

Human placenta from full‐term, caesarean birth was obtained with informed consent from healthy donors with ethical clearance (HREC/09/QRBW/14). The donors were split into samples, with the first sample containing a single donor. The second sample contained mixed CV from three different donors. CV were separated interchangeably by Dr Abbas Shafiee and Seen Lin Sim (UQCCR) using established methods [73], [101] . Briefly, the umbilical cord and amniotic membrane were removed and discarded by placing the placenta fetal side up (Figure 9). Exposed chorion was then removed by leaving a ~5 mm layer of maternal tissue. The tissue was kept wet and briefly washed with Hanks Balanced Saline Solution (HBSS) at 4 °C. CV was then cut to ~ 3 cm3 portions using a scalpel, frozen and stored at ‐80 °C until further use.

Chapter 2: Materials and Methods 17

Figure 9: Anatomy of Placenta [108]. Stored tissue was placed into a high‐speed food processor (Russell Hobbs, UK) on pulse setting until tissue had a puree consistency. Tissue was transferred to 1 L glass Schott bottles, split evenly per sample (65 g – 80 g) then placed on a shaker plate at 4 °C for decellularisation.

Tissue was then washed with ultrapure water (ddH2O) three times a day, for 2 days. Tissue was then decellularised using two different detergents, respectively (Decellularisation Method 1 or Decellularisation Method 2). Samples were taken each day for analysis and washed 10 times to remove decellularisation solution. After decellularisation, the tissue was labelled dECM.

Decellularisation Method 1

This decellularisation method was based on established protocol previously published by

Skårdal et al. (2015) [18] (Table 2). Following the previous ddH2O wash step for 2 days, the tissue was immersed in 2% (v/v) TX, exchanging twice a day, for 8 days. Then tissue was immersed in 2% (v/v) TX, 0.1% (v/v) NH4OH (Sigma‐Aldrich, USA) exchanging twice a day, for one day. Tissue was then washed again with ddH2O for 2 days, washing as needed until no bubbles remained after shaking.

Decellularisation Method 2

This decellularisation method was based on established protocol previously published by

Francis et al. (2017) [13] (Table 2). Following the previous 2‐day ddH2O wash step, the tissue was immersed in 2% (w/v) SLS exchanging twice a day, for 9 days. Tissue was then washed again with ddH2O for 2 days, washing as needed until no bubbles remained after shaking.

Table 2: Timeline of decellularisation

18 Chapter 2: Materials and Methods

2.2 DETERMINATION OF DNA AND SGAG CONTENT

To determine the efficacy of decellularization and polysaccharide retention, the DNA and sGAG (sulphated glycosaminoglycan) content was explored. Determination of DNA and sGAG content was based on well‐established protocols [138], [139]. Spectrophotometry was performed on a POLARstar spectrophotomer (BMG Labtech, Germany). Briefly, tissue and ECM samples were frozen at – 80 °C, lyophilised, and the dry weight was recorded. Samples were then digested with 0.5 mg/ml Proteinase‐K (Thermofisher Scientific, USA) in phosphate buffered ethylenediaminetetraacetic acid (PBE, pH 7.1) overnight at 56 °C under constant agitation at 900 rpm (Eppendorf ThermoMixer, USA).

DNA concentration in the digests was measured using the Quant‐iT™ PicoGreen® dsDNA quantification assay (Invitrogen, USA) following the manufacturer’s instructions. DNA assay samples were diluted to 1:1000 for all time points except day 11, which used a 1:10 dilution. 100 µl of diluted digest was pipetted into a black 96 well plate (supplied by Regenerative Medicine group, Australia) containing 100 µl of 2x Picogreen dye. Fluorescence was measured for 485 nm excitation and 520 nm emission. The average of each sample was blanked against the mean of a 0 ng/ml concentration for DNA. sGAG concentrations were measured using the dimethyl‐methylene blue (DMMB) assay at pH 1.5 [144], [145]. All sGAG assay samples were diluted 1:10. 30 µl of sample was pipetted into a clear 96 well plate (Corning, USA) with 300 µl of DMMB dye (supplied by Cartilage Research Laboratory group, Australia). A standard curve was generated using shark chondroitin sulfate (CS; Sigma‐Aldrich, USA). Absorbance was measured at 525 Nm and 595 then subtracting the 595 nm results from the 525 nm results [145]. The mean of the blank was subtracted from each sample.

2.3 dECM SOLUBILISATION

Following decellularisation, the dECM was lyophilised to determine the dry weight for digestion. Enzymatic digestion was performed to produce soluble polypeptides that could be functionalised, then photocrosslinked. The solubilisation method based on a well‐established protocol previously published by Freytes et al. (2016) [137]. Briefly, lyophilised dECM was digested with pepsin (Sigma‐Aldrich, USA) at a ratio of 1 g dECM/100 mg pepsin in a volume of 100 mL of 0.01 M hydrochloric acid (HCl). This mixture was stirred for 3 days at room temperature until the digest became viscous and cloudy. Digestion was terminated by

Chapter 2: Materials and Methods 19

deactivation of pepsin by titration of the digest to pH 8 using 1 M NaOH. Samples were labelled sECM (solubilised ECM).

2.4 FUNCTIONALISATION OF sECM WITH METHACRYLOYL GROUPS

Following solubilisation, the sECM solution was functionalised, using the previous lyophilised dry weight to determine the amount of methacrylic anhydride (MAAh; Sigma‐ Aldrich, USA) to add. The functionalisation method was based on established protocols previously published by Loessner et al. (2016) [16] and Annabi et al. (2014) [17]. These steps were as follows. sECM was diluted with carbonate‐bicarbonate buffer (CBB; pH 9) at 4 °C to a ratio of 1 g sECM/ 100 ml CBB. MAAh was weighed and added to the sECM solution at a ratio of 0.6 g MAAh/1 g sECM. The solution was stirred on ice (2‐10 °C), protected from light to prevent photo‐interference and gelation of material. The pH was regularly adjusted to 8 ‐ 9 using 5M NaOH for 8 hours in order to favour the reaction going forward. The functionalised material was then transferred to benzoylated dialysis tubing (2000 MWCO; Sigma‐Aldrich, USA) to remove unreacted methacrylic acid. Dialysis was performed against ddH2O for 7 days, changing water twice daily. The product was then frozen at ‐80 °C and lyophilised.

2.5 1HYDROGEN NUCLEAR MAGNETIC RESONANCE

1Proton Nuclear Magnetic Resonance (1H‐NMR) was performed using a Bruker Advance 600 (Bruker, USA) at the QUT NMR facility (QUT, Brisbane) to confirm successful methacryloyl functionalisation of sECM. Briefly, sECM and PlacMA samples were dissolved at 1% w/v in

1 90% v/v ddH2O/10% v/v D2O (Sigma‐Aldrich, USA) and 0.05M of HCl, and H‐NMR spectra were obtained using water suppression.

2.6 DETERMINATION OF THE DEGREE OF sECM FUNCTIONALISATION

To quantitatively determine the degree of amines functionalised with methacryloyl groups (DoAF; methacrylation), a 2,4,6‐Trinitrobenzenesulfonic acid solution/ Picrylsulfonic acid (TNBS; Sigma‐Aldrich, USA) assay was used [146]. A stock solution was prepared from each sample before (sECM) and after functionalisation (PlacMA). A standard curve was created

20 Chapter 2: Materials and Methods

using 1:2 dilutions ranging from 500 µg/ml to 31.25 µg/ml of unfunctionalised sECM and PlacMA, respectively. Then, 0.05% w/v TNBS was added to each well and incubated at 37 °C for 2 h protected for light. Absorbance was read at 335 Nm.

The DoAF was calculated using the slope (m) of each dilution curve in the equation below.

𝑚 𝐷𝑜𝐴𝐹 100% 1 (1) 𝑚

2.7 PREPARATION OF PLACMA HYDROGELS

Lyophilised PlacMA was dissolved in cold PBS (4 °C, pH 7.5) at 0.5% w/v and 1% w/v respectively, overnight protected from light. 0.15% w/v lithium phenyl‐2,4,6‐ trimethylbenzoylphosphinate (LAP; TCI Chemicals, Japan) was added before crosslinking and thoroughly mixed via Gilson positive displacement pipetting, then kept on ice and protected from light until crosslinking. Hydrogels were cast on a Teflon surface (custom made by IHBI, Australia), covered with a glass slide (Figure 10, panel A) using two microscope slides as height spacers (2mm). Hydrogel precursor solutions (with or without cells) were photocrosslinked by exposure to 405 nm light in a Luna Crosslinker (Gelomics, Australia), and transferred to PBS (cell‐free hydrogels) or cell culture media (cell‐laden hydrogels).

Gels were prepared using PlacMA concentrations of 0.5% w/v and 1% w/v, and photocrosslinked for 15, 30, and 45 seconds, respectively (Table 3). 40 µl of solution was used per gel. Gilson positive displacement pipettes were used to handle the viscous solution. After casting, hydrogels were weighed and incubated overnight in 37 °C PBS. Gels were weighed again and the change in mass recorded.

2.8 EFFECTIVE SWELLING

The effective swelling of PlacMA hydrogels was determined as the ratio of the hydrogel wet weight following overnight incubation in PBS at 37 °C on a shaker plate, protected from light to wet weight immediately following hydrogel preparation.

Chapter 2: Materials and Methods 21

2.9 MECHANICAL TESTING

The hydrogel surface area was determined by using stereomicroscopy images (Nikon SMZ 745T) and FiJi (1.52p) prior to mechanical testing (Figure 10, panel B). The compressive modulus of hydrogels submerged in PBS at 37 °C was measured in unconfined compression tests using an Instron 5848 microtester equipped with a 5 N load cell (Instron, Melbourne, VIC, Australia) (Figure 11, panel C). Blue food dye (Queen, Australia) was added to the PBS to facilitate location of gel when in the water bath. Constructs were compressed at 0.01 mm/s using a non‐porous aluminium indenter and the modulus was determined as the slope of the stress‐strain curve from 0.1 – 0.15 mm/mm strain [16]. To accurately determine hydrogel strains, first, the height of each gel was determined by plotting force/extension data (Figure 10, panel D). Then the ‘toe’ and linear loading sections of the curve were identified via convergence of data. Height of the gel was determined as the transition point into these sections. Failure stress and failure strain were determined by detection and appearance of a sudden drop in the stress/strain curve.

22 Chapter 2: Materials and Methods

Figure 10: PlacMA hydrogel casting and determination of physical properties.

A) Crosslinked hydrogels (arrow) on Teflon surface, beneath glass slide, and removed from Teflon (inset). B) Surface area determination via stereomicroscopy, (b1) hydrogel proper (b2) blue‐dyed PBS. C) Mechanical testing setup, (c1) load cell (c2) indenter (c3) waterbath D) Force/ Extension curve showing the toe (d1) and linear loading (d2) sections. Blue line indicates mean force and grey band is 95% confidence interval.

2.10 CELL EXPANSION

Bone marrow Mesenchymal Stem Cells (MSCs; provided by Regenerative Medicine group, IHBI, Australia) at passage 8 were thawed in a 37 °C waterbath and suspended in 10ml media (Gibco alpha minimum essential media (MEM); Thermofisher, USA) and centrifuged for 5 minutes 300 relative centrifugal force, at room temperature. Supernatant containing dimethyl sulfoxide was discarded. 32ml of media containing 15% fetal bovine serum (FBS; provided by Regenerative Medicine group, IHBI, Australia) and 1% v/v penicillin/ streptomycin was added to the pellet and mixed. Cells containing media was then evenly split

Chapter 2: Materials and Methods 23

to 2, 75ml tissue culture flasks and monitored daily. Cells were incubated at 37 °C, 5% CO2 until 70% confluency, in a humidified cell culture incubator.

2.11 CELL ENCAPSULATION AND CULTURE

To understand how PlacMA hydrogels would perform as a 3D culture platform, MSCs were trypsinised (trypsin provided by Regenerative Medicine group, IHBI, Australia), counted (Countess Automated Cell Counter, Thermofisher, USA), and resuspended in hydrogel precursor solutions at 0.75 x 106 cells/ml with 94% viability immediately following trypsinisation. Hydrogels constructs were prepared as outlined in section 2.7, and cell‐laden hydrogels were cultured in 15% v/v FBS MEM media at 37 °C, 5% CO2 for 6 days.

2.12 CELL VIABILITY

At day 1 and 6 of culture, cell‐laden gels were removed from media and washed with PBS. Living and dead cells were visualised using fluorescein diacetate (FDA; provided by Regenerative Medicine group, IHBI, Australia) and propidium iodide (PI; provided by Regenerative Medicine group, IHBI, Australia), respectively. Briefly, two gels per group were incubated in PBS with 10 μg/mL FDA and 5 μg/mL PI for 5 min at room temperature and transferred to fresh PBS until imaging. Imaging was performed on a Zeiss Imager M2 (Carl Zeiss, Germany) using a 5x objective and images were generated using Zen software (Zen 2.3, Carl Zeiss Microscopy GmbH, Jena, Germany). A z‐stack was obtained over two channels, AlexaFluor 488 nm (AF488) for FDA and AF568 for PI. Images generated from each channel were combined in ImageJ to a maximum projection. These maximum projections were then combined to create a merged image, and cells were counted via use of a script to analyse particles (Appendix D: ImageJ Code), or manually using the cell counter function depending on the level of background staining. Viability was determined as the by proportion of living cells over total cells counted, following below formula:

𝐿𝑖𝑣𝑒 𝐶𝑒𝑙𝑙𝑠 𝑉𝑖𝑎𝑏𝑖𝑙𝑖𝑡𝑦 100% (2) 𝐿𝑖𝑣𝑒 𝐶𝑒𝑙𝑙𝑠 𝐷𝑒𝑎𝑑𝐶𝑒𝑙𝑙𝑠

24 Chapter 2: Materials and Methods

2.13 CELL MORPHOLOGY

To understand how cells would perform in PlacMA hydrogels, confocal imaging was carried out after fluorescence staining to examine cell morphology. Gels were made using a concentration of 0.5% w/v and 1% w/v, using crosslinking times of 15, 30, and 45 seconds (s) per concentration for two time points of 1 day and 6 days. For confocal imaging 2 gels from each group were fixed in 4% v/v paraformaldehyde (provided by Regenerative Medicine group, IHBI, Australia) at day 6 of culture and stained for fluorescently labelled.

Hydrogels were incubated in 5% v/v goat serum (provided by Regenerative Medicine group, IHBI, Australia) and 0.1% v/v TX in PBS for 2 hours at 4 °C to block and permeabilise. Hydrogels were then incubated with primary antibody was anti α‐Human β1 integrin antibody (ab134179; 1:500 dilution; Abcam, UK) with 1% v/v goat serum in PBS overnight at 4 °C. Samples were then washed with 1% goat serum in PBS for 30 minutes at 4 °C and then again for 4 hours at 4 °C. Hydrogels were then incubated with AF488 labelled secondary goat‐ α‐ rabbit (1:200 dilution; Thermofisher, USA) with phalloidin AF518 (1:50 dilution; provided by Regenerative Medicine group, IHBI, Australia) and DAPI (4′,6‐diamidino‐2‐phenylindol; provided by Regenerative Medicine group, IHBI, Australia) in PBS at 4 °C overnight. Samples were then incubated in 4 °C PBS for 2 hours, exchanging solution after 1 hour.

2.14 STATISTICAL ANALYSIS

Plots of data and residuals were examined to determine if the data followed a normal distribution. If there was no evidence of normal distribution, non‐parametric tests were used to determine significance, such as Kruskal‐Wallis and Multiple Mean Comparisons for post‐ hoc analysis. For parametric tests, data was analysed using ANOVA with Tukey’s Honestly Significance Difference test (Tukey HSD) for post hoc means comparison. Significance of α = 0.05 was used. Statistics collected were mean (µ), standard deviation(StdDev), variance (Var), count (n), and standard error of the mean (SEM). Results were visualised using Rstudio (1.2.1578) using the ‘ggplot2’, ‘ggsignif’, ‘ggpubr’ libraries.

Chapter 2: Materials and Methods 25

Chapter 3: Results

3.1 DECELLULARISATION EFFICACY

In order to produce PlacMA hydrogels, human CV tissue was first decellularised to purify the ECM. Decellularisation was performed using N‐lauroyl sarcosine (sarkosyl; SLS) and Triton X‐ 100 (TX) detergent‐based methods. In order to determine the efficacy of decellularization and polysaccharide retention, the DNA and sGAG content was assessed using Picogreen and DMMB assays, respectively. Visually, placental tissue displayed a slight loss in red colour between day 0 to day 2 in ddH2O, with the supernatant appearing substantially clearer. The addition of the respective decellularisation solutions (DCS; Figure 11, panel C) led to an increase in the turbidity of the both, SLS and TX, solutions. Focusing on the images for a qualitative assessment, the SLS method produced a more rapid colour change from red to white until day 9 (Figure 12, panel F). On day 9 the TX and ammonium hydroxide solution was exchanged, and the paler colouration was observed.

Two days of ddH2O water washing did not significantly (p value = 0.2) lower the DNA content (Figure 12, panel A) compared to native placental tissue (day 0). After 24 hours of washing with decellularisation solution (DCS) containing 2% w/v sodium lauroyl sarcosine (SLS) (day 3) there was no significant change (p value = 0.6232), however, a clear decreasing trend appears following this time point. The final sample showed a statistically significant decrease of 74.7 ± 19.8% in total DNA content when normalising the data to native tissue.

Samples washed with 2% v/v Triton X‐100 (TX) showed a significant increase in the reported DNA content in the first day of adding the DCS (day 3 of protocol). A clear decreasing trend appears for the TX method also. The final sample showed a significant decrease of 94.2 ± 3.1% for the TX decellularization method (Figure 12, panel B). This data demonstrated significantly better DNA removal compared to the SLS method (p value < 0.01; see Appendix A, Table 3 for data).

Chapter 3: Results 27

Figure 11: Decellularisation of placental tissue over time.

Qualitative display of human placental tissue being decellularised using N‐lauroyl sarcosine and Triton X‐100 based decellularisation solutions, respectively. For panels C‐G) Two leftmost bottles contain tissue decellularised with N‐lauroyl sarcosine, two rightmost bottles contain tissue decellularised using Triton X‐100. A) Day 0, Immediately after tissue was

immersed in ddH2O. B) End of day 2, before decellularisation solution added. C) Day 3, after 1 day of decellularisation solution added. D) Day 5. E) Day 7. F) Day 9, end of decellularisation solution. Two rightmost bottle solutions were exchanged with Triton X‐100

and ammonium hydroxide. G) Day 11, after 2 days in ddH2O. 4 bottles used as 2 donor tissues underwent 2 decellularisation solution types.

Retention of sulphated glycosaminoglycans (sGAGs) is displayed in Figure 12, panel C. The sGAG content of raw placental tissue was 20.51 ± 0.50 µg/g of dried tissue. TX derived dECM retained significantly more sGAGs (3.12 ±0.06 µg/g) than SLS derived dECM (1.68 ±0.33 µg/g). See Appendix A, Table 3 for data.

28 Chapter 3: Results

Figure 12: Decellularisation efficacy by presence of DNA. Quantitative results of decellularisation. A) Efficacy as determined by mass of DNA per mass of total dry tissue over time; the tissue is only washed in ddH2O for the first days before being immersed in decellularisation solution. B) Efficacy as determined by percentage of total DNA removed. C) Retention of sulphated glycosaminoglycans. Error bars show standard error. Each of the two samples from different biological donors with 6 technical replicates, respectively, were used for each decellularization method. n = 6 per donor, per decellularisation method. ‘**’ denotes p value < 0.01 between groups.

3.2 SOLUBILISATION OF DECM AND FUNCTIONALISATION WITH METHACRYLIC ANHYDRIDE

Following decellularization with SLS and TX, respectively, dECM was solubilised using pepsin digestion. Visually, dECM was gradually digested and the supernatant became increasingly clear and viscous. Solubilised ECM (sECM) was then functionalised by reaction with MAAh to introduce photocrosslinkable methacryloyl moieties. 1H‐NMR was performed to confirm

Chapter 3: Results 29

successful synthesis of methacryloyl‐substituted placental ECM (PlacMA). The spectra corresponding to sECM and PlacMA are shown in Figure 13 (panels A‐C, pre‐functionalised‐

Figure 13: 1H‐NMR Spectra for pre‐ and post‐functionalised sECM.

A‐B) Triton X‐100 derived sECM. C‐D) N‐lauroyl sarcosine derived sECM. A) pre‐ functionalised sECM. B) Functionalised sECM (PlacMA). C) pre‐functionalised sECM. D) Functionalised sECM (PlacMA). E) Representative map of functionalised peptide to 1H‐NMR Spectra. Image modified from [90]. (v) aromatic amino acid peaks (w) methacrylate and methacrylamide peaks (x) unknown peak (y) lysine amino acid peak (z) methyl peak.

30 Chapter 3: Results

‐sECM; panels B‐D, PlacMA), along with the corresponding illustration of chemical structure (panel E). Group w and z are present in the functionalised samples and absent in the pre‐functionalised samples. These peaks correspond to the introduced methacryloyl groups in the product PlacMA [147], [148]. Group y corresponds to the lysine amino acid and the presence has been reduced in the functionalised spectra. This reduction together with the presence of w groups is an indicator that some level of functionalisation of methacrylate groups on the protein has occurred [147]–[149].

Figure 14: Results of TNBS Assay.

Determination of the level of functionalisation for solubilised placenta. A) Concentration gradient used to determine degree of functionalisation with methacryloyl from normalised absorbance values. B) Degree of functionalisation for decellularisation type. Error bars show standard error. A‐B) For each group n =6. C) n = 4, ‘***’ denotes p value < 0.001.

The degree of amines functionalised with methacryoyl groups (DoAF; methacrylation) was quantitatively assessed using a TNBS assay (Figure 14). All samples were normalised according to the respective decellularisation method and pre‐/post‐functionalisation stage (Figure 14, panel B). In panel C, the PlacMA synthesized from SLS sECM had a DoAF of 88.2 ± 2.3%. TX‐ derived PlacMA produced a significantly higher DoAF (p value = <2 x 10‐16 ) of 93.5 ± 0.1%.

Chapter 3: Results 31

3.3 MECHANICAL TESTING

Following successful synthesis of SLS‐ and TX‐derived PlacMA, hydrogels were prepared using visible light crosslinking, and incubated overnight at 37 °C in PBS to allow for effective swelling. Hydrogels from each sample were combined to display the variability of the material as a whole. Unconfined, uniaxial compression testing was performed to explore the mechanical properties of crosslinked PlacMA hydrogels. The compressive moduli, as well as failure strain and stress were recorded to characterise properties of PlacMA hydrogels. Figure 15 shows representative stress‐strain curves of SLS and TX‐derived PlacMA hydrogels. Panels A, C, E show 0.5% w/v PlacMA hydrogels (gels), while panels B, D, F show 1% w/v gels crosslinked by exposure to 405 nm light for 15, 30, or 45 s. All stress‐strain curves displayed a J‐shaped curve, characteristic for biological materials. Linear elastic behaviour occurred between 0.1 and 0.15 mm/mm strain and the slope between these points was used to determine the compressive moduli. In this region the fibrils of the gel rearrange. In all cases, the slope of TX gels is greater, indicating stiffer gels. TX gels also undergo more strain than SLS gels before failure.

Figure 16 Panel A shows the compressive moduli (kPa) of 0.5% w/v PlacMA gels crosslinked for varying amounts of time. TX gels produced significantly stiffer gels (1.45 ± 0.29 kPa to 7.65 ± 0.48 kPa; mean (µ) ± standard error of the mean (SEM)) for all crosslinking durations compared to SLS gels (1.09 ± 0.2 kPa to 3.27 ± 0.36 kPa). As the modulus continues increasing with crosslinking time, complete crosslinking has not been achieved. Appendix 10 Table 5 contains modulus data.

Figure 16 Panel B shows the compressive moduli (kPa) of 1% w/v PlacMA gels crosslinked for varying amounts of time. TX gels produced significantly stiffer gels (20.87 ± 0.57 kPa to 44.3 ± 0.99 kPa) for all crosslinking durations compared to SLS gels (4.61 ± 0.81 kPa to 16.9 ± 2.04 kPa). As the modulus continues increasing with crosslinking time, complete crosslinking has not been achieved. Appendix 10 Table 5 contains modulus data.

Figure 16 Panel C shows the stress failure (kPa) of 0.5% w/v PlacMA gels crosslinked for varying amounts of time. There was no significant difference between, when controlling for crosslinking time, TX gels (2.72 ± 1.28 kPa to 36.91 ± 8.55 kPa) for all crosslinking durations compared to SLS gels (8.64 ± 2.28 kPa to 34.23 ± 5.72 kPa). Appendix 10 Table 7 contains stress failure data.

32 Chapter 3: Results

Figure 16 Panel D shows the stress failure (kPa) of 1% w/v PlacMA gels crosslinked for varying amounts of time. TX gels produced significantly stiffer gels (45.94 ± 6.07 kPa to 82.8 ± 22.41 kPa) at 15 seconds crosslinking duration (p value < 0.001) however, there was no significant difference for other crosslinking durations compared to SLS gels (19.84± 8.15 kPa to 37.52 ± 8.63 kPa).

Figure 15: Representative Stress/ Strain behaviour of PlacMA hydrogels.

Representative stress (kPa)/ strain (mm/mm) graphs from uniaxial, unconfined compressive testing of PlacMA hydrogels of different concentrations (% w/v) and photocrosslinking times (seconds) under 405 Nm light. Hydrogels were compressed at 0.01 mm/s until 4.5 N or 0.01 mm extension. Column A‐E) 0.5% w/v PlacMA B‐F) 1% w/v PlacMA. Row A‐B) 15 second crosslink C‐D) 30 second crosslink E‐F) 45 second crosslinking. Representative graphs taken from sample with compressive modulus closest to the group mean. Each group n = 12.

Chapter 3: Results 33

Figure 16: Mechanical Data for PlacMA Hydrogels.

Mechanical data obtained from uniaxial, unconfined compressive testing of PlacMA hydrogels of different concentrations (% w/v) and photocrosslinking times (seconds) under 405 Nm light. Hydrogels were compressed at 0.01 mm/s until 4.5 N or 0.01 mm extension. Group n= 12. Column A‐E) 0.5% w/v PlacMA Hydrogel B‐F) 1% w/v PlacMA Hydrogel. Row A‐B) Compressive modulus (kPa) between 10%‐15% strain. C‐D) Failure stress (kPa) E‐F) Failure strain (mm/mm).

34 Chapter 3: Results

Figure 16 Panel E shows the strain failure (mm/mm) of 0.5% w/v PlacMA gels crosslinked for varying amounts of time. TX gels produced significantly weaker gels (0.25 ± 0.09 mm/mm to 0.66 ± 0.05 mm/mm) at 15 seconds crosslinking duration (p value < 0.001) but not for other crosslinking durations compared to SLS gels (0.53 ± 0.1 mm/mm to 0.79 ± 0.02 mm/mm).

Figure 16 Panel F shows the failure strain (mm/mm) of 1% w/v PlacMA gels crosslinked for varying amounts of time. There was no significant difference between TX gels (0.53 ± 0.03 mm/mm to 0.65 ± 0.06 mm/mm) for all crosslinking durations compared to SLS gels (0.37 ± 0.04 mm/mm to 0.67 ± 0.06 mm/mm). Appendix 10 Table 6 contains strain failure data.

Collectively, this data demonstrates that PlacMA hydrogels mechanical properties are highly tuneable to suit a large variety of applications and cells types.

3.4 EFFECTIVE SWELLING

To determine the effective swelling, 40 µl PlacMA hydrogels were weighed immediately after crosslinking and again after swelling overnight in PBS at 37 °C. The difference in wet weights was expressed as a percentage and referred to hereafter as effective swelling (Figure 17, Panels A and B). Appendix B Table 8 contains group effective swelling data.

Figure 17: Effective swelling of PlacMA hydrogels.

Effective swelling data obtained by weighing 40µl hydrogels immediately following crosslinking and after being immersed in 37 °C PBS overnight on a shaker plate. The post‐immersion weight was divided by the post‐crosslinking weight, n= 12 . A) 0.5% w/v PlacMA Hydrogel B) 1% w/v PlacMA Hydrogel.

Chapter 3: Results 35

Figure 17 Panel A shows the effective swell rate of 0.5% w/v PlacMA gels. There was no significant difference between TX gels (‐14.65 ± 1.18 mg to 0.89 ± 2.59 mg) for all crosslinking durations compared to SLS gels (‐28.85 ± 6.73 mg to 5.44 ± 2.24 mg).

Figure 17 Panel B shows the effective swell rate of 1% w/v PlacMA gels. There was no significant difference between TX gels (‐8.44 ± 3.62 mg to ‐1.83 ± 1.52 mg) for all crosslinking durations compared to SLS gels (‐9.04 ± 1.82 mg to 4.22 ± 2.09 mg).

3.5 CELL CULTURE

To assess compatibility of SLS‐ and TX‐derived PlacMA hydrogels, human bone marrow mesenchymal stem cells (MSCs) were encapsulated in PlacMA hydrogels and cell viability was determined on day 1 and 6 of culture through FDA/PI, live/dead staining. Hydrogels were crosslinked for 15 second and 45 second durations to generate the low and high modulus gels, respectively.

Figure 18 shows fluorescent staining using FDA and PI on crosslinked hydrogels on day 6 after seeding. The inset are images taken after 1 day of seeding. Appearance of stretched cell morphologies were observed in SLS‐derived gels at day 6. Please see Appendix: Table 9 for results.

After one day of culture (panel column C‐L), there was no significant difference (p value = 0.10) in viability between TX derived PlacMA hydrogels (84.41 ± 1.44% to 95.35 ± 0.81%) and SLS derived PlacMA hydrogels (83.6 ± 1.22% to 89.75 ± 0.92%).

After day 6 of culturing there was significant difference (p value < 0.001) between the groups. TX derived PlacMA hydrogels (87.66 ± 2.84% to 92.01 ± 1.8%) supported higher viability and SLS derived PlacMA hydrogels (89.34 ± 1.24% to 92.65 ± 1.23%).

Collectively, this data demonstrates that PlacMA hydrogels can maintain a high degree of cellular viability in all conditions.

36 Chapter 3: Results

Figure 18: Cell Viability in PlacMA Hydrogels.

Cell viability data of bone marrow MSCs cultured in PlacMA hydrogels of 0.5% w/v and 1% w/v, crosslinked at 15 seconds and 45 seconds, n= 6 per group. Confocal images are representative of live/ dead stain performed. Columns A‐K) Qualitative live/dead staining performed with FDA (green) and PI (red). Scale bars = 300µm. Panel was taken on day 6 with the inset image (bottom left on each panel) taken on day 1. Column A‐ J) SLS derived PlacMA hydrogels. Column B‐K) TX derived PlacMA hydrogels. Column C‐L) Viability data of live/ dead staining obtained by counting and dividing all live cells by total number of cells counted. Row A‐C) 0.5% w/v PlacMA, crosslinked for 15 seconds. Row D‐F) 0.5% w/v PlacMA, crosslinked for 45 seconds. Row G‐I) 1% w/v PlacMA, crosslinked for 15 seconds. Row J‐L) 1% w/v PlacMA, crosslinked for 45 seconds.

3.6 CELL MORPHOLOGY

Confocal imaging was carried out to examine cell morphology at day 6 following DAPI and phalloidin staining (Figure 19). Blue (DAPI) indicates the nuclei and green (AF488 phalloidin) indicates the presence of actin. Hydrogels were crosslinked for 15 second and 45 second‐

Chapter 3: Results 37

Figure 19: Cell Morphology of Cultured PlacMA Hydrogel.

Representative confocal imaging of PlacMA hydrogels cultured for 6 days. Cultured hydrogels were stained with DAPI (blue; nuclear) and phalloidin (green; actin). Scale bars = 300µm. Columns A‐G) SLS derived PlacMA hydrogels; B‐H) TX derived PlacMA hydrogels. Section A‐D) 0.5% w/v concentration of PlacMA; E‐H) 1% w/v concentration of PlacMA. Row A‐B), E‐F) 15 seconds crosslinked; C‐D), G‐H) 45 seconds crosslinked. Arrows indicate cells expanding on surface of hydrogel. ‐durations for the low compressive modulus to promote cell expansion and migration. SLS gels generally presented more spindle‐like and stretched morphologies when compared to TX gels. Stiffer gels generally resulted in more rounded morphologies.

38 Chapter 3: Results

Figure 19 shows the confocal imaging of cells taken at the end of the cell study. Blue (DAPI) indicates the nucleus and green (phalloidin) indicates the presence of actin. SLS gels present with more spindle and stretched morphologies over TX gels. Stiffer gels present with more rounded morphologies. In panel B) and panel F) arrows indicate the surface of hydrogels that are supporting cells, which are adopting flat morphologies.

Chapter 3: Results 39

Chapter 4: Discussion

4.1 DECELLULARISATION ASSESSMENT

Decellularisation was performed over the course of this project using SLS‐based (sarkosyl; N‐ lauroyl sarcosine) and TX‐based (Triton X‐100) detergents, respectively (Figure 5). Both destabilise the cell wall, allowing cellular material removal in subsequent washings. [64], [77], [115]–[119]. SLS is an ionic detergent with a polar head region that does interferes with protein‐protein interactions [115]–[117]. TX is considered a mild detergent as it is non‐ionic and causes minimal damage to ECM proteins by primarily interfering with lipid‐lipid, and lipid‐protein interactions [18], [77], [119]. Ammonium hydroxide, added in the final day of TX decellularisation solution, is a strong alkali that interferes with the cell wall and denatures dual stranded DNA [114], [150].

Assessing the efficacy of methods using different decellularisation solutions (DCS) is useful to inform about possible immunoinflammatory effects when interacting with other biological systems and any cell‐instructive properties [6], [13], [119]. Hydrogels produced by Francis et al. (2017) [13] demonstrated that using human placenta as a biomaterial can create xeno‐ and tumor‐material‐free platforms rich in structural proteins. Human placental biomaterial was also demonstrated to possess many growth factors that upregulate regeneration, angiogenesis, stem cell recruitment and reduce scarring and inflammation in injury sites [13]. Human placenta already has demonstrated clinical use and it typically discarded after delivery, making it an excellent source for a biomaterial [121].

By visual inspection, it appears the two days of ultrapure water (ddH2O) washing removes much of the residual extracellular waste with the solution becoming less turbid (Figure 11, panel A – B). In panel C, the solutions show a higher turbidity, likely as a result of cellular debris from DCS action. The tissue washed with SLS visibly shows a deeper colouring, likely as a result of the ionic action of SLS solubilising material faster, when compared with TX [117], [119]. On day 8 the TX DCS was supplemented with 0.1% w/v of the strong base ammonium hydroxide (NH4OH). After one day this has produced a more pale colouration (Figure 11, panel F, G) and a lower average DNA content (Figure 11, A) compared to SLS.

Starting weight of DNA was 3.4 ± 0.25 µg/ 1 mg of dry tissue. The first day (day 3) of adding DCS saw an increase in the presence of DNA in the tissue being decellularised with 2% v/v TX

Chapter 4: Discussion 41

(Figure 12, panel A). This is likely due to insufficient washing of the samples, as TX is a known contaminant for Picogreen assays [151]. The vendor lists an acceptable concentration of 0.1% v/v TX increasing the signal by 7% [151]. A clear downward trend in the amount of DNA is still observed despite this. TX DCS reduced DNA content to 0.2 ± 0.03 µg/mg, while SLS DCS reduced DNA content to 0.56 ± 0.11 µg/mg. Francis et al. (2017) [13] reduced DNA content of human chorionic plate to 0.12 µg/mg using an SLS based method and a dedicated decellularisation reactor. Coronado et al. (2017) [114] reduced DNA content of porcine liver to <0.05 µg/mg using an TX/ ammonium hydroxide‐based method and a dedicated decellularisation reactor. Skårdal et al. (2015) [18] decellularised hepatic, cardiac, and skeletal muscular tissue, however, did not quantitatively assess the efficacy. Decellularisation was instead reported using the reduced presence of nuclear material via histological staining.

In a review, Crapo et al (2011) [77] explored a range of decellularisation techniques and outlined a level of decellularisation, based on adverse host reactions. Quantitatively these levels were defined as <50 ng dual stranded DNA per 1 mg of sECM dry weight, <200 base pair DNA fragment length [152], [153]. Qualitatively it was stated there should be no visible material when tissue sections when stained for nuclear material (DAPI, or hematoxylin and eosin stains). Using this guideline, additional optimisation of the decellularisation process will be required to improve the DNA removal from placental tissue.

For measuring potential cell‐instructivity, sGAGs were measured via DMMB assay (Figure 12, panel C). sGAGs are highly polar, linear, sugars that modulate growth factor binding and bioactivity within the ECM [154], [155]. sGAGs help facilitate cohesion of the ECM and attract water molecules to allow low friction surfaces and mechanical load dispersal [154], [155]. Starting weight of sGAGs in human placenta measured at 20.51 ± 0.5 µg/mg dry tissue. TX based decellularisation method had significantly higher retention at 3.12 ± 0.06 µg/mg, while the SLS based method retained 1.68 ± 0.33 µg/mg. Francis et al. (2017) [13] performed an ELISA directly and found the presence of many growth factors for cell instructivity such as; VEGF‐A,‐B, HGF, bFGF, EGF, PDGFs, IGF‐II and Human Placental Lactogen. It was also demonstrated the placenta was rich in collagen‐I,‐IV, laminin γ‐1 and fibronectin. Skårdal et al. (2015) [18] also used an ELISA to demonstrate the presence of growth factors the persisted through decellularisation such as; bFGF, BDNF, BMP‐5,‐7, FGF‐4,‐7, HB‐EGF, HGF. Coronado et al. (2017) [114] reported 10.35 mg/mg sGAGs from porcine liver. While growth factors were not tested for in this project, these papers demonstrate that growth factors are retained through these respective decellularisation methods.

42 Chapter 4: Discussion 43

There were two papers that formed the basis of the decellularisation methods used in this project. The first is Skårdal et al. (2015) [18] that used a TX based method on liver, skeletal and cardiac muscle for 8 days but did not publish any decellularisation data. Francis et al. (2017) [13] used a SLS based method on human placental tissue for 3 days and removed 99 % of DNA. Figure 12, panel B shows the TX (94.2 ± 3.1 %) removing significantly more than SLS (74.7 ± 19.8 %). It should also be noted that decellularisation carried out in this project was performed in 1 L glass bottles on a shaker plate. There was no dedicated reactor able to deliver vigorous agitation and thus decellularisation took longer than previously published. Having a dedicated reactor or a perfusion pump for tissue decellularisation should yield higher decellularisation rates [77], [119].

Assessing the effects of the different decellularisation detergents, it was found that the TX‐ based method yielded significantly greater removal of DNA and with less variation over the SLS‐based method. The TX‐based method also yielded greater retention of sGAGs over the SLS‐based method. For these reasons, the TX based method is considered superior in terms of decellularisation for human placenta tissue for purposes of collecting dECM.

For any studies that look to further optimise decellularisation using TX, should be performed using more optimised equipment to produce more vigorous agitation and lower decellularisation time. In addition, the samples removed for decellularisation analysis should also undergo a separate ddH2O washing over several days to remove all residual TX.

4.2 FUNCTIONALISATION ASSESSMENT

Functionalisation was performed to offer control over the mechanical properties of the hydrogel. By changing the compressive modulus of a hydrogel, the platform can better accommodate the physicochemical needs of a given cell culture [27]. The dECM was functionalised using MAAh using a method based on well‐established protocols, below 4 °C, keeping the pH above 8 using 5 M NaOH for 5 hours [16], [17]. The pH was adjusted throughout functionalisation as the isoelectric point (pI) of lysine is 9.74, with pKa values of the lysine side chain 10.67 [147]. Functionalisation was performed at 4 °C to prevent gelation of sECM at physiological temperatures.

Characterisation was performed to better understand the functionalised product. This was performed by TNBS assay and proton nuclear magnetic resonance spectroscopy (1H‐NMR). TNBS binds to primary amines, such as the side group on an unfunctionalised lysine sidechain,

Chapter 4: Discussion 43

and optimally absorbs light at 335 nm [156]. The drop in value indicates functionalisation has occurred, while measuring the gradient of a curve allows the determination of DoAF [16].

Examining the 1H‐NMR spectra (Figure 13 panel E), no significant differences between SLS‐ or TX‐ based methods were apparent. Functionalisation can be confirmed by presence of methacrylate and methacrylamide peaks (w, z) and decrease in lysine peaks (y) as the side chain is the most reactive [149]. These peaks are indicative as described by the existing literature [70], [147], [149]. sECM samples (Figure 13 panel A, C) also show many peaks not present in functionalised samples (Figure 13 panel B, D). These are likely to be small molecular species (MW <2000 Da) of peptides and amino acids. As there is no dialysis between digestion and functionalisation, these species are present here. However, dialysis following functionalisation may lead to loss of low molecular weight polypeptides, explaining why they are no longer present in the functionalised sample.

The TNBS assay (Figure 14) on PlacMA show TX derived PlacMA (93.5 ± 0.1%) have a significantly higher DoAF compared to SLS derived PlacMA (88.2 ± 2.3%), when functionalised with 0.6mg MAAh /g sECM for 5 hours. Annabi et al. (2013) [17] functionalised recombinant tropoelastin (10% w/v in PBS) for 12 hours using 8% v/v, 15% v/v, 20% v/v and reported DoAF levels of 31%, 44%, and 48% respectively at 4 °C. GelMA is typically functionalised at 50 °C with a high functionalisation (0.6 g MAAh/g gelatin) with 75% DoF, or low functionalisation (0.06 g MAAh/g gelatin) with 31% DoF [16]. This demonstrates that placental sECM can be highly functionalised under cold conditions.

Further work on this project should include adjusting the functionalisation levels. High DoAF correlates with increase in crosslinking density, stiffer gels, and reduced pore size [16], [17], [157]. This can be performed through the variation of the MAAh/sECM ratio, pH, and length of reaction time [16], [17], [147]. Physicochemical needs of cells must be met to effectively culture and study them [27], [158]. Softer gels are ideal for studying angiogenic cells and other soft tissues [27]. A low‐functionalised PlacMA, being derived from highly vascular tissue, with a low compressive modulus could prove beneficial for studying formation of blood vessels.

44 Chapter 4: Discussion 45

4.3 MECHANICAL ASSESSMENT

Preliminary testing with PlacMA found that a stock solution could only be made at 2% v/v with 1x PBS. 5% v/v and 10% v/v solutions would not solubilise, and 2% v/v was still very viscous and could only be reliably handled using positive displacement pipettes.

Experimenting with crosslinking times first started with 3, 5 and 10 minutes however the gels produced were too brittle to be removed from the casting surface. It was also found that TX derived PlacMA hydrogels could be cast starting at 3 second crosslinking duration but could not be replicated for SLS derived PlacMA. These potential hydrogels could not reliably form a sufficient crosslinking density due to the lower DoAF in the short amount of time and would either fail to cast, or dissolve entirely when in 37 °C PBS overnight. The 15 second crosslinking duration time represents the first stable, workable and reproducible time for all tested concentrations of PlacMA using a 40 µl to form a hydrogel. (Figure 10, panel A). Hydrogels from each donor were cast and then tested as per decellularisation method to investigate variability across donors. Figures for mechanical testing can be seen in Appendix B.

Compressive Moduli of PlacMA Hydrogels

Hydrogel compressive moduli were determined from mechanical data and determination (Figure 10, panel B, C, D) of each gel’s surface area and height. Panel D, (d1) show the ‘toe’ of the curve, as indicated by a substantial increase in the slope of the stress‐strain curve in this region [159]. This region has high variability as the fibrils rearrange and twist in response to an applied load [159]. Reliable determination of gel height was therefore difficult and best accomplished through mechanical data examination, as the high‐water content of gels gives false readings when using digital callipers. Attempts to automate gel height discovery have not yet been successful due to relatively high variation across samples for this small area.

Examining the mechanical testing for the compressive moduli (Figure 16, panel A, B) increasing PlacMA concentration and crosslinking time produced stiffer (higher modulus) gels on average. This is due to the density of crosslinked networks increasing and providing greater resistance to deformation [1], [16], [157]. TX PlacMA produces stiffer gels compared to SLS gels. This is likely due to a combination of the DCS action and functionalisation level of SLS derived PlacMA being lower. The action of SLS damaging proteins through conformational changes to linearise proteins, while TX functions mainly by disruption lipid‐ lipid and lipid‐protein interactions [115], [117], [118].

Chapter 4: Discussion 45

PlacMA shows a wide range of possible compressive moduli. The softest gels, 0.5% SLS PlacMA crosslinked for 15 s shows 1.09 ± 0.68 kPa, while the stiffest, 1% TX PlacMA crosslinked for 45 s shows 44.30 ± 2.80 kPa. The softest TX gel is 2.39 ± 0.70 kPa, and the stiffest SLS gel is 16.90 ± 7.36 kPa. The increasing moduli with increasing times indicate that complete crosslinking is not achieved over this range of exposure times. This range can be compared to unmodified gelMA which has been demonstrated to have a compressive moduli range of 5‐180 kPa[70], [79], [160] and MeTro which has demonstrated a compressive moduli range of 8.8 to 159 kPa [17]. This range of compressive moduli makes PlacMA suitable for a range of soft tissue cell culture such as; non‐coronary vessels, breast tissue, nerve, adipose, soft muscle and cartilage, and MSC differentiation studies [27]. The increase in hydrogel stiffness with concentration, crosslink duration, and DoAF is consistent with literature [16], [17]. The low standard error of the mean shows little donor‐to‐donor variability.

Mechanical Failure of PlacMA hydrogels

Mechanical properties at failure (Figure 16, panel C, D, E, F) were obtained by allowing testing to complete at the prescribed end rate. Failure was automatically detected and confirmed by manual inspection, of a negative gradient >0.01 kPa in the strain curve.

Failure stress of PlacMA hydrogels ranges from 2.72 ± 1.28 kPa to 82.8 ± 22.41 kPa while failure strain ranges from 0.25 ± 0.09 mm/mm to 0.79 ± 0.02 mm/mm. Meaningful statistical relationships of the mechanical failure data (stress and strain) were difficult to determine. This was due to the high variation within each sample group and no clear correlation between crosslink time, failure data, and DoAF. The cause of this variation is likely due to the non‐ optimised methodology for producing PlacMA and the biological nature of the source material also [14], [50]. For example, the extra time required to decellularise the placenta (11 days) when compared to other methods such as Skardal et al. (2015; 8 days) [18] and Francis et al. (2017; 3 days) [13]. MeTro fail stress was demonstrated to range from 12.5 to 39.3 kPa [17].

For cell cultures requiring very soft gels, SLS would be recommended as TX produced the stiffer gels for all tested concentrations and crosslinking times. Another option could be low functionalised TX gels, which were not explored here. This could be ideal for low compressive moduli (<1 kPa) cultures, such as they are required for endothelial cells [27].

More data from additional samples of placenta and an optimised decellularisation method are required to fully understand the physical properties of PlacMA. Additional data and

46 Chapter 4: Discussion 47 optimisation may allow for a parametric model to be constructed. This would afford an end‐ user who desires specific mechanical properties, to know what concentration and crosslinking time to achieve those properties.

Automation of mechanical results is desirable as this is a time consuming and non‐precise practise. It is often performed ‘by eye’ of the operator. Combining successful automation of results with a script to calculate surface area would allow a user to order all the data (mechanical and stereomicroscopy) together and have a program return results, along with settings used. This would allow consistent recreation of results from any user.

Effective Swelling

Characterisation of swelling in a hydrogel network is important as it affects mechanical and biological properties and influences the final dimensions of hydrogel constructs [1], [7]. Gel swelling degree is a result from the interplay of crosslinking degree, polymer concentration, crosslink duration, polymer‐solvent interactions and pore sizes created by the network formation [161], [162]. This information is important for hydrogel drug delivery release rates, in vivo space filling properties for injury‐site integrity, and platform stability when being studied for TERM [7], [50], [157].

DoAF affects crosslinking density, which in turn affects pore size and the swelling characteristics of a hydrogel [17]. TX derived PlacMA is has a significantly higher DoAF over SLS derived PlacMA but does not significantly swell less.

4.4 CELL VIABILITY AND MORPHOLOGICAL ASSESSMENT

Fluorescence staining was performed to assess viability of cells cultured in PlacMA hydrogels. By counting the number of live and total cell counts, viability was expressed as a percentage. FDA is a substrate that is taken in by live cells and is cleaved to a fluorescent product, indicating a cell is metabolically active and has an intact membrane [163]. PI is a DNA dye that is non‐permanent to living cells [164]. Using these together allows creating a set of images containing living cells, and another set containing dead cells. These images have then been combined (Figure 18). Cell counts were taken and used to assess the viability of each gel, the summarised data is found in Table 9: Viability Data and is represented in Figure 18.

Viability of MSC cultures after 6 days in PlacMA hydrogels was very good across all conditions, ranging from 83.6 ±1.22% to 95.35 ± 0.81%, with TX derived PlacMA hydrogels demonstrating higher viability over SLS derived PlacMA hydrogels. MeTro hydrogels demonstrated excellent

Chapter 4: Discussion 47

(>92%) viability over 7 days for 3T3 fibroblasts and HUVEC cell types. GelMA, reinforced with hyaluronic acid had average viability (<75%) 1 day after encapsulation and dropped for <60% after 28 days [70]. Francis et al. (2017) [13] reported significant increase in metabolic activity over the control. High viability demonstrated in the cell culture makes PlacMA hydrogels an excellent choice for cell studies.

Confocal fluorescence staining is performed to assess the morphology of the cell (Figure 19). DAPI is a nucleic acid stain that displays the presence of eukaryote nuclei[165] while Phalloidin is an F‐actin stain that displays the cytoskeleton allowing morphological assessment [166]. Channels for each stain were formed into a maximum projection (as with fluorescence staining) and then combined to a single image using ImageJ.

Morphology of encapsulated cells corresponds to stiffness of the gel. Despite the loss in viability, SLS gels displayed more spread, networked morphologies as they were able to move and remodel the gel with less resistance [167], [168] (Figure 19, row A‐G). TX gels, stiffer than the SLS counterparts, provided more resistance to the MSCs, keeping them in a rounded shape longer (Figure 19, row B‐H). The flat morphologies as indicated by arrows in panel B) and F) show that cells also grow on the surface of the hydrogel.

High viability and the spread morphologies of cells demonstrated PlacMA as a cyto‐ compatible and highly tuneable 3D cell culture platform.

4.5 FURTHER WORK

This section highlights the future work that should be carried out to improve characterisation of PlacMA. This work was beyond the scope of this project which primarily focused on establishing methods for decellularisation, solubilisation, and functionalisation with MAAh of placenta and other soft human tissues.

 Proteomics should be carried out for high fidelity characterisation of PlacMA at the molecular level. This would inform the content of the biopolymer, such as structural proteins and growth factors, and functionalisation of polypeptides present in PlacMA.  Performing an ELISA (enzyme‐linked immunosorbent assay) could also inform the presence of growth factors and their potential activity.  Rheology would characterise viscoelastic‐flow behaviour. This would allow determination of PlacMA as a viable source for bioprinting.

48 Chapter 4: Discussion 49

 Electron microscopy could be employed to investigate the morphology of the PlacMA hydrogel matrix as demonstrated by Francis et al., (2017) [13] among other papers.  Endotoxin and mycoplasma testing should be carried out to assess risk associated with this biomaterial in vivo.  Comparative performance of PlacMA should be studied against other popular polymers, such as gelMA for semi‐synthetic hydrogels, gelatin or placental biopolymers for biological hydrogels, or polycaprolactone for synthetic hydrogels.  Suitability for wound healing could be explored as Francis et al., (2017) [13] used a catheter to inject placental hydrogel into a cardiac injury in mice models and this resulted in reduced scarring.  Degradation studies should be performed to determine how the hydrogel will perform in vivo, as outlined by Oryan et al., (2018) [88], Yoon et al., (2016) [71] and others. Degradation is assessed as the loss of hydrogel mass over time.  Longer exposure times for photocrosslinking should be carried out to determine when complete crosslinking is achieved, and the mechanical properties of the hydrogel at that stage. Weighing a completely crosslinked gel wet, compared to the dehydrated weight would provide additional characterisation data on the sol‐gel fraction value [169], [170].

Method Optimisation

Methods used in this project could be used to explore other tissue, such as functionalising sECM derived from umbilical cord, a source of stem cells and growth factors, and amniotic membrane which already has demonstrated clinical use such as, for burns and ophthalmology [76], [121], [140]. A placental‐protein‐expression difference from mothers experiencing gestational diabetes might allow for a crosslinkable hydrogel tailored for studying insulin resistance [171].

To create softer gels, which are ideal for studying differentiation and vascularisation, changes to the functionalisation method, or an addition to the casting could be included[27][16]. The ratio of MAAh to dECM could be lowered from 0.6 g/g to 0.06 g/g and/ or less reaction time could be allowed, for example reducing from 5 hours to 1 hour[16].

A common step in producing sECM hydrogels is comminution of dECM prior to digestion through mesh size 20 (particle size 0.841 mm) but was not performed here [6], [8]–[10]. Brown et al (201) [172] explored the effect of particle size of adipose dECM in composite semi‐synthetic hydrogels. The decellularised tissue was exposed to either a bead‐beater or

Chapter 4: Discussion 49

cryomill, with the cryomill producing the smallest particle sizes. It was found that smaller sizes improved cell attachment, proliferation, and differentiation of adipose‐derived stem cells cultured in the hydrogel. Use of a cryo‐mill on PlacMA could further improve the cell culture potential of PlacMA.

50 Chapter 4: Discussion

Chapter 5: Conclusion

This research project sought to explore the effects of different detergent‐based decellularisation methods on novel, semi‐synthetic PlacMA hydrogels. This was achieved through three aims;

1. Semi‐synthetic photocrosslinkable hydrogels were created from human placental tissue using different decellularisation methods where it was found TX derived PlacMA removed more DNA and retained more sGAGs

2. Physicochemical properties of resulting PlacMA hydrogels were characterised where it was found that the hydrogels had a range of tuneable compressive moduli suitable for soft tissue culture. TX derived PlacMA were demonstrated to have higher DoAF and produce stiffer gels.

3. Viability of cells cultured in PlacMA hydrogels were assessed where it was found SLS based PlacMA promoted faster cell morphological changes but TX derived PlacMA demonstrated higher cell viability.

In conclusion, this work demonstrated that TX‐based decellularisation is the most promising for producing for PlacMA. This method removed more DNA, retained more sGAGs, supported higher cell viability, and produced stiffer hydrogels but this could be amended by lowering the DoAF. During the work performed it was discovered PlacMA could be suitable for further study for TERM research as a physicochemically tuneable, cytocompatible hydrogel.

Chapter 5: Conclusion 51

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Appendices

Appendix A Decellularisation Data

Please note: SEM, Standard Error of the Mean; StdDev, Standard Deviation; Var, Variance

Table 3: DNA Content, Before/ After decellularisation. DNA/ Tissue, ng/mg Time Sample Method (days) Average SEM StdDev Var n Labelled D109 ‐ 0 3.62 0.49 1.19 1.42 6 DM ‐ 0 3.18 0.11 0.27 0.07 6 D109 7 11 0.50 0.16 0.40 0.16 6 Sample 1 D109 6 11 0.25 0.05 0.13 0.02 6 Sample 2 DM 6 11 0.62 0.15 0.36 0.13 6 Sample 4 DM 7 11 0.14 0.01 0.03 0.00 6 Sample 3

By Decellularisation Type Time Sample Method (days) Average SEM StdDev Var n Native ‐ 0 3.40 0.25 0.86 0.73 12 TX 6 11 0.20 0.03 0.11 0.01 12 SLS 7 11 0.56 0.11 0.37 0.13 12

Table 4: sGAG Content, Before/ After decellularisation. sGAG/ Tissue, µg/mg Time Sample Method (days) Average SEM StdDev Var n D109 ‐ 0 21.81 0.55 1.35 1.83 6 DM ‐ 0 19.21 0.36 0.88 0.77 6 D109 7 11 1.39 0.50 1.23 1.51 6 D109 6 11 3.16 0.09 0.23 0.05 6 DM 6 11 3.07 0.08 0.20 0.04 6 DM 7 11 2.03 0.43 0.96 0.92 5

By Decellularisation Type Time Sample Method (days) Average SEM StdDev Var n Native ‐ 0 20.51 0.50 1.74 3.02 12 TX 6 11 3.12 0.06 0.21 0.04 12 SLS 7 11 1.68 0.33 1.11 1.23 11

Appendices 67

Appendix B Mechanical Data

Table 5: Compressive Modulus of PlacMA hydrogels (kPa)

Compressive Modulus (kPa) D/C PlacMA Crosslink time Average Mthd. % (s) (kPa) Std.Dev Std.Er Variance n SLS 0.5 15 s 1.09 0.68 0.20 0.46 12.00 SLS 0.5 30 s 2.39 0.70 0.20 0.49 12.00 SLS 0.5 45 s 3.27 1.25 0.36 1.55 12.00 TX 0.5 15 s 1.45 1.01 0.29 1.02 12.00 TX 0.5 30 s 5.78 1.78 0.51 3.18 12.00 TX 0.5 45 s 7.65 1.59 0.48 2.53 11.00 SLS 1.0 15 s 4.61 2.92 0.81 8.52 13.00 SLS 1.0 30 s 8.31 2.54 0.73 6.46 12.00 SLS 1.0 45 s 16.90 7.36 2.04 54.18 13.00 TX 1.0 15 s 20.87 1.79 0.57 3.19 10.00 TX 1.0 30 s 33.83 6.85 2.07 46.98 11.00 TX 1.0 45 s 44.30 2.80 0.99 7.85 8.00

Table 6: Strain rate (%) of PlacMA hydrogels at failure

Strain at Failure (%) D/C PlacMA Crosslink time Average Mthd. % (s) (%) Std.Dev Std.Er Variance n SLS 0.5 15 s 0.68 0.14 0.04 0.02 12.00 SLS 0.5 30 s 0.79 0.08 0.02 0.01 11.00 SLS 0.5 45 s 0.53 0.32 0.10 0.10 11.00 TX 0.5 15 s 0.25 0.30 0.09 0.09 11.00 TX 0.5 30 s 0.66 0.17 0.05 0.03 11.00 TX 0.5 45 s 0.58 0.23 0.07 0.05 10.00 SLS 1.0 15 s 0.55 0.16 0.04 0.02 13.00 SLS 1.0 30 s 0.67 0.18 0.06 0.03 10.00 SLS 1.0 45 s 0.37 0.14 0.04 0.02 12.00 TX 1.0 15 s 0.65 0.18 0.06 0.03 9.00 TX 1.0 30 s 0.55 0.14 0.04 0.02 10.00 TX 1.0 45 s 0.53 0.07 0.03 0.01 6.00

68 Appendices 69

Table 7: Stress rate (kPa) of PlacMA hydrogels at failure

Stress at Failure (kPa) D/C PlacMA Crosslink time Average Mthd. % (s) (kPa) Std.Dev Std.Er Variance n SLS 0.5 15 s 8.64 7.90 2.28 62.35 12.00 SLS 0.5 30 s 34.23 18.96 5.72 359.57 11.00 SLS 0.5 45 s 15.91 20.54 6.19 422.09 11.00 TX 0.5 15 s 2.72 4.25 1.28 18.10 11.00 TX 0.5 30 s 36.91 28.37 8.55 804.89 11.00 TX 0.5 45 s 34.78 42.64 13.48 1818.19 10.00 SLS 1.0 15 s 24.24 26.09 7.23 680.47 13.00 SLS 1.0 30 s 37.52 27.28 8.63 744.29 10.00 SLS 1.0 45 s 19.84 28.23 8.15 797.06 12.00 TX 1.0 15 s 82.80 67.23 22.41 4520.29 9.00 TX 1.0 30 s 60.12 51.81 16.38 2684.22 10.00 TX 1.0 45 s 45.94 14.87 6.07 221.16 6.00

Table 8: Relative Swelling Change (mg) in Mass in PlacMA Hydrogel MassChange (Relative)

D/C Mthd. PlacMA % Crosslink time (s) Average (mg) Std.Dev Std.Er Variance n SLS 0.5 15 s ‐28.85 24.26 6.73 588.56 13.00 SLS 0.5 30 s 5.44 7.42 2.24 55.06 11.00 SLS 0.5 45 s ‐12.48 4.28 1.29 18.32 11.00 TX 0.5 15 s ‐14.65 3.92 1.18 15.39 11.00 TX 0.5 30 s 0.89 8.58 2.59 73.69 11.00 TX 0.5 45 s ‐8.96 2.62 0.83 6.87 10.00 SLS 1.0 15 s ‐3.83 13.13 3.64 172.29 13.00 SLS 1.0 30 s 4.22 6.93 2.09 48.02 11.00 SLS 1.0 45 s ‐9.04 6.55 1.82 42.84 13.00 TX 1.0 15 s ‐1.87 5.75 1.92 33.11 9.00 TX 1.0 30 s ‐1.83 4.82 1.52 23.25 10.00 TX 1.0 45 s ‐8.44 9.57 3.62 91.56 7.00

Appendices 69

Appendix C Cellular Viability Data

Table 9: Viability Data as Live/Total cell count (%) Viability DCM PlacMA Crosslink Mean Mthd (%) (s) Day (%) StdEr StdDev Var n TX 0.5 45 1 84.41 1.44 3.54 12.52 6 TX 0.5 15 1 90.94 0.97 2.36 5.59 6 TX 1 45 1 87.32 1.36 3.34 11.16 6 TX 1 15 1 92.01 1.80 4.41 19.42 6 SLS 0.5 45 1 89.35 1.94 4.76 22.63 6 SLS 0.5 15 1 91.28 1.60 3.92 15.38 6 SLS 1 45 1 89.34 1.24 3.05 9.30 6 SLS 1 15 1 92.65 1.23 3.01 9.03 6 TX 0.5 45 6 93.11 1.38 3.39 11.47 6 TX 0.5 15 6 90.59 2.09 5.11 26.09 6 TX 1 45 6 87.66 2.84 6.97 48.51 6 TX 1 15 6 95.35 0.81 1.98 3.94 6 SLS 0.5 45 6 83.60 1.22 3.00 8.98 6 SLS 0.5 15 6 86.43 1.21 2.97 8.80 6 SLS 1 45 6 86.62 1.49 3.66 13.38 6 SLS 1 15 6 89.75 0.92 2.26 5.12 6

70 Appendices 71

Appendix D ImageJ Code

Batch Max Projection

inputDir = getDirectory("Choose a Source Directory");

outputDir = getDirectory("Choose an Output Directory"); ;

setBatchMode(true);

list = getFileList(inputDir);

for (i = 0; i < list.length; i++) {

path = inputDir+list[i];

run("Image Sequence...", "open=[path] number=1000 starting=1 increment=1 scale=100 file=c1 sort");

run("Z Project...", "start=1 stop="+nSlices+" projection=[Max Intensity]");

name=getTitle;

outputAppendix = "_dead";

saveAs("Tiff", outputDir+name+outputAppendix);

close();

close();

run("Image Sequence...", "open=[path] number=1000 starting=1 increment=1 scale=100 file=c2 sort");

run("Z Project...", "start=1 stop="+nSlices+" projection=[Max Intensity]");

name=getTitle;

outputAppendix = "_live";

saveAs("Tiff", outputDir+name+outputAppendix);

close();

close();

}

setBatchMode(false);

Particle Count

run("Subtract Background...", "rolling=50");

setOption("BlackBackground", false);

run("Make Binary");

run("Watershed");

run("Analyze Particles...", "pixel show=Outlines display summarize");

Appendices 71

Annex 1

This Annex was is based upon the research performed in 2018 while at Hubrecht Institute, Universitair Medisch Centrum Utrecht, Nederland which was been previously submitted to fulfil the requirements of the Master of Biofabrication at Universitair Medisch Centrum Utrecht.

72 Appendices 73

Chapter 1, version A

David Pershouse, BSc, BMath

Supervisors; Susana Piluso PhD, Petra deGraaf PhD

Major Research Project, Biofabrication Masters

March 2019

Appendices 73

Keywords Biofabrication, Biosignalling, Cell Instructivity, Corpus Spongiosum, Crosslink, Decellularised ECM, Decellularized ECM, Electrospin, Gelatin, Hypospadias, Membrane, Neo‐ Urethra, Non‐woven Fibre, Placenta, Polycaprolactone, Regenerative Medicine, Silk Fibroin, Solution Electrospinning, Tissue Engineering, Urethra, Urethral Reconstruction, Urethral Stricture, Urethroplasty

74 Appendices 75

1. Abstract Urethral stricture and hypospadias (congenital malformation) patients undergo urethral reconstruction (urethroplasty) utilising grafts that replace the epithelial layer only, not the supportive surrounding tissue. Electrospun membranes based on decellularised extracellular matrices (dECM) represent a promising source for creation of cell‐instructive platform for tissue engineering and regenerative medicine. Biochemical signalling molecules retained by dECM spur cellular proliferation, differentiation, and function for specific urogenital tissue tracts. However, electrospun membranes generated solely from dECM are quickly degraded and mechanically weak. Here, the morphology and mechanical properties of placental dECM is explored through combination with silk fibroin, PCL, and gelatin and electrospun in single‐ and multi‐layer configurations to produce membranes. Generated membranes are potentially cell‐instructive and physicochemically‐robust constructs for research into urethral repair. Such membranes could be rolled to create a urethral lumen while also integrating and supporting into the surrounding tissue.

Appendices 75

5.1 1.1 LAYMAN’S SUMMARY

Hypospadias describes the urethral opening incorrectly forming on the underside of the penis and is one of the most common genital birth defects in males. Treatment involves creating a new length of urethra to a new opening at the correct location. Treatment can leave scar tissue that narrows the urethra, described as urethral stricture. To fix stricture, the scar tissue is removed, and the gap replaced with a small slice of the patients own tissue, usually from the mouth. Scar tissue can be form for many different reasons, such as infection and even the surgery to fix the original stricture.

Tissue engineering and regenerative medicine are two areas of research that investigate how new materials and technologies can be combined to provide better treatments for disease. These fields are important for any transplant‐related procedure as distance, donor compatibility, and donor‐site morbidity can cause problems or prevent treatment to a patient. This project looks at some of these materials and technologies and how they could be used to create a new membrane for treating hypospadias and stricture. The method used was electrospinning, as it can reproduce membranes pf ultra‐thin, non‐ woven fibres.

The membranes created for this project were electrospun using combined decellularised tissue, gelatin, polycaprolactone (PCL) and silk fibroin. This could allow the membranes to be bio‐compatible, strong, also to promote healing and cell growth. How the fibres looked and their resistance to force was measured.

76 Appendices 77

Table of Contents

1. ABSTRACT ...... 75

1.1 Layman’s Summary ...... 76

2. INTRODUCTION ...... 83

3. MATERIALS AND METHODS ...... 90

3.1 Tissue Isolation ...... 90

3.1.1 Urethra Isolation, Decellularisation, and Digestion ...... 90

3.1.2 Placental Isolation and Decellularisation, and Digestion ...... 91

3.2 Silk Fibroin Extraction ...... 91

3.3 Single Layer Electrospinning ...... 92

3.4 PCL Reinforced Multiple Layer Electrospinning ...... 92

3.5 Crosslinking ...... 93

3.6 Scanning Electron Microscopy ...... 9 4

3.7 Mechanical Testing ...... 94

3.8 Analysis ...... 94

4. RESULTS AND DISCUSSION ...... 94

4.1 Decellularisation ...... 94

4.2 Electrospinning ...... 96

4.2.1 Early Electrospinning Attempts ...... 99

4.2.2 Electrospinning with Formic Acid ...... 101

4.2.3 Electrospinning with Silk Fibroin ...... 104

4.2.4 Electrospinning with dECM ...... 108

4.2.5 Effects of Electrospin Settings on dECM ...... 110

4.2.6 Multilayered Electrospinning ...... 114

4.3 Stabilising Gelatin via Crosslinking ...... 117

4.4 Mechanical Testing of Electrospun Membranes ...... 120

5. CONCLUSIONS AND FUTURE WORK ...... 123

6. REFERENCES ...... 124

Appendices 77

78 Appendices 79

List of Tables

Figure 1: Hypospadias Regions...... 83

Figure 2: Anatomy of Stricture ...... 84

Figure 3: Anatomy of the Human Penis ...... 85

Figure 4: Human Urethra and CS, Verhoff Strain...... 86

Figure 5: Electrospin Setup...... 87

Figure 6: Equine Urethra and CS, Trichrome stain ...... 88

Figure 7: Decellularisation ...... 95

Figure 8: Morphology of initial membranes ...... 99

Figure 9: Fibre Diameter of Initial Membranes...... 100

Figure 10: Membranes generated using formic acid as a solution...... 102

Figure 11: Fibre diameter of membranes created from polymers solubilised in formic acid ...... 103

Figure 12:Membranes generated using silk fibroin in formic acid ...... 105

Figure 13: Fibre diameter of membranes created from fibroin polymers solubilised in formic acid ...... 106

Figure 14: Fibre Morphology of gelatin and gelatin/dECM blend membranes ...... 108

Figure 15: Fibre diameter of gelatin and gelatin/ dECM at 15% and 35% w/v...... 109

Figure 16: Effect of different electrospinning parameters...... 111

Figure 17: Effect of different spinning parameters on fibre diameter ...... 112

Figure 18: Mid‐spin feed‐rate Issues...... 113

Figure 19: Effect of multilayering with PCL...... 115

Figure 20: Fibre diameter for membranes with‐ and without membrane reinforcement...... 116

Figure 21: Cross section of single layer and reinforced membrane...... 118

Figure 22: Effect of different crosslinking methods...... 119

Figure 23: Effect of crosslinking, cross section...... 120

Figure 24: Elastic modulus of non‐ and reinforced membranes...... 122

Appendices 79

List of Abbreviations

Abbreviation Full Name Abbreviation Full Name

AcOH Acetic Acid PBS Phosphate Buffered Saline

CaCl2 Calcium Chloride PCL Polycaprolactone

CC Corpus Cavernosum RM Regenerative Medicine

CS Corpus Spongiosum SDS Sodium Dodecyl Sulphate dECM Decellularised ECM SEM Scanning Electron Microscopy dH2O Distilled Water sGAG Sulfated Glycosaminoglycans

DMF Dimethylformamide TE Tissue Engineering

DNA Deoxyribonucleic Acid TX100 Triton X100

ECM Extracellular Matrix UMC Universitair Medisch Centrum

ES Electrospinning v/v Volume Per Volume

EtOH Ethanol w/v Weight Per Volume

FDA Food and Drug Administration WKZ Wilhemina Kinderziekenhuis

GLT Glutaraldehyde

H Humidty

HCL Hydrochloric Acid

HFIP Hexafluoroisopropanol kDa Kilodalton kV Kilovolts

LD50 Lethal Dose, 50 ml/h Millilitres Per Hour

Mpa Megapascal mQH2O Milli‐Q Ultra‐Pure Water

MtOH Formic Acid

80 Appendices 81

Na2CO3 Sodium Carbonate

NH4OH Ammonium Hydroxide

Statement of Original Authorship

The work contained in this thesis has not been previously submitted to meet requirements for an award at this or any other higher education institution. To the best of my knowledge and belief, the thesis contains no material previously published or written by another person except where due reference is made.

Signature: ______

Date: ______

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Acknowledgements

This project could not have been completed without the help of Nicoline Korthagen, PhD, Mr. Lois van de Boom, Bart Leeman, PhD, and Ms. Nikola du Preez of the Department of Equine Sciences, UMC (Nederlands) for provision of equine tissue. Kristine Janssen, MD of Wilhelmina Kinderziekenhuis, for provision of placenta. Marco Viveen, PhD of the UMC Microbiology lab for use of SEM. Mies van Steenberger, Department of Pharmacology UMC for use of the mechanical tester. Merle, PhD for use of the tissue homogeniser.

Thank you to Miguel Castillo, PhD, Paul, Chris, Quentin Pfiffer, and Margot general advice. Joanne Richardson QUT, Roos Nieuwenhuis and Jacqueline Alblas, PhD UMC for their near‐miraculous ability to handle the red tape involved in the international biofabrication program. As a general thank you for continued support, fellow biofabrication and regenerative medicine masters students and tau‐tau Bilstraaters, Franzi and Novella. With a special thank you to Sophie for saving me from a London coach terminal and Samira for the many adventures.

With deep apologies to my long‐suffering supervisors Petra and Susana whom, if I took their advice, would’ve submitted on time.

82 Appendices 83

2. Introduction This introduction outlines urethral disease, current treatment, and issues within that treatment. Paths for research into a new treatment are then presented, followed by a brief discussion on the urethra itself to understand the requirements for these treatments to meet. New technologies, materials, and research fields are then presented and their relevance to research for treatments. This section ends by proposing a hypothesis and the aims to support the proposal.

Hypospadias is a congenital deformation of the urethral meatus and typically requires surgical correction [1‐5]. Rather than occurring at the distal point of the penis, the urethra opens on the underside of the penis and is further classified according to which region the urethra opens (see Figure 20)[1]. Occurring in approximately 5 in 1000 USA male births and increasing each year, hypospadias is one of the most common birth defects and frequently occurs with penile curvature, an incomplete foreskin, with a narrower and more rigid urethra [3‐6]. The exact cause is unknown, however it’s a multifactorial disease phenotype resulting from the genital plate not closing correctly while in utero [1, 7]. Hypospadias is treated through the creation of neo‐urethra, often from the patient’s prepuce [2, 3]. Neo‐urethra patients unfortunately experience higher rates of stricture and fistulae, even into adulthood [2, 5].

Figure 20: Hypospadias Regions. Stadler, Nature reviews genetics., 4 (2002): 478‐482 [1] Urethral stricture (stricture) describes a narrowing of the urethra from fibrotic tissue blocking the lumen, which restricts urine flow (see Figure 21) [8]. This can be caused by Insertion and use of devices, such a catheter or endoscope, trauma, cancer and radiation therapies, STIs, or previous surgery [8, 9]. Stricture can lead to inflammation, further

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infection, and generally presents a decrease in a patient’s quality of life [8, 9]. Short urethral stricture can be treated by excision and joining, where the limitations of tissue elasticity and structural integrity are negligible [9]. In longer stricture however additional tissue, usually an autologous buccal mucosal graft or urothelium allograft, is required to keep functional continuity of the tissue [10]. This graft replaces the urethral epithelium only, not the surrounding tissue [11, 12]. This substitute material is often an autologous graft [8, 13]. As with many secondary procedures, the buccal mucosal graft is subject to donor‐site morbidity, with recovery sometimes taking months [14‐16].

Figure 21: Anatomy of Stricture. Mayo Clinic, Urethral Stricture, (online) 2018 [8] Tissue engineering and regenerative medicine are two interlinked fields of research that offer new materials and methods for future disease treatment [16, 17]. Tissue engineering explores cell‐compatible constructs, such as an artificial membrane, and bioactive molecules in combination to mimic and replace native tissue [16, 17]. Regenerative medicine uses tissue engineering and advances by with research on how the body self‐heals, how healing can be improved, and how to incorporate this knowledge to rebuild tissue and organs [17]. A paper by de Kemp et al (2015) discussed how these fields could be used to create a standard treatment for hypospadias and stricture [12]. Despite the autologous buccal mucosal graft is a common treatment for stricture, there are over 300 urethroplasty procedures [12]. This paper stresses the point that the number of procedures available show that this is a complex disease with no unified, ideal treatment [12]. The desirable properties for a new membrane using the techniques and materials developed by tissue engineering and regenerative medicine are also described and are summarised below.

 Integrative; the membrane should allow host tissue to bond to and be metabolised. Being biodegradable and/ or bioresorbable allows cells to constructively remodel ECM structures to suit [12, 18, 19].

84 Appendices 85

 Inductive; the membrane should drive new cell growth. This can be accomplished through inclusion of bio‐signalling molecules, such as MMPs, cytokines, and growth factors [12, 18, 19].  Conductive; autologous cell culture should be supported. By expanding autologous cells, the need for a graft is reduced to a biopsy, reducing donor‐site morbidity [12, 16]. Cells will also form the fluid‐proof epithelial layer of the urethra [12, 20].  Mechanically suitable; the material produced should withstand the normal function of the urethra but also surgical handling during surgery [12, 18, 19].  Fabricated; the material should be able to be replicated without human error to help guarantee good manufacturing practices and high‐quality products for patients [12, 16, 18, 19].

The human male urethra functions as the passage for fluid to be expelled while protecting the surrounding tissue [21, 22]. During this function, the urethra will expand both circumferentially and longitudinally as a response to fluid pressure from inside the lumen and/ or elongation from the supporting corpus spongiosum (CS) and corpus cavernosum (see Figure 22) [21]. These tissues are high vascularised with the CS supplying the urethra with blood flow [23]. The urethra protects tissue through intercellular binding proteins, such as tight junction and uroplakins, forming an impermeable layer [24]. Understanding the forces acting upon and the performance of the urethra is important in creating a replacement membrane [12].

Figure 22: Anatomy of the Human Penis. Boeckx, 2017 [22]. The urethra is part of the lower urinary tract and consists of four areas and is developmentally different from the upper urinary tract [21, 25, 26]. Luminal lining of the ureter, bladder, and upper urethra are composed of urothelium, stratified transitional epithelial cells that change shape according to forces applied [21, 25]. These changes send nervous and hormonal signalling that help regulate void and fill phases of the bladder [27, 28]. Beginning at the bladder, the urethra starts with the pre‐prostatic and prostatic urethra,

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both lined with urothelium [21, 25]. The subsequent urethral sections are the membranous and spongy urethra, formed by pseudostratified columnar epithelium, becoming more squamous near the urethral opening [21, 22, 25]. The upper urinary tract of the kidneys and ureters develops from the mesoderm embryonic layer; the lower urinary tract, bladder and urethra, develops from the endoderm layer [26]. The epithelial layer is surrounded first by a layer of connective tissue, the lamina propria, then smooth muscle cells connecting to the CS (see Figure 23) [21, 22, 25]. Understanding these differences in epithelial lining and the vascularisation via surrounding tissue is important as previous inconsideration may be responsible for urethroplasty failure [11].

Figure 23: Human Urethra and CS, Verhoff Strain. de Graaf et al., Tissue Engineering B 23 (2018) [23] Electrospinning (ES) has great interest in tissue engineering as it can create membranes of thin, non‐woven fibres [29‐33]. The technique involves a solubilised polymer being pumped via electronic control through a spinneret from a syringe (see Figure 24) [32]. A high voltage is then applied to the spinneret while an inverse voltage is applied to the collection plate creating an electrical potential [32, 33]. The potential forms the solution into a fine filament as it travels and deposits onto the plate [32, 33]. The solvent evaporates, leaving fibres of the material which can then be removed and further studied [32, 33]. The fibre diameters can tailored to be in the tens to hundreds of nanometres, similar to ECM [29]. Through precision mechanical control of these settings, human error is minimised allowing the membranes to be fabricated in this manner [34]. Mechanical and biological properties of the membrane can alter by selection of polymers in the solution [32, 33]. This makes ES an excellent choice for the study of tissue engineered urethra [35, 36].

86 Appendices 87

Figure 24: Electrospin Setup. Adapted from Bhardwaj and Kundu, Biotechnology Advances, (2010)[32]

Polycaprolactone (PCL), gelatin, decellularised ECM (dECM) and silk fibroin (fibroin) have established use for tissue engineering and regenerative medicine [20, 36‐41]. These polymers possess different mechanical strengths and levels of integrative, conductive, and inductive properties and can each be solubilised for electrospinning separately or as a combination [20, 36‐41].

PCL is a polyester which is slowly, and safely metabolised in vivo [16]. This material already has FDA approval and is used for implants, controlled drug release, and as sutures [42]. While PCL is well tolerated in vivo, it has poor conductive and inductive properties, as it suffers from poor cell‐adhesion and possesses no bio‐signalling cues [40, 43, 44]. PCL can be blended with other materials such as gelatin to improve these shortcomings [40, 43]. Some of the research with PCL has been used to generate membranes for skin and blood vessels [40, 44‐46].

Gelatin is a commercially available biopolymer with good conductive and integrative properties [47‐49]. This material offers the structural support of the parent native tissue by retaining cell binding sites and being easily remodelled by cells to better suit their particular niche [48]. Gelatin has poor mechanical properties however, and is in liquid phase at 37°C, requiring further modification such as crosslinking to be viable for research with tissue engineering and regenerative medicine [20, 43, 47, 50, 51]. Crosslinking improves mechanical properties and prevents solubility [20, 47, 50, 51]. Gelatin is a promising material for creation

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of novel membranes as it can be combined with other polymers, such as dECM [20, 30, 31, 39, 40, 43, 52‐54].

The use of dECM can drive induction as it has excellent cell compatibility properties [52, 55‐57]. The cellular microenvironment of binding sites, and bio‐signalling molecules persist through decellularisation and can make for a potent, if limited, material for research [52, 55‐57]. Once decellularised, the tissue can be repopulated with a desired cell population [20, 30, 57, 58]. This has been demonstrated in decellularised brain, arteries, and cardiac tissue being used for tissue engineering replacement membranes [20, 30, 58].

Equine urethra is a valid source of material for tissue engineering [21]. Compared to human urethra, equine urethra is similar in use of functional process, and is histomorphometrically similar with luminal lining and surrounding CS and smooth muscle cells (see Figure 25) [21, 59, 60]. Equine and human urethra undergo similar forces during function with the distal circumference changing the least and proximal length changing the most [21]. Neither horses, or humans have a sigmoidal flexure, a length of urethra that orientates to facilitate erectile changes [61]. Once decellularised, the urethra could then be repopulated, retaining the structure and composition [57].

Figure 25: Equine Urethra and CS, Trichrome stain. Natali et al., Exp Physiol, 101, 641‐56 (2016) [21]

Using similar tissue can provide site specific growth factors however non‐analogous tissue has also been decellularised and studied [20, 30, 45, 46, 55, 56]. Placenta has been valued in regenerative medicine for its high progenitor cell count, however it also functions as an endocrine organ and has a broad array of growth factors, meaning the dECM could be

88 Appendices 89 likewise valuable to tissue engineering [62‐64]. Some of the novel membranes studied have demonstrated use for angiogenic cardiac and dermal patches, neural tubes, and blood vessels [30, 45, 46, 56, 58]. There exist many protocols for creating decellularised material, consisting of mechanical, chemical, and/ or enzymatic methods however most are very time consuming and take several days [65, 66]. Decellularisation is vital to remove the immune‐ inflammatory remnants of cells which would hinder the growth of newly seeded cells [65, 66].

Fibroin is an easily produced biopolymer from silk cocoons, with mechanical tensile properties that exceed many synthetic fibres [36, 67, 68]. This material also has demonstrated good biocompatibility, with cell attachment, cellular remodelling by ECM proteolytic enzymes and low immunogenicity [36, 69]. Fibroin has been used in the study of bone, blood vessels, bladder, and urethra [36, 70‐72].

In this introduction the target diseases, current treatments, and their implications have been discussed. Tissue engineering and regenerative medicine have been introduced along with promising materials and techniques for generating thin membranes to create a new treatment. A new treatment could therefore resemble the following; a hypospadias or stricture patient would have a small biopsy taken for cheek cells, the cells would then be cultured and expanded while a new membrane is created. This membrane would be electrospun using a combination of materials to support treatment. The expanded culture would then be transferred to the membrane, then further cultured until the membrane can perform as functional tissue. The membrane would then be transferred to the patient, whereby the injury site is stimulated to heal and incorporate the new membrane. Use of the tissue is restored and the new membrane is remodelled and biodegraded.

Using this information, a hypothesis can be formed, ‘Can decellularised ECM be used to create cell‐instructive electrospin membranes for researching treatment of hypospadias and stricture? To answer this, the following aims are put forward. The first aim is to produce decellularised tissue. The second aim is to produce electrospun membranes viable for testing. The third aim is that the generated membrane is to be tested for mechanical and biological properties.

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3. Materials and Methods All materials used from Sigma Aldrich (USA), unless stated otherwise. Type B gelatin from bovine skin (225g Bloom) used (gelatin).

3.1 TISSUE ISOLATION Equine samples collected via donation program at the Veterinary Hospital, Department of Equine Sciences, UMC Utrecht. Horse owners gave informed consent for use of the cadavers for scientific research. Human placenta samples were collected from the maternity ward of the Wilhelmina Children's Hospital (WKZ, Utrecht), patients gave broad informed consent.

5.1.1 3.1.1 Urethra Isolation, Decellularisation, and Digestion

Urethra was isolated based on previously published procedure by Natali et al., (2016) [21]. First, skin and any extraneous tissue was removed while the remaining tissue was preserved with cold phosphate buffered saline (PBS). During dissection, cold PBS was used to keep tissue preserved. Next a plastic pipette was inserted into the proximal end of the urethra to provide stability while handling tissue. An incision was then made into the lateral aspect of the tissue between the tunica albuginea and bulbospongiosus muscle surrounding the corpus spongiosum (CS) and urethra. Beginning at the proximal end, the urethra was then isolated from the penis.

Isolated urethra tube was cut into 5cm lengths and opened to a mucosal sheet. The tissue sheet was then spread, and any remaining spongy urethra was then removed. The urethra was further cut into 5mm pieces and stored in ‐20°C or proceeded to decellularisation.

Urethra isolated based on previously published procedure by Johnson et al., (2015) [73]. 10ml of the isolated and stored urethra (see Error! Reference source not found.) were placed into 50ml falcon tube with sterile dH2O (distilled/ demi‐water) for 1 hour. Tubes were then centrifuged (900G, 5 minutes, room temperature) and the supernatant discarded. 1% (w/v) sodium dodecyl sulphate (SDS) in sterile PBS was then added to the tissue and placed on a rocker protected from sunlight. SDS solution was changed after 24h (hours), 48h, 24h.

SDS solution was then replaced with sterile milliQ ultrapure water (mQH2O) for 24 hours. The tissue was washed until no bubbles formed. Tissue was then lyophilised.

90 Appendices 91

1g of lyophilised tissue was digested with 100mg pepsin (from porcine gastric mucosa) in 0.1M hydrochloric acid (HCl) and stirred for 3 days at room temperature. Digestion was ended by raising pH to 8.0 to inactivate pepsin. Digested material was then lyophilised.

5.1.2 3.1.2 Placental Isolation and Decellularisation, and Digestion

Placental tissue isolated based on previously published procedure by Pelekanos et al., (2016) [63]. The placenta was placed maternal side down and the amniotic membrane and umbilical cord were removed and disposed of. The remaining tissue, decidua basalis and chorion, were cut into large portions and washed twice with 4°C PBS to elute blood. Large portions were then cut into 5mm cubes. Cubes were then stored in ‐20°C or proceeded to decellularisation.

Placental tissue decellularisation based on previously published procedure by Skardal et al., (2015) [74]. 10ml of isolated placental tissue were added to a 50ml falcon tube with sterile dH2O overnight. Tubes were then centrifuged (900G, 5 minutes, room temperature) and the supernatant discarded. Sterile dH2O was replaced 3 times a day, for 2 days and then frozen for 2 days. 2% Triton X100 (TX100) in sterile dH2O was then added to tissue and replaced 2 times a day, for 5 days, then frozen for 2 days. 2% TX100, 0.1% ammonium hydroxide (NH4OH) in sterile dH20 was then added to tissue and replaced twice a day for two days. Sterile mQH2O was then added to tissue and replaced 3 times a day, for 2 days to remove TX100 and NH4OH. Tissue was then freeze dried.

Lyophilised decellularised placental tissue was digested using the same method as equine urethra (see 3.1).

3.2 SILK FIBROIN EXTRACTION Silk fibroin extraction based on previously published procedures by Rockwood et al.,

(2011), Xie et al., (2013), Zhang et al., (2015) [36, 67, 68]. 2L of mQH2O was brought to boil and 4.24g sodium carbonate (Na2CO3). While boiling, Bombyx mori cocoons were cleaned of silkworms and cut to 20mm squares. The internal layer of silk was discarded as this improved the solubility and clarity of silk fibroin. 5g of cocoon was added to boiling 0.02M Na2CO3 and stirred for 30 minutes.

Degummed silk fibroin was removed and rinsed with cold mQH2O. Na2CO3 was discarded. Excess water was squeezed out before transferring silk to 1L of mQH2O and stirring

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for 20 minutes. Silk fibroin was rinsed and stirred a total of 3 times. Silk fibroin was then spread out on foil and dried overnight in a fume hood.

Dried silk fibroin was then dissolved with 8% w/v calcium chloride (CaCl2) in formic acid at room temperature in a petri dish. Dissolved silk fibroin was left to dry overnight in a fume hood. Dried material was then stored for electrospinning.

3.3 SINGLE LAYER ELECTROSPINNING Electrospun membranes that used a 9:1 acetic acid to formic acid solution were created using previously published procedures by Wei et al., (2015), Xie et al., (2013) [36, 41]. Solutions were prepared by dissolving gelatin and PCL and electrospun. Electrospin settings were 11kV, 15cm collector distance, 0.1ml/h flow rate, needle internal diameter 0.413mm (22G).

Electrospun membranes that used a formic acid solution were created using previously published procedures by Kia et al., (2005) [53]. Solutions were prepared by dissolving gelatin, PCL, silk fibroin and electrospun. Electrospin settings were 14kV, 15cm collector distance, 0.1ml/h flow rate, needle internal diameter 0.603mm (20G).

Electrospun membranes that used a 90% v/v acetic acid solution to test placental dECM potential were created using previously published procedures by Baiguera et al., (2013) [20]. Solutions were prepared by sonicating dECM and gelatin and electrospinning. Electrospin settings were 12kV, 10cm collection distance, 0.4ml/h flow rate and 0.838mm needle internal diameter (18G).

Membranes were dried overnight in a fume hood after spinning to remove traces of solvent.

3.4 PCL REINFORCED MULTIPLE LAYER ELECTROSPINNING Multiple layer membranes were created by first spinning alternate layers of PCL and gelatin, or PCL and gelatin/ dECM, see Error! Reference source not found., Table 10. Solutions were placed into a 3ml syringe.

Electrospun membranes that used a 50% v/v acetic acid solution were created using recommendations at UMC, Utrecht. dECM was dissolved 5% w/v then gelatin dissolved 30% w/v in solution then electrospun. Electrospin settings were 7kV, 10cm collection distance, 1.2ml/h flow rate and 0.413mm needle internal diameter (22G). See Table 11. Membranes were dried overnight in a fume hood after spinning to remove traces of solvent.

92 Appendices 93

Table 10: Multi‐layer Electrospin Times, 120 minutes

Sample GM.10.120 GM.5.120 GM.3.120 dM.10.120 Layer 1 PCL 10min PCL 10min PCL 10min PCL 10m Gelatin/ Gelatin 120min Gelatin 120min Gelatin 120min 120 Layer 2 dECM Layer 3 PCL 10min PCL 10min PCL 10min PCL 10m Gelatin/ Gelatin 120min Gelatin 120min Gelatin 120min 120 Layer 4 dECM Layer 5 PCL 10min

Layer 6 Gelatin 120min

Table 11: Multi‐layer Electrospin settings Material PCL Gelatin Gelatin/ dECM

% (w/v) 20 35 35

3:1 Chloroform, 50% Acetic Acid 50% Acetic Acid Solvent DMF

Overnight, 45 minutes, 1 hour, stirring Solubilisation time stirring sonicator

Needle Internal 0.514 0.413 0.413 Diameter, mm

Flow Rate, ml/h 0.5 1.2 1.2

Voltage, kV 11 14 14

Distance, cm 20 10 10

3.5 CROSSLINKING Method 1 ‐ Membrane was frozen at ‐20°C before 10% glutaraldehyde (GLT) at 4°C was slowly pipetted onto the membrane. Membrane was crosslinked for 3 days at 4°C then

washed with mQH2O.

Method 2 ‐ Membrane was frozen at ‐20°C before 9:1 ethanol: GLT solution was slowly pipetted onto the membrane. Membrane was crosslinked for 3 days at 4°C then washed with

mQH2O.

Method 3 ‐ Membrane was placed in a sealed glass chamber containing 25% GLT

vapor. Membrane was crosslinked overnight then washed with mQH2O.

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Method 4 ‐ Membrane was placed in a sealed glass chamber containing 1:9 GLT: ethanol vapor. Membrane was crosslinked overnight then washed with mQH2O.

3.6 SCANNING ELECTRON MICROSCOPY Samples were mounted on stainless steel stubs, fixed with carbon tape, and coated with 6nm gold particle using a Q150R‐S (Quorum, UK). A single sample from each membrane was taken from the centre.

SEM image samples were typically taken at 150x, 500x, 800x, 3000x,5000x,8000x,10000x, 15000x, using a SEM Phenom II (Thermo Fisher Scientific, USA) at 5kV.

3.7 MECHANICAL TESTING Membranes were tested for their tensile properties using a DMA Q800 (Texas Instruments, USA). The following parameters were used; Ramp force, 0.1 N/ min ‐> to 18N; Preload force 0.001 N. Samples were measured to find the width and thickness of the sample using a digital calliper.

Elastic modulus was determined by plotting a stress/ strain curve and measuring the slope between 5 and 10% strain.

3.8 ANALYSIS Statistical analysis (t test, ANOVA, Tukey’s pairwise analysis) was performed using RStudio (Version 1.1.463), using the ggplot2 library.

4. Results and Discussion 4.1 DECELLULARISATION Decellularisation is crucial to remove various biological molecules which elicit an immunoinflammatory response. Methods chosen were based on materials, equipment availability, and efficacy regarding growth factor retention. SDS and TX100 are strong detergents that solubilise cellular membranes so that subsequent washes can remove intercellular molecules and cell surface markers which may trigger immunoinflammatory responses [75‐77]. Growth factors, which are stored in the ECM, show insignificant decline when using these methods. [73, 74].

94 Appendices 95

Decellularisation was performed on equine urethra, using sodium dodecyl sulphate (SDS), and human placenta using TritonX100 (TX100) as described in materials and methods. Both methods were able to reduce native tissue to a white material that was then digested and lyophilised (see Figure 26). Digested tissue was then tested for gelation by solubilising 20% w/v in phosphate buffered saline (PBS) at 37°C. The first two equine samples were not able to produce a hydrogel, while the third equine sample and placental sample did produce hydrogels.

Figure 26: Decellularisation. A) Equine penis, dermis removed. B) Human placenta after amniotic sac and umbilical cord removed. C) Equine urethra after 3 days of SDS washes. D) Pepsin digest of decellularised urethra. E) Lyophilisation Hydrogels formed indicate the presence of intact crosslinking proteins such as collagen telopeptides, laminins, proteoglycans, elastins, fibronectins, and glycosaminoglycans [55, 78‐ 81]. Inability to form a hydrogel from the first samples may indicate insufficient SDS washing

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as the residual ionic detergent would leave a negative charge on the proteins resulting in the peptides repelling each other [76, 82].

The ideal tissue to use would have been the target tissue itself as the ECM would have host‐specific biosignalling molecules. There was however no supply of human male urethra for this project. The next tissue was male equine urethra as it is both histomorphometrically and functionally similar to human urethra [21]. Supply of equine tissue was however not ideal. Only the third sample produced a viable material, and this arrived late in the project. Human placenta was decided upon after the failure of the first two equine samples as it has a high density of growth factors and had better availability [62, 63].

All electrospun dECM membranes discussed here are made using placental tissue. Digested urethral dECM was produced but not used owing to time constraints.

In this step, the first aim of producing decellularised tissue was partially achieved. By digesting and lyophilising the remaining tissue, a fine powder was produced that could be solubilised for electrospinning. No qualifying or quantifying data in the form of a nucleic acid stain or DNA assay, is present however

4.2 ELECTROSPINNING Electrospinning was used to generate membranes for this project using PCL, gelatin, silk fibroin and dECM solubilised polymers. The technique was selected as it can fabricate non‐woven fibres with diameters similar to ECM and has already shown promise in urethral tissue engineering [29, 32, 34‐36]. Materials were chosen for their properties to make the membranes mechanically suitable for research, conductive to support cell culture, integrative to become part of a host, and inductive to drive new cell growth [12]. PCL is a bioresorbable plastic with good mechanical properties that already has FDA approval, however it is hydrophilic often requiring further modification for use in cell culture [43, 83]. Gelatin is a widely available material that retains some of the integrative and conductive properties of the original collagen, it has poor mechanical properties and persists in a liquid state at 37°C, requiring further modification [20, 52]. Silk fibroin has high conductive and integrative properties and has shown much promise for many tissue engineering fields [67, 84, 85]. Specific properties of dECM depend on the origin tissue and decellularisation method but will typically retain the cell binding sites and growth factors, giving it excellent conductive, integrative, and inductive properties, however it is a limiting reagent requiring

96 Appendices 97 long times to process [20, 30, 55, 86]. Combining this technique and materials could produce a membrane suitable for tissue engineering.

The solvents used for the polymer solutions were chosen for their availability and toxicity. Hexafluoroisopropanol (HFIP) is a common solvent for electrospinning, however it has been shown this degrades secondary and tertiary higher protein structures, weakening them [31, 87, 88]. Acetic acid was chosen as it can solubilise ECM and related proteins, such as growth factors, with limited denaturing [30, 89]. Formic acid was used as it can solubilise silk fibroin [67, 85]. HFIP vapours are a known health hazard and the solvent has a higher toxicity (LD50) at 1500mg/kg, compared to acetic acid (3310mg/kg) and formic acid (1800 mg/kg) in rat models [90‐92]. By focusing on acetic acid as a solvent, a membrane with better mechanical and conductive properties could be obtained.

Fibre diameter and morphology is the result of a complex relationship between different electrospinning parameters involved during creation of the membrane [32, 33, 93‐ 95]. These parameters can be ordered to three general groups; solution, processing, and environmental [32, 94, 96]. Solution parameters describe the physicochemical properties of the prepared material; viscosity, surface tension, electrical conductivity, solvent volatility, and the concentration of polymers used [32, 94, 96]. Processing parameters are the settings used during electrospinning; applied voltage to the spinneret, internal diameter of the spinneret, distance to the collector plate, flow‐ or feed‐rate of the polymer in the syringe [32, 94, 96]. Environmental parameters include the localised, ambient conditions around the electrospinning; temperature, relative humidity, air flow, atmospheric composition, and atmospheric pressure [32, 94, 96]. Electrospinning for this project was performed in an unsealed cabinet, making the environmental parameters dependent upon the laboratory air conditioning facilities to control.

Viscosity is known to be the most important factor for electrospinning with higher solution viscosities producing thicker fibre diameters [32, 33, 93‐95, 97, 98]. Viscosity is an interplay between the concentration, and molecular weight, of the polymer, and solvent used to solubilise [94, 97, 98]. The high surface tension of a viscous solution will offer more resistance to the applied charge, producing thicker fibre diameters [94, 97, 98]. Force of surface tension reverts a solution to a low energy sphere shape while the electrical potential draws the solution into a jet to form an elongated fibre [94, 97, 98]. A viscosity that exceeds the force of the potential can form bead‐, blob‐, or pearl‐like morphologies in the depositing membrane from the high surface tension, if fibres form at all [94, 97, 98]. Low polymer

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concentration, however, will lack sufficient interaction and entanglement to produce continuous fibres, instead producing jets of droplets described as, ‘electrospraying’ [96].

Fibres in the membrane are also influenced by the volatility of the solvent and electrical conductivity of the solution [38, 94, 96, 97, 99, 100]. Through addition of an ionic reagent, such as monopotassium phosphate or sodium chloride salts, the electrical resistance of the solution, reducing the effect of viscosity and allowing thinner fibres to form as a result [38, 96, 97]. Thinner fibres can also be obtained by using more potent solvents to increase solution conductivity [38, 95]. Rate of evaporation, or volatility, of the solution affects both diameter and morphology of the fibre [94, 99, 100]. Slow evaporation in hydrophilic polymers leads to thicker fibres as water is absorbed when in high relative humidity (>60%) conditions slows precipitation of the polymer [94, 100]. In hydrophobic polymers, high relative humidity causes the material to precipitate sooner creating fibres with thinner diameters that are prone to forming pores [94, 99, 100]. The solutions used in this project were based on existing methods to generate electrospun membranes for tissue engineering research [20, 36, 41, 53].

The settings used during electrospinning also factor into the fibre diameter and morphology [38, 50, 96, 97, 101, 102]. The flow‐ or feed‐rate of the solution through the spinneret minimises the effect of evaporation by providing consistent new material [97, 101]. A low feed rate gives thicker fibre diameters as the viscosity of the droplet is increased with evaporation [97, 101]. A larger internal diameter of the spinneret allows a larger droplet to form also reducing evaporation [97, 101]. A small droplet evaporates faster, increasing viscosity and resulting in thicker fibre diameters [97, 101].

Increasing working distance between the spinneret tip and collector plate can affect fibres through a decrease in force of the electrical potential, more mid‐flight evaporation, and additional stress as the filament travels [50, 93, 102]. The weakened potential and extra evaporation can create thicker fibres, while the travelling can split the filament [50, 93, 102]. Thinner fibres can be obtained by increasing the applied voltage to the spinneret [50, 102]. The more elongated fibres are a result of surface tension being minimised in comparison to the charge [50, 102]. There is a point, however where the applied charge far exceeds surface tension, beginning to produce beads akin to electrospraying [93, 96, 103]. Process parameters used were based on existing methods to generate electrospun membranes for tissue engineering research [20, 36, 41, 53].

98 Appendices 99

After electrospinning, the foil collector plate was removed and placed in a fume hood overnight to dry, allowing any remaining solvent to evaporate. Membranes were then removed from the foil and small samples were taken for scanning electron microscopy. These samples were coated with 6nm gold nanoparticles (Quorum, UK) to allow for clearer imaging at the UMC (Nederlands) microbiology lab. Samples were analysed for fibre morphology and diameter with a SEM‐Phenom II (Thermo Fisher, USA) using a 5kV electron beam (see Materials and Methods). Multiple magnifications were used to assess the fibres however for consistency and comparison 5000x is used in this discussion.

4.2.1 EARLY ELECTROSPINNING ATTEMPTS Initial membranes (see Method 3.5) were created by solubilising 20% w/v precursor materials in acetic acid: formic acid solution (9: 1) and electrospinning (see Single Layer Electrospinning, 3.3). Settings used were; 0.1 ml/h flow rate, 11kV applied voltage, 15cm collection distance, needle internal diameter 0.603mm, for 6 hours. 20% weight/ volume (w/v) of polymers were used to generate membranes. Polymers used were PCL, gelatin, and 75:25 PCL: gelatin, with the latter being mixed after solubilising.

Figure 27: Morphology of initial membranes. 5000x, scale bar = 10µm. Electrospin settings; 20% w/v in 9:1 acetic acid: formic acid, 0.1ml/h, 11kV, 15cm, needle internal diameter 0.603mm, 6 hours. A) PCL‐only. B) Gelatin‐only. C) 75: 25 PCL: Gelatin blend.

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Figure 28: Fibre Diameter of Initial Membranes. n = 50, α = 0.05

Examining the initial membranes (see Figure 27) the PCL membrane (A; 452 ± 211nm) displays non‐straight fibres and defects. The gelatin membrane (B; 424 ± 57nm) shows smooth straight fibres. PCL‐gelatin membrane (C; 515 ± 191nm) had smooth, straight fibres. Fibre diameters were not found to be significant (see Figure 28; n = 50, α = 0.05) between gelatin and PCL‐gelatin membrane but were found to be significantly different between PCL and gelatin, and PCL and PCL‐gelatin.

The many blob defect morphologies and non‐straight fibres are indicative of a solution with high viscosity [33, 50, 54, 93, 97, 104]. A jet still formed from the spinneret however the low flow rate and collection distance may have allowed excessive evaporation to occur to the filament mid‐flight [50, 93]. The increasing surface tension has then reverted elongated, straight fibres to thicker, rounder morphologies. Thin filaments between the blobs have likely undergone more whipping force mid‐flight and have resulted in non‐straight fibres [50, 93]. Reducing the collection distance and increasing flow rate will likely overcome these issues if this method is tried in the future [39, 93, 98].

Fibre diameters were not found to be significant (see Figure 28) between gelatin and PCL‐gelatin membrane but were found to be significantly different between PCL and gelatin, and PCL and PCL‐gelatin. Gelatin has a typically lower molecular mass (20‐100 kDa) compared to PCL (80 kDa) and hydrolyses when in aqueous solution [31, 39, 105]. This results in a gelatin

100 Appendices 101 solution with less viscosity and higher electrical conductivity than a similar concentration of PCL solution [31, 39]. In turn this creates a membrane with lower fibre diameters [31, 39]. Combining two polymers results in a solution with different properties, proportional to their mix ratio [31, 39, 106]. During this attempt it was found that a 75: 25 PCL: gelatin mix was enough to significantly decrease the fibre diameter from PCL but not significantly increase from gelatin. Both gelatin and PCL‐gelatin present with smooth, straight fibres indicating that the electrospinning was successful [32, 94, 107].

Membranes generated this way were thin, brittle, and would break when attempting to remove from the foil collector plate. It was hypothesised that blending silk fibroin to the solution would improve mechanical properties. When blending silk fibroin to this mix the acetic acid was unable to solubilise the silk and a gel formed. No dECM was used at this stage as these were attempts to find suitable electrospin settings and dECM is not abundant.

This method was abandoned as the membranes produced were too fragile to be repeatably handled and thus not suitable for further testing. Silk fibroin was not able to be included in the spin and thus could not improve mechanical properties. As formic acid is used in the production of silk fibroin, a new method only using formic acid was devised.

4.2.2 ELECTROSPINNING WITH FORMIC ACID The new method was designed around using formic acid to solubilise silk fibroin, PCL and gelatin (see method 3.3). This was to improve mechanical properties of the membranes and allow them to be handled. Electrospin settings used were; 0.1 ml/h flow rate, 14kV applied voltage, 15cm collection distance, needle internal diameter 0.603mm, for 6 hours. Precursor materials were solubilised at different concentrations; 20%, 15%, 10% w/v. Membranes generated were; PCL; 75:25 PCL: gelatin. Gelatin membranes were not produced as the solution was not viscous enough electrospin. Solutions were mixed after solubilising. No dECM was used in this method as the material was not ready and this method was untested.

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Figure 29: Membranes generated using formic acid as a solution. 5000x scale bar=10µm. Electrospin settings; various % w/v in formic acid, 0.1ml/h, 14kV, 15cm, needle internal diameter 0.603mm, 6 hours. A) 20% PCL B) 15% PCL C) 10% PCL D) 20% 75:25 PCL: Gelatin E) 15% 75:25 PCL: Gelatin F) 10% 75:25 PCL: Gelatin.

102 Appendices 103

Figure 30: Fibre diameter of membranes created from polymers solubilised in formic acid. n = 50, α = 0.05

Analysing the PCL membranes (see Figure 29 ) A (20% w/v, 125 ± 31nm) presents thin, straight fibres with some blob defects present. B (15% w/v, 341 ± 187nm) presents non‐ straight, thin and thick fibres. C (10% w/v, 128 ± 52nm) presents also with non‐straight, thin and thick fibres with some blob defects. Analysing the PCL‐gelatin (75:25) membranes (see Figure 29) D (20% w/v, 114 ± 67nm) presents non‐straight thick and thin fibres with bead morphologies. E (15% w/v, 234 ± 78nm) presents straight fibres. F (10% w/v, 232 ± 133nm) presents non‐straight, thick and thin fibres with bead morphologies.

Comparing the fibre diameters (see Figure 30), increasing concentration of PCL from 10% (w/v) to 15% gave significantly thicker fibres, increasing to 20% saw a significant decrease while 10% and 20% PCL fibres are not significantly different. Blending PCL and gelatin (75:25) saw a significant increase in fibre diameter at 10%, a significant decrease at 15%, and an insignificant decrease at 20%.

Thin and thick fibre diameters (as seen in B, C, D, F, in Figure 29) are indicative of an electrospin that has undergone splaying [50, 93]. This occurs when the filament splits into

Appendices 103

multiple, uneven fibres due to axial stress generated from the whipping motion mid‐flight [50, 93]. Should the stress exceed the surface tension of the solution, splits will occur and deposit on the collector plate [50, 93]. Splaying is typically a result from insufficient viscosity and/ or over application of voltage [50, 93]. This is further supported by the formation of bead morphologies, indicating the polymer solution was near to undergoing electrospraying [98].

The counterintuitive results between PCL and PCL‐Gelatin membrane fibre diameters may be explained with environmental factors [99, 100]. Literature and previous electrospinning on the project support an increase in polymer concentration increasing mean fibre diameter [32, 33, 93‐95, 97, 98]. This is seen between 10% and 15% for both PCL and PCL‐gelatin membranes, however this is not evidenced for the increase between 15% and 20%. Addition of gelatin to PCL is likewise supported and known to decrease fibre diameter as viscosity changes [54]. This is clearly seen in 15% yet the opposite is true at 10%. As there was no means of isolating, controlling, or recording the electrospinning chamber atmosphere at this point in the project, it’s unknown what effects the microenvironment may have played during the electrospinning for these membranes [99, 100]. A warm, humid day would increase evaporation rates while the humidity would absorb into the solution and precipitate the PCL, creating fibres with thicker diameters [99, 100]. What may be likely is that 10%, 15%, and 20% w/v polymer concentration is insufficient to produce significantly different fibre diameters as formic acid is stronger than acetic acid [54, 90, 92, 93]. Environmental factors that favour quick evaporation occurred on the days that 10% PCL‐Gelatin, 15% PCL, and 15% PCL‐Gelatin were electrospun. Any future attempts with this method should start with a higher polymer concentration to resist splaying and higher flow rate to lessen effects of evaporation [97, 99‐101].

Membranes produced here were part of a larger set that was to include silk fibroin. Issues with morphology were not yet known as the SEM was not available until later in the project.

4.2.3 ELECTROSPINNING WITH SILK FIBROIN Silk fibroin was prepared (see method 3.3) by degumming Bombyx mori cocoons and then solubilised as a precursor polymer in formic acid. PCL and gelatin were also solubilised as separate precursors and then mixed prior to electrospinning. Solutions prepared were at 15% w/v. Membranes generated were silk fibroin, and PCL‐gelatin‐silk fibroin at 80:10:10, 70:20:10. Electrospin settings used were; 0.1 ml/h flow rate, 14kV applied voltage, 15cm

104 Appendices 105 collection distance, needle internal diameter 0.603mm, for 6 hours. Membranes were analysed using SEM and calculating mean fibre diameter (n = 50, α = 0.05).

Figure 31:Membranes generated using silk fibroin in formic acid. 5000x scale bar=10µm. Electrospin settings; 15% in formic acid, 0.1ml/h, 14kV, 15cm, needle internal diameter 0.603mm, 6 hours. A) 15% Silk Fibroin) B) 15% 80:10:10 PCL:Gelatin:Silk fibroin blend C) 15% 70:20:10 PCL:Gelatin:Silk fibroin blend

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Figure 32: Fibre diameter of membranes created from silk fibroin polymers solubilised in formic acid. n = 50, α = 0.05

Inspecting morphologies of 15% w/v electrospun membranes (see Figure 31) silk fibroin (A, 147 ± 56nm) shows some blob‐ and bead‐shapes with non‐straight fibres. PCL‐ gelatin‐silk fibroin (80:10:10; B, 134 ± 68) shows very fine fibres with blob shapes. PCL‐ gelatin‐silk fibroin (70:20:10; C, 186 ± 114) shows thick and thin non‐straight fibres. SEM images D and E are from previous Figure 29 but included here for ease of comparison.

Comparing mean fibre diameters (see Figure 32) the silk fibroin produces thinner fibres than PCL and PCL‐gelatin membranes. Mixing in PCL and gelatin, 80:10:10 membrane fibre diameters are not significantly smaller than the single silk polymer membrane, while 70:20:10 fibre diameters are not significantly larger.

Decrease in mean fibre diameter between PCL and silk fibroin membranes could be likely due to the increase in conductivity of the solution [41, 84, 85]. Silk fibroin is a protein structure containing positive and negative amino acids, similar to gelatin, and allows for the applied charge to further elongate the filament [41]. This could however be evidence for environmental factors coming into play as 15% PCL and 15% PCL‐gelatin give significantly thicker fibres compared to 15% silk fibroin [41]. Combining PCL, gelatin, and silk should yield thicker fibres compared to silk alone, however there is no significant change here [41]. Silk fibroin membrane, and membranes containing silk, at 15% w/v have similar fibre diameters to PCL membranes at 10% (128 ± 52nm) and 20% (125 ± 31nm). Together with the presence

106 Appendices 107 of blob shapes, this may indicate a higher concentration was needed to successfully electrospin the solution [98].

As the molecular weight of silk fibroin is dependent upon the initial degumming preparation, this method could also be attempted using silk that has been boiled in calcium chloride for less time [67, 108, 109]. Longer degumming stages reduces the molecular weight of silk fibroin, which results in thinner fibre diameters [67, 108, 109].

Membranes produced using this method were also to include electrospinning with 10% and 20% w/v solutions. Upon analysing the fibres with SEM, the then‐current state of the project and further complexity required to obtain the settings for a successful electrospin became apparent. Silk fibroin containing membranes produced here were not valid for future use as the beaded morphology is poor for mechanical properties and cell culture [110]. Time required to find the optimal settings for a successful electrospun membrane would require time that was not available. It was decided to focus on developing a novel, cell instructive membrane from dECM and reinforcing that membrane with layers of PCL to achieve better tensile properties.

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4.2.4 ELECTROSPINNING WITH DECM Initial attempts to generate a membrane using only dECM resulted in electrospraying. The first successful membrane generated using dECM and gelatin was created using a method described by Baiguera et al. (2014)[20]. Briefly, precursor polymers (15% w/v; gelatin, dECM at 14:1) were solubilised in 90% acetic acid then mixed and electrospun using; 0.4ml/h flow rate, 12kV applied voltage, 10cm collection distance and 0.838mm needle internal diameter, for 2 hours. A second round of successful membranes were generated by solubilising polymers (35% w/v; gelatin, dECM at 30:5) in 50% acetic acid and electrospinning at 1.2 ml/h flow rate, 14kV applied voltage, 10cm collection distance, and 0.413mm needle internal diameter, for 2 hours. Membranes were analysed using SEM and calculating mean fibre diameter (n = 50, α = 0.05).

Figure 33: Fibre Morphology of gelatin and gelatin/dECM blend membranes. 5000x scale bar =10µm. Electrospin settings (A, B); 15% w/v in 90% acetic acid, 0.4ml/h, 12kV, 10cm, needle internal diameter 0.838mm, 2 hours. Electrospin settings (C, D); 35% w/v in 50% acetic acid, 1.2ml/h, 14kV, 10cm, needle internal diameter 0.413mm, 2 hours. A) Gelatin. B) 14:1 Gelatin:dECM blend. C) Gelatin D) 30:5 Gelatin:dECM

108 Appendices 109

Figure 34: Fibre diameter of gelatin and gelatin/ dECM at 15% and 35% w/v. n = 50, α = 0.05

Inspecting fibre morphologies (see Figure 33), smooth straight fibres are present in all membranes; 15% gelatin (A, 213 ± 22nm, 35% Humidty (H)), 15% gelatin‐dECM (B, 342 ± 73nm, 36% H), 35% gelatin (C, 514 ± 82nm, 31%H), 35% gelatin‐dECM (D, 391 ± 43nm, 36%H). Fibre diameters for A, B similar to published results [20].

Comparing fibre diameters (see Figure 34) at 15% w/v, the mixing of dECM to gelatin resulted in significantly thicker fibre diameters, while at 35% mixing in dECM resulted in significantly thinner fibres. 35% solutions gave significantly thicker diameters over 15%.

These novel membranes confirm that placental dECM can be electrospun albeit with gelatin. The first attempts that resulted in a low‐quality electrospray informed that a higher concentration of polymer, as changing electrospinning settings was unable to form a filament to deposit the polymer as fibres. The method was scaled up to 35% w/v, using 30:5 gelatin: dECM to create thicker fibres for better handling during testing.

The increase in mean fibre diameter from 15% to 35% can be explain by both the addition of extra material, and less acetic acid in the solution reducing conductivity [38, 94, 96, 97, 99, 100]. At 15%, mixing gelatin and dECM (14:1) creates thicker fibres, explained through an increase in viscosity [20]. Enzymatically digested tissue retains many ECM proteins that are removed during hydrolysis to produce gelatin production, creating a more viscous solution [20, 55, 105]. At 35%, mixing gelatin and dECM (30:5) creates thinner fibres, due to the presence of salt [38, 96, 97]. Decellularised tissue is digested with pepsin and

Appendices 109

hydrochloric acid. This digest is ended by neutralising with sodium hydroxide, however there is no dialysis step afterwards to remove the sodium chloride that would form as a result [33, 97, 101].

As this was a novel solution, membranes produced under different electrospin settings were explored [50, 97].

4.2.5 EFFECTS OF ELECTROSPIN SETTINGS ON DECM Gelatin‐dECM was solubilised (35% w/v, 30:5) in 50% acetic acid and then electrospun using differing settings. The default (control) setting was 1.2 ml/h flow rate, 14kV applied voltage, 10cm collection distance, and 0.413mm needle internal diameter, for 2 hours. Flow‐ rate, needle internal diameter, distance, and applied voltage settings were altered, one per membrane, to generate a high and low for each parameter. Membranes were analysed using SEM and calculating mean fibre diameter (n = 50, α = 0.05).

110 Appendices 111

Figure 35: Effect of different electrospinning parameters. Electrospin settings for control membrane; 35% gelatin/ dECM membrane in 50% acetic acid, 1.2ml/ h flow rate, 10cm distance, 14kV applied voltage, needle internal diameter 0.413mm, 2 hours. Parameter changes were; A) no change B) flow rate decreased to 0.5ml/h C) distance increased to 20cm D) voltage decreased to 12kV E) voltage increased to 16kV F) needle

Appendices 111

internal diameter decreased to 0.260 mm G) needle internal diameter increased to 0.686 mm.

Figure 36: Effect of different spinning parameters on fibre diameter. n = 50, α = 0.05. Default electrospin settings; 35% w/v in 50% acetic acid, 1.2ml/h, 14kV, 10cm, needle internal diameter 0.413mm, 2 hours. A) Flow rate B) Internal diameter C) distance to collector plate from spinneret. D) Applied voltage. Examining the fibre morphologies (see Figure 35) and fibre diameters (see Figure 36) the control membrane (A, 391 ± 43nm, 36%H) shows smooth, straight fibres indicative of a successful spin. Decreasing the flow rate (B, 519 ± 72nm, 26%H) significantly increases mean fibre diameter and shows smooth straight fibres. Increasing the distance (C, 481 ± 182nm, 34%H) significantly increases mean fibre diameter and shows some flat, “ribbon like” non‐ straight fibres. A decrease of applied voltage (D, 448 ± 57nm, 26%H) significantly increases mean fibre diameter shows smooth, straight fibres, while increasing the applied voltage (E, 469 ± 78nm, 26%) gives the same result. Decreasing needle internal diameter (F, 448 ± 57nm, 32%H) significantly increases mean fibre diameter and shows smooth, non‐straight fibres, while increasing the diameter (G, 469 ± 78nm, 27%H) significantly decreases mean fibre

112 Appendices 113 diameter and shows smooth straight fibres. “Ribbon like” fibres noted in B, D, G however these are single instances and not indicative of the membrane.

Increase in fibre diameter from a decrease in flow rate (Figure 35, B) is likely a result from evaporation of the droplet creating a more viscous solution, a higher flow rate supplies new solution to reduce this effect [97, 101]. The settings used (1.2 and 0.5 ml/h) represent rates that could be reliably left unmonitored for several minutes. Below 0.5ml/h the droplet would dry, preventing the jet and filament to form (see Figure 37, A). Above 1.2 ml/h the droplet excess would form separate jets and create a membrane from the collector plate to the stage below the spinneret (see Figure 37, B). Other rates would electrospin but required the machine to be frequently stopped to remove dried filaments.

Figure 37: Mid‐spin feed‐rate Issues. A) Filament dries mid‐flight B) Falling droplets are turned to fibres connecting from the connection plate to the stage below. The “ribbon‐like” fibres on the increased distance (Figure 35, C ) are a result of the outside of the filament drying mid‐flight and forming a ‘skin’ that then collapses when depositing on the collector plate [50, 102]. This is a result of the electrical potential weakened by the distance making a slower filament [50, 102]. A distance of 5cm was tested but this produced electrospraying. A 15cm distance was also used while electrospinning however this membrane was damaged irreparably damaged before SEM could take place and there was insufficient time to create another membrane at this setting.

Both decreasing applied voltage (D) and increasing (E) has increased fibre diameter. This is likely due to the environmental conditions as the control (14kV) membrane was spun at 36% humidity while the 12kV membrane (D) was spun at 26% humidity and the 16kV membrane (E) was spun at 27% humidity. This decrease in humidity has likely been sufficient to increase evaporation of the solution, increasing viscosity, therefore creating thicker fibres [94, 99, 100]. As viscosity is a critical factor in determining fibre characteristics, evaporation

Appendices 113

is a stronger force than the change in electrical field strength during electrospinning over this range of voltages [32, 101].

Size of the internal needle diameter has been shown here to effect fibre diameter with smaller sizes creating smaller droplets, allowing for more evaporation, in turn raising viscosity, resulting in thicker fibre diameters [97, 101].

By using different electrospin settings, how this novel solution performs is better understood. Membranes produced could tolerate removal from foil collector plate and be reliably handled.

4.2.6 MULTILAYERED ELECTROSPINNING Membranes were generated through alternating layers PCL for mechanical support with gelatin‐dECM for cell combability (see method 3.4) as PCL, gelatin‐dECM, PCL, gelatin‐ dECM. Gelatin and dECM were solubilised (35% w/v, 30:5) in 50% acetic acid and then electrospun using the settings; 1.2 ml/h flow rate, 7kV applied voltage, 10cm collection distance, and 0.413mm needle internal diameter, for 2 hours per layer. PCL (20% w/v) was solubilised in DMF: chloroform 3 :1) and electrospun using the settings; 0.5 ml/h flow rate, 11kV applied voltage, 20cm collection distance, and 0.514mm needle internal diameter, for 10 minutes per layer. Membranes are then transferred to a fume hood to allow the rest of the solvent to evaporate. Membranes were analysed using SEM and calculating mean fibre diameter (n = 50, α = 0.05).

114 Appendices 115

Figure 38: Effect of multilayering with PCL. 5000x scale bar 10 µm. Electrospin settings; 35% w/v in 50% acetic acid. 1.2ml/h, 14kV, 10cm,0.413mm internal diameter, 2 hours. A) Gelatin only B) Gelatin/ dECM blend C) PCL reinforced gelatin D) PCL reinforced gelatin/ dECM. E) PCL‐only membrane

Appendices 115

Figure 39: Fibre diameter for membranes with‐ and without membrane reinforcement. n = 50, α = 0.05 Comparing fibre morphologies (see Figure 38) and fibre diameters (see Figure 39), smooth straight fibres for PCL reinforced gelatin (C, 335 ± 43nm, 46%H) and PCL reinforced gelatin‐dECM (D, 307 ± 78nm, 45%H). Smooth fibres with twists present in PCL membrane (E, 2014 ± 422nm, 30%H). Images A (514 ± 82nm, 31%H) and B (391 ± 43nm,36%H) are from previous Figure 33 but included here for ease of comparison. Adding dECM to gelatin had significantly lowered mean fibre diameter previously, however when creating a membrane reinforced with PCL, there was no significant decrease between the fibre diameters.

The significant decrease in mean fibre diameter between the reinforced and non‐ reinforced membranes may be a result of gas composition and humidity [32, 94, 100]. SEM images of the membrane are only from the topmost layer, electrospinning however had been taking place for 4 hours prior from previous layers. The extended time would have allowed for a higher presence of acetic acid vapor in the electrospin cabinet [32, 94, 100]. This change may have slowed both evaporation of acetic acid and absorption of water vapor into the membrane, resulting in thinner fibres [32, 94, 100]. The twists present in the PCL membrane are likely a result of the filament drying mid‐flight and undergoing axial stress, this can be likely prevented in the future by decreasing the collection distance [50, 93]. The decrease in humidity may also have been the cause of an increase of fibre diameters between the non‐ reinforced (31, 36%H) and reinforced (46, 45%H) [94, 99, 100].

116 Appendices 117

In this step, the second aim of producing a viable membrane for further testing was partially achieved. The new method using 35% w/v of dECM and gelatin, or just gelatin, in 50% acetic acid produced membranes that could be handled both with and without PCL reinforcement.

4.3 STABILISING GELATIN VIA CROSSLINKING While the generated membranes were able to be handled, a considerable portion was gelatin and would dissolve in warm aqueous solutions, such as during cell culture. To allow cell studies to be performed on these membranes, the gelatin would need be stabilised through crosslinking [20, 50]. Crosslinking prevents the solubility of gelatin by joining polymer chains together through bonding [47, 111]. Glutaraldehyde crosslinks gelatin by reacting with the nucleophilic protein groups [51, 112]. The two reactive ends of glutaraldehyde covalently bond with the gelatin amines creating a large insoluble polymer [51, 112].

Membranes were cut to strips and exposed to glutaraldehyde either in solution, or by vapor, to crosslink. Membranes were then washed with cold milliQ water to remove traces of the crosslinker and left in a fume hood overnight to dry. Membranes were analysed using SEM from both top‐down and cross sectional to qualitatively determine how successful crosslinking was.

Appendices 117

Figure 40: Cross section of single layer and reinforced membranes. 500x, scale bar = 100µm. A) 35% Gelatin. B) 35% Gelatin/ dECM. C) PCL reinforced gelatin. D) PCL reinforced gelatin/ dECM The cross sections (see Figure 40) show fibres converging to a single plane for gelatin membranes (A), two planes of convergence for gelatin‐dECM membranes with many fibres between them (B), three layers for PCL reinforced gelatin membranes with some fibres between them (C), two layers with a single plane of convergence with some fibres between them for PCL‐reinforced gelatin‐dECM (D). The cross‐section images were created by cutting the dry membrane with a scalpel at room temperature, coating with gold, and then imaged. The vacuums created for SEM and gold coating have damaged the membranes by splitting them into layers, while the planes of convergence are created by cutting with a scalpel [113]. To eliminate the cutting, freeze fracturing should be used in the future [113].

118 Appendices 119

Figure 41: Effect of different crosslinking methods. 5000x scale bar 10µm. A) 35% Gelatin membrane soaked in 10% glutaraldehyde at 4°C for 3 days. B) 35% Gelatin/ dECM membrane exposed to 9:1 ethanol: glutaraldehyde vapor overnight at room temperature. C) 35% Gelatin membrane soaked in 9:1 ethanol: glutaraldehyde at 4°C overnight. D) 35% Gelatin‐dECM membrane exposed to 25% glutaraldehyde vapor overnight at room temperature. Comparing the morphology of the crosslinked membranes (see Figure 41) both solution‐based crosslinking (A, C) attempts have resulted in a solid film. Vapor‐based crosslinking (B, D) produced a porous film, with some fibre shapes still present.

Solution based crosslinking was performed by cooling the material and reagents down to ‐20°C, or 4% for the solution, to prevent freezing solid. As this should have prevented solubilisation, it’s possible that the rinsing has partially solubilised the gelatin [20, 39]. Porosity is important as this allows cells to facilitate biocommunication, angiogenesis, and vascularisation [32, 41]. Without porosity, the solution‐based membranes are not fit for research.

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Figure 42: Effect of crosslinking, cross section. A) PCL reinforced gelatin, 3 days in 10% glutaraldehyde, 4°C solution. B) Gelatin/ dECM blend, overnight in 25% glutaraldehyde vapor. C) Gelatin/ dECM blend, overnight in 9: 1 ethanol: glutaraldehyde, vapor. D) Gelatin/ dECM blend, overnight in 9: 1 ethanol: glutaraldehyde vapor. The crosslinked cross‐section (see Figure 40) a stratum of solid layers can be seen (A, B, C) in both solution‐based and vapor‐based methods. In D, a non‐smooth surface is present. Porosity has been maintained in the ethanol‐glutaraldehyde (9:1) vapor crosslinking.

All membranes crosslinked would shrivel and fold while rinsing out the glutaraldehyde. This made them difficult to further test. Vapor crosslinking is shown to be the most promising here however the exact method is still yet to be determined. Other crosslinking agents or functionalisation with chemical groups that can be covalently crosslinked are likely to be worth exploring also [47, 111].

In this step, the second aim of producing a viable membrane for testing was still addressed. The project produced membranes but failed to produce viable membranes.

4.4 MECHANICAL TESTING OF ELECTROSPUN MEMBRANES There are five main mechanical strengths a material may be attributed with. Compression, tensile, torsion, and shear, each describing the resistance to the respective

120 Appendices 121 orientated force with yield denoting resistance to deformation. Compression describes force pressing inward, such as a knee supporting the body [114]; tensile is pulling, like a tendon in use[115], torsion is twisting, such as a spine during turning[116]; shear is forces applied in opposition, as what can be seen in bone fracture[117]. Lastly yield strength is the force required to irreversibly stress a material as plastic deformation. Elastic deformation is where a material will return to its original shape and is described by the elastic modulus[118]. By understanding this information, the mechanical requirements of a new membrane treatment are understood.

Performing mechanical testing informs how the membrane tolerates physical forces through generation of an elastic modulus [21, 24]. This property denotes the force a material can withstand before changing shape [21, 24]. Tensile strength is an important property as native urethra expands and stretches during use [21, 24]. Electrospun membranes were tested for their tensile mechanical properties to see the effect of PCL reinforcement. Membranes were tested in a dry, un‐crosslinked state, as they were not able to be handled when wet. No PCL membranes were tested as they were too thin (<0.01mm) to be measured with a digital calliper.

Tensile forces were investigated for this project, as the urethral epithelium expands during function. Mueller et al., (2008) [119] report a tensile modulus between 5 and 21 kPa, which was validated in part by Masri et al., (2018) [120] who found distal regions to be more elastic and longitudinal directions less elastic. This is explained by the corpus cavernosum applying more force longitudinally during erection compared to circumferential force during micturition [21]. The reported human moduli are more rigid than equine urethra, which is between 4.46 and 14.41 kPa [21]. This bears mentioning as the future of this project will involve decellularised equine urethra for generating membranes.

Appendices 121

Figure 43: Elastic modulus of non‐ and reinforced membranes. Membrane strips of 5mm by 20mm were subjected to 0.1N/min ramp force. Moduli were recorded between 5% and 10% strain. A) non‐reinforced membranes B) Membrane reinforced with PCL. Comparing the elastic modulus (see Figure 43) it was found that gelatin membranes (0.05 ± 0.01 MPa) could be significantly improved upon by adding dECM (0.75 ± 0.70 MPa) significantly improved the modulus of gelatin membranes (0.05 ± 0.01 MPa), without PCL reinforcement. With PCL reinforcement, gelatin membranes (38.35 ± 8.64 MPa) could again be significantly improved with dECM (72.32 ± 10.42 MPa). Reinforcing membranes with PCL has significantly improved the elastic moduli of membranes through its own strong tensile properties [39, 41, 43]. The mix of other structural ECM proteins, such as laminin and elastic, allow for additional polymer bonds to form and improve mechanical resistance [20, 53, 55, 111].

The elastic modulus target of 14.41 kPa for urethra and 0.8 MPa for surgical handling has been achieved by using PCL reinforcement [21, 36, 37, 121]. These properties could again be improved with successful crosslinking [47, 111].

In this step, the third aim of testing the membranes for biological and mechanical properties was partially addressed. By reinforcing the membranes with layers of PCL membranes displayed a higher resistance to mechanical forces, above the reported values for functional tissue and surgical handling.

122 Appendices 123

5. Conclusions and Future Work This project investigated the hypothesis, ‘can decellularised ECM be used to create cell‐instructive electrospin membranes for the treatment of hypospadias and stricture?’ This question was to be answered with three aims; first, to generate decellularised tissue; second, to electrospin a membrane viable for testing; third, was to test the membranes for mechanical and biological properties.

On the first aim, both equine urethra and human placenta appeared decellularised, as the resulting product was white to off‐white in colour. No DNA assay or histochemical staining was carried out to verify this procedure, however. On the second aim, many membranes were produced over the course of this project. Single‐ and multi‐layered membranes with dECM were produced that could be handled. The fibre diameter varies according to the settings used. No successful crosslinking was performed which was vital for performing further studies however a vapor‐based method appeared most promising. On the third aim, no cell studies could be performed as the membranes remained water soluble. Mechanical tests were performed but will serve only as a baseline once an effective crosslinking method is established but did quantitatively show the membranes were sufficiently resistant to tensile forces when incorporating PCL reinforcement layers.

To conclude, the hypothesis of this project remains unproven. Effective crosslinking that allows further testing is still required, followed by cell studies to verify the ‘cell instructivity’ and mechanical properties. Additional testing is required to for presence of DNA and sGAG proteins.

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