Natural Marine Derived Bioactive Glass Based Scaffolds with Improved Functionalities

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Natural Marine Derived Bioactive Glass based Scaffolds with Improved Functionalities

Bioaktive glasbasierte Scaffolds aus Naturschwamm mit verbesserter Funktionalität

Der Technischen Fakultät

der Friedrich-Alexander-Universität Erlangen Nürnberg

zur Erlangung des Grades

DOKTOR-INGENIEUR

vorgelegt von

Elena Boccardi aus Clusone

Als Dissertation genehmigt von der Technischen Fakultät der Friedrich-Alexander-Universität Erlangen-Nürnberg

Tag der mündlichen Prüfung: 19.12.2016

Vorsitzende/r des Promotionsorgans: Prof. Dr.-Ing. Reinhard Lerch Gutachter: Prof. Dr.-Ing. Aldo R. Boccaccini Prof. Dr. Enrica Vernè

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Acknowledgments

Firstly, I would like to express my sincere gratitude to my supervisor Prof. Aldo R. Boccaccini, who gave me the opportunity to do my PhD research at the Institute of Biomaterials. I want to thank him for his support during the past three year and for the opportunities to take part to numerous international conferences, which had a great impact on my knowledge and personal experience.

My gratitude also for trusting me to work for two European projects in the past two years, ReBioStent and Mozart, which helped me to get more confidence in my capabilities and the opportunity to see the link between university and industry.

Besides my advisor, a special thank you goes to the coordinators of the European project GlaCERCo, which founded my first year of PhD. Thanks to Prof. Monica Ferraris, Prof. Milena Salvo and Cristiana Contardi for this amazing opportunity and experience. Special thanks also to all the students that took part to project: Dr. Anahí Philippart, Dr. Luca Bertolla, Dr. Sandra Cabañas-Polo, Dr. Auristela de Miranda, Tassos Toulitsis, Dr. Rama Krishna Chinnam, Dr. Ines Ponsot, Parisa Eslami, Dr. Harshit Porwal, Dr. Ben Milsom, Paulina Perdika, Dr. Richa Saggar, Panagiota Kleverkloglou. It was a great experience. I would like to thank Stafanos Giannis and all the people in Element who hosted me for two week during my secondment. Thanks to Prof. Chiara Vitale-Brovarone, Dr. Giorgia Novajra, Dr. Francesco Baino and Lucia Pontiroli for helping me during my stay in Turing and for the fruitful collaboration.

A special mention is for Dr. Lina Altomare and Dr. Luigi De Nardo, whom supported me through the bachelor and the mater thesis and gave me the confidence to not give up and look for a PhD. I have also to thank you because you gave me the possibility to come back to my old lab for cell culture study, supporting my work once again. Thanks to all the group of Mancinelli, especially to Monica Moscatelli and Dario Picenoni. Thanks to Dr. Virginia Melli for all the nice discussions and opinion exchanges, it has been a pleasure to work with you. When we were talking together sometimes I was thinking “She is crazier than me, I like that!”

This work would not have been possible without the help of many people which actively contributed to make this PhD come true. I want to thank all the collaborators that helped me with the analyses. Thanks a lot to Dr. Judith A. Roether, Dr. Yaping Ding, Dirk Dippold, Dr. Jochen Schmidt, Dr. Liliana Liverani, Barbara Myszka, Nicoletta Toniolo, Dr. Mirza Mackovic, Christian Dolle, Dr. Luis Cordero, Kai Zheng, Francesca Ciraldo. Special

v mention to Dr. Ana M. Beltrán because no matter in which country she is working, she has always the time and the pleasure to help me with enthusiasm!

Thanks to all the member of the Biomaterial Department. Special mention to Heinz Maler and Dr. Julia Will for their support and infinite kindness. Thanks to Dr. Gerhard Frank because he was always ready to help. Thank you for your time, especially for have been a great SEM teacher. To Frau Bärbel Wust, who was always kind and ready to fight for me! Thanks to Jasmin, who is the pillar of the lab and will become rich writing a book on students´ “adventures” (or many of them… you have more than enough material). Thanks to all my colleagues, especially to the Spanish group and mensa group! Anahí, Luis, Sandra, Michael, Giulia, Valentina thanks for making me feel welcome and more proud to be Italian =). I know that sometimes it is difficult to believe, until you don´t hear my accent, but yes I am Italian!!!!! Special thanks to Anahí and Kai, who are not only incredible colleagues but amazing friends. Thanks to Dr.Dr. Sandra and Luis (Dr. Cordero when he doesn’t answer to email) for our lovely fight, the institute was not the same without the two of you. A special thanks to Judith, because you always listened during my “…nothing is working…”. Worst you had to put up with my amazing English… also now that you are reading these acknowledgments (sorrryyy).

Thanks to my students, whom had the bad experience to have me as supervisor. Thanks to Jonas Hahn, Selma Uzun, Verena Hahn, Ben Huang, Roman Guenther, Johanna Dueck Bularca and Laine Paidere.

Thanks to “le donnine dei Marinoni e di Premolo”, Crisitna e Fabiana, for your friendship also now that we are leaving far away. The University is almost over, so for the first time in I don´t even remember how many years you will not have to listen to me complaining about exams, PhD, professors or students! We have to organize a wedding, so let´s change topic!

Thanks to my family, especially my parents for your support and love. The best help you ever gave to me was to trust my decision and support me without interfering. Grazie, vi voglio bene!

Thanks to Gabriele for always being at my side. Ti amo.

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List of Publications

E. Boccardi, I.V. Belova, G.E. Murch, A.R. Boccaccini, T. Fiedler, “Oxygen diffusion in marine-derived tissue engineeringscaffolds”, Journal of Materials Science Materials in Medicine, 26(6): 200 (2015), DOI: 10.1007/s10856-015-5531-2

E. Boccardi, A. Philippart, J.A. Juhasz-Bortuzzo, G. Novajra, C. Vitale-Brovarone, A.R. Boccaccini, “Characterization of Bioglass based foams developed via replication of natural marine sponges”, Advances in Applied Ceramics, 114: S56-S62 (2015), DOI: 10.1179/1743676115Y.0000000036

E. Boccardi, A. Philippart, J.A. Juhasz-Bortuzzo, A.M. Beltran, G. Novajra, E. Spiecker, A.R. Boccaccini, “Uniform surface modification of 3D Bioglass®-based scaffolds with mesoporous silica particles (MCM-41) for enhancing drug delivery capability”, Frontiers in Bioengineering and Biotechnology, 3 (2015), DOI: 10.3389/fbioe.2015.00177

E. Boccardi, A. Philippart, V. Melli, L. Altomare, L. De Nardo, G. Novajra, C. Vitale- Brovarone, T. Fey, A.R. Boccaccini, ”Bioactivity and Mechanical Stability of Natural Marine Sponges based Bioglass® scaffolds”, Annals of Biomedical Engineering, 2016, DOI: 10.1007/s10439-016-1595-5

Nomenclature and Abbreviations

3D Three dimensional σ Stress state ε Deformation � Porosity µ-CT Micro Computed Tomography °C Degree Celcius Ag_MCM-41 Silver doped ordered mesoporous silica particles Al2O3 Aluminium oxide BET Brunauer Emmet Teller BGs Bioactive glasses BJH Barrett, Joyner and Halenda BG-PU Bioglass®-based scaffolds prepared with polyurethane as sacrificial template BG-SA Bioglass®-based scaffolds prepared with Spongia Agaricina as sacrificial template BG-SL Bioglass®-based scaffolds prepared with Spongia Lamella as sacrificial template BTE Bone tissue engineering CaCl2 Calcium chloride CaO Calcium oxide Co-MBG Sol-gel glass composition 78SiO2 20CaO 1.2P2O5 0.8CoO (mol%) Co-MBG BG-PU Bioglass®-based scaffolds prepared with polyurethane as sacrificial template coated with Cobalt doped ordered mesoporous sol-gel glass synthesis solution (78SiO2 20CaO 1.2P2O5 0.8CoO mol%) Co-MBG BG-SA Bioglass®-based scaffolds prepared with Spongia Agaricina as sacrificial template coated with Cobalt doped ordered mesoporous sol-gel glass synthesis solution (78SiO2 20CaO 1.2P2O5 0.8CoO mol%) CTAB n-decyltrimethylammonium bromide Deff Effective diffusivity Dt Bulk diffusivity DC Dip coating method DCM Direct coating method ECFs Extracellular fluid ECM Extracellular matrix EDS Energy dispersive spectroscopy EISA Evaporation-induced self-assembly EPD Electrophoretic deposition ERMI® Endosseous Ridge Maintenance Implant F127 Pluronic® copolymer FDA Food and Drug Administration FT-IR Fourier transformed infrared spectroscopy GPa Giga pascal HA hydroxyapatite HCA hydroxycarbonate apatite HRTEM High resolution transmission electron microscopy Ib Bioactivity index t0.5bb 50% of the interface is bond to tissue IGFs Insulin-like growth factors

IGF-I Insulin-like growth factor I IGF-II Insulin-like growth factor II ICP-OES Inductively coupled plasma IMP Impregnation method IUPAC International Union of Pure and Applied Chemistry KBr Potassium bromide KCl Potassium chloride K2HPO4 (3H2O) Potassium Hydrogen Phosphate Trihydrate kN Kilo newton M Molarity (mol/L) MCM-41 Ordered mesoporous silica particles MCM-41_A 29 mL H2O, 18.5 mL NH3, 0.2 g CTAB, 1 mL TEOS MCM-41_B 11 mL H2O, 18 mL EtOH, 18.5 mL NH3, 0.2 g CTAB, 1 mL TEOS MCM-41_C 4 mL H2O, 25 mL EtOH, 18.5 mL NH3, 0.2 g CTAB, 1 mL TEOS MCM-41_D 11 mL H2O, 19 mL EtOH, 3.3 mL NH3, 0.2 g CTAB, 1.2 mL TEOS MBGs Mesoporous bioactive glasses MEP® Middle ear prosthesis MgCl2 (6H2O) Magnesium Chloride Hexahydrate MPa mega pascal N Newton NaCl Sodium chloride NaHCO3 Sodium bicarbonate Na2O Sodium oxide Na2SO4 Sodium sulfate µm Micrometer nm Nanometer P123 Pluronic® copolymer P2O5 Phosphorous oxide PCM Post coating method PCM_A Post coating method with wet particles PCM_B Post coating method with dried particles PEI Polyethyleneimine ppi Parts per inch PTFE Polytetrafluoroethylene PU Polyurethane PVA Polyvinyl alcohol rcf Relative centrifugal force rpm Revolutions per minute RT Room temperature (R´O)3SiR Organosilanes SA Spongia Agaricina SL Spongia Lamella SAXRD Small angle X-Ray diffractometry SBF Simulated body fluid SDD Silicon drift detector SEM Scanning electron microscopy SiC Silicon carbide SiO2 Silicon dioxide Si-O-Si Siloxane bonds Si(HO)4 Hydrated silicon dioxide Si3N4 Silicon nitride SL Spongia Lamella

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SrCu-MBG Sol-gel glass composition 78SiO2 20CaO 1.2P2O5 0.4CuO 0.4SrO (mol%) SrCu-MBG BG-PU Bioglass®-based scaffolds prepared with polyurethane as sacrificial template coated with Strontium and Cupper doped ordered mesoporous sol-gel glass synthesis solution (78SiO2 20CaO 1.2P2O5 0.4CuO 0.4SrO mol%) SrCu-MBG BG-SA Bioglass®-based scaffolds prepared with Spongia Agaricina as sacrificial template coated with Strontium and Cupper doped ordered mesoporous sol-gel glass synthesis solution (78SiO2 20CaO 1.2P2O5 0.4CuO 0.4SrO mol%) TE Tissue engineering TEM Transmission electron microscopy TEOS Tetraethyl orthosilicate TEP Triethyl phosphate TIE Thermal Ion Exchange Tg Temperature of glassy transition TGFβs Transforming growth factors beta Tm Melting temperature TRIS Tris(hydroxymethyl)aminomethane UHMWPE Ultra-high-molecular-weight polyethylene Wt.% weight percent XRD X-Ray diffractometry

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Abstract

Since its development in late 1960s, 45S5 Bioglass® is one of the most studied biomaterials for bone defect repair due to its property to bond with both soft and hard tissues and the capability to release ions overtime that have a positive effect on osteoblast gene activation. This material has been successfully used in clinical applications as particulates or monolithic, but porous 3D Bioglass® scaffolds are not yet available for clinical applications due to the relatively low mechanical properties associated to the high porosity (> 90%) of the constructs.

The present work developed 45S5 Bioglass®-based foams via a replica method using natural marine sponges as sacrificial templates. The goal was to produce Bioglass®-based scaffolds with improved mechanical properties (up to 4 MPa) due to a reduction of the total porosity (68%) but without affecting the pore interconnectivity (> 99%), which plays a key role in cell ingrowth and neovascularization. Natural marine sponges, thanks to the millenarian evolution in water filtration, were considered adequate sacrificial template for production of bone tissue engineering scaffolds due to their efficient interconnected porous structure. The obtained scaffolds were characterized by high bioactivity and mechanical stability also after one month in simulated body fluid (SBF), achieving compressive strength values of 1.2 ± 0.2 MPa.

In order to enhance even more the mechanical properties of these Bioglass® scaffolds based on natural marine sponges and to introduce functional ions and drug release capability, different functional coatings with ordered mesoporous materials were developed. The approach was to combine the bioactivity of the bioactive glass based scaffolds with the drug uptake and release capability of ordered mesoporous silica based materials.

Ordered mesoporous silica particles, MCM-41 and MCM-41 doped with Ag (Ag_MCM-41), were synthetized and used as a functional coating with the effect of increasing the mechanical properties, up to 11 MPa, of the Bioglass®-based scaffolds, enhance the bioactivity of this combined system and to introduce the possibility to release locally molecules (e.g. antibiotics) and Ag ions for antibacterial properties.

Ordered mesoporous sol-gel bioactive glasses, MBGs, containing functional Co2+, Sr2+ and 2+ Cu ions were also produced. Two different MBGs compositions were considered: 78SiO2 2+ 20CaO 1.2P2O5 0.4CuO 0.4SrO mol% and 78SiO2 20CaO 1.2P2O5 0.8CoO mol%. Co and Cu2+ are both known for their capability to enhance angiogenesis due to their hypoxia- mimicking effect. Moreover, Cu2+ ions can also significantly inhibit bacterial viability. Sr2+

xv ions, as other divalent ions, e.g. Zn2+ and Mg2+, can stimulate the osteogenic/cementogenic differentiation of human bone marrow stromal cells. The two sol-gel glass solutions were successfully used to produce homogeneous coatings on the surface of the Bioglass®-based scaffolds and also in the inner core of the scaffolds. The systems showed high bioactivity and superior mechanical properties (compressive strength up to 0.1 MPa for scaffolds prepared using polyurethane as template).

Summarizing, the present research project was successful in producing a new family of scaffolds for bone tissue engineering based on 45S5 Bioglass®, and exhibiting improved mechanical properties, bioactivity and drug release capability.

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Zusammenfassung

Seit seiner Entwicklung in den späten 1960er Jahren ist 45S5 Bioglass® eines der meist studiertesten Biomaterialien für die Instandsetzung von Knochendefekten aufgrund der Eigenschaft sowohl an hartes wie auch an weiches Gewebe binden zu können und der Fähigkeit, Ionen freizusetzen, die positiv auf die Genaktivierung in Osteoblasten wirken. Dieses Material konnte schon in Form von Partikeln oder monolithisch erfolgreich in klinischen Anwendungen eingesetzt werden. Dreidimensionale und poröse Bioglass® Scaffolds sind aber wegen der relativ geringen mechanischen Festigkeiten, zurückzuführen auf die hohe Porosität (> 90%) der Konstrukte, noch nicht für den klinischen Einsatz verfügbar.

Die vorliegende Arbeit beschreibt die Entwicklung 45S5 Bioglass®-basierter Schäume mit Hilfe der Schaumreplikationsmethode unter Verwendung von Naturschwamm als Template. Das Ziel war die Produktion von Bioglass® Scaffolds mit verbesserten mechanischen Eigenschaften (bis zu 4 MPa) durch die Reduzierung der Porosität (68%) bei gleichzeitiger Einhaltung der Poreninterkonnektivität (> 99%), welche eine große Rolle beim Einwachsen von Zellen und der Neovaskularisation spielt. Die Naturschwämme wurden als potenzielle Vorläufer für die Produktion von Scaffolds für den Bereich Knochen Tissue Engineering in Erwägung gezogen, da sie dank ihrer tausendjährigen Evolution in der Wasserfiltration eine effiziente interkonnektierende Porenstruktur aufweisen. Die erhaltenen Scaffolds sind gekennzeichnet durch hohe Bioaktivität und mechanische Festigkeit im Bereich von 1,2 ± 0,2 MPa sogar nach einem Monat in simulierter Körperflüssigkeit (SBF).

Um die mechanischen Eigenschaften der Bioglass® Scaffolds auf der Basis von Naturschwamm weiterhin zu erhöhen und um funktionelle Ionen- und Wirkstofffreisetzung einzubringen, wurden unterschiedliche funktionelle Beschichtung mit mesoporösen Materialien entwickelt. Diese Herangehensweise sollte die Bioaktivität der bioaktiven glasbasierten Scaffolds mit der Fähigkeit der Wirkstoffaufnahme und –freisetzung von geordneten mesoporösen silikatischen Materialien vereinen.

Geordnete mesoporöse Silicapartikel, MCM-41 und MCM-41 dotiert mit Silber (Ag_MCM- 41), wurden synthetisiert und als funktionelle Beschichtung angewandt mit dem Ergebnis, dass die mechanischen Eigenschaften der Bioglass® Scaffolds bis auf 11 MPa erhöht werden konnten. Auch konnte innerhalb des kombinierten Systems die Bioaktivität erhöht

xvii werden und die Möglichkeit der lokalen Freisetzung von Medikamenten (z.B. Antibiotika) oder Silberionen (antibakterielle Eigenschaften) eingeführt werden.

Zusätzlich wurden auch geordnete mesoporöse bioaktive Sol-Gel Gläser (MBGs) hergestellt, die funktionelle Co2+, Sr2+ und Cu2+ Ionen enthielten. Zwei unterschiedliche MBGs

Zusammensetzungen wurden dabei in Betracht gezogen: 78SiO2 20CaO 1,2P2O5 0,4CuO 2+ 0,4SrO (in mol%) und 78SiO2 20CaO 1,2P2O5 0,8 CoO (in mol%). Sowohl Co wie auch Cu2+ sind für ihre Fähigkeit bekannt, durch die Nachahmung von Hypoxie die Angiogenese zu verbessern. Darüber hinaus können Cu2+ Ionen die Lebensfähigkeit von Bakterien signifikant herabsetzen. Sr2+ Ionen, wie auch andere zweiwertige Ionen (z.B. Zn2+ und Mg2+), können die osteogene/cementogene Differenzierung von humanen stromalen Stammzellen des Knochenmarks anregen. Die zwei Sol-Gel Glas Lösungen wurden erfolgreich verwendet, um homogene Beschichtungen auf der Oberfläche und auch im Inneren der Bioglass® Scaffolds herzustellen. Das System zeigte hohe Bioaktivität und überragende mechanische Festigkeiten auf (Druckfestigkeiten bis zu 0,1 MPa für Scaffolds, die mit Hilfe von Polyurethan als Template hergestellt wurden).

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Table of Contents

Acknowledgments v

List of Publications ix

Nomenclature and Abbreviations xi

Abstract xv

Zusammenfassung xvii

Table of Contents xix

1. Introduction 1 1.1. Clinical Problem 1 1.2. Aim of the Thesis 2

2. Fundamentals 5 2.1. Bone Tissue: Functions, Composition, Structure and Mechanical Properties 5 2.2. Bone Tissue Engineering (BTE) 8 2.3. Classification of Biomaterials 10 2.4. Bioceramics 11 2.4.1. Structure of Glass 13 2.4.2. Silicate Bioactive Glasses 17 2.4.3. The Mechanism of Bioactivity 19 2.4.4. Two Classes of Bioactive Glasses 22 2.4.5. Clinical Applications of Bioactive Glasses 24 2.4.6 Production of Porous Bioactive Glass Scaffolds 26 2.5. Ordered Mesoporous Silicate Materials 28 2.5.1. Ordered Mesoporous Silica Nanoparticles. 31 2.5.2. Ordered Mesoporous Bioactive Sol-gel Glasses 32 2.5.3 Ordered Mesoporous Silica Materials as Drug Delivery System 34

3. Characterization Techniques 39 3.1. Light Microscope 39

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3.2. Scanning Electron Microscopy (SEM) – Energy Dispersed X-Ray 39 Spectroscopy (EDS) 3.3. High Resolution Transmission Electron Microscopy (HRTEM) 39 3.4. Micro Computed Tomography (micro-CT) 40 3.5. X-Ray Diffraction (XRD) 40 3.6. Fourier Transform Infrared Spectroscopy (FT-IR) 41 3.7. Nitrogen Sorption Analysis 41 3.8. Inductively Coupled Plasma Atomic Emission Spectroscopy (ICP-OES) 42 3.9. Compression Test 42

4. Scaffolds Production and Characterization 45 4.1. Introduction 45 4.2. Materials and Methods 46 4.2.1. Scaffold Production 46 4.2.2. Bioactivity Study and Characterization 48 4.2.3. In Vitro Cell Culture Test 49 i. Indirect Cell Culture Test 49 ii. Direct Cell Culture Test 50 4.2.4. Oxygen Diffusion Evaluation 50 4.3. Results 51 4.3.1. Templates and Scaffolds Architecture 51 4.3.2. Micro-CT Analysis 57 4.3.3. XRD Analysis 59 4.3.4. Mechanical Test 59 4.3.5. Bioactivity and Mechanical Stability 60 i. Weight and Porosity Variation 60 ii. Mechanical Stability 61 iii. Ion Release and pH Variation 62 iv. Surface Modification Analysis: SEM and FT-IR 63 v. Micro-CT Analysis 66 vi. Cell Culture Study 68 vii. Oxygen Diffusivity 69 4.4. Discussion 72 4.5. Conclusions 77

5. MCM-41 Coating of BG-based Scaffolds 79

5.1. Introduction 79 5.2. Materials and Methods 81 5.2.1. Synthesis of MCM-41 Particles 81 5.2.2. Synthesis of Ag_MCM-41 Particles 82 5.2.3. Composite System Preparation 82 5.2.4. Drug Release Test 84 5.2.5. Bioactivity and Stability of the Composite System 84 5.2.6. Characterization Techniques 85 5.3. Results 86 5.3.1. Optimization of MCM-41 Particles Production Process 86 5.3.2. Drug Release of MCM-41 particles 90 5.3.3. Bioactivity and Degradation of MCM-41 Particles in SBF 91 i. Surface Modification 91 ii. Ordered Mesostructure Modification 93 iii. FT-IR Analysis 96 iv. Ion Release and pH Variation 96 5.3.4. Ag nanoparticles-doped MCM-41 particles 97 i. Degradability of Ag_MCM-41 Particles 100 5.3.5. MCM-41 Particles Coating of BG Scaffolds 104 i. Mechanical properties 111 ii. Bioactivity in SBF 111 iii. pH Variation 114 iv. Drug release Capability 115 5.4. Discussion 116 5.5. Conclusions 121

6. Sol-Gel Glass Coating of BG Scaffolds 123 6.1. Introduction 123 6.2. Materials and Methods 124 6.2.1. Co-MBG Synthesis 124 6.2.2. SrCu-MBG Synthesis 124 6.2.3. MBG coated BG-based scaffolds Preparation 125 6.2.4. Bioactivity Test in SBF 125 6.2.5. Drug Up-take and Release Capability of MBGs 126 6.2.6. Capillarity Test 127 6.3. Results 128

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6.3.1. Co-MBG and SrCu-MBG: Preparation and Characterization 128 i. Bioactivity Evaluation 129 ii. Drug Test 131 6.3.2. Double Thermal Treated BG Pellets: Bioactivity Evaluation 132 6.3.3. MBG Coating by Dip Coating (DC) Method 133 i. Coating Optimization 133 ii. Mechanical Properties Evaluation 134 iii. Bioactivity Test 135 6.3.4. MBG Coating by Impregnation (IMP) Method 135 i. Coating Optimization 135 ii. Mechanical Properties Evaluation 138 iii. Bioactivity Test of MBG coated BG-based Scaffolds 138 in SBF iv. Capillarity Test 148 6.4. Discussion 149 6.5. Conclusions 153

7. Conclusions and Future Directions 155 7.1. Conclusions 155 159 7.2. Future Work

References 163

Appendix – Oxygen Diffusion Evaluation 173

List of Figures 175

List of Tables 185

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1. Introduction 1.1. Clinical Problem Bone is a dynamic connective tissue that undergoes continuous remodeling during the entire lifetime of a mammal1,2. Its high regenerative capability means that the majority of fractures can heal without the need of surgical intervention. Despite this, large bone defects associated with accidents or tumor resection, lack the required template for complete regeneration of the tissue and surgery is often necessary2–4.

The present “gold standard”, the autograft, involves harvesting “donor” bone from a non- load-bearing site and transplanting it into the defect site of the same patient1. An alternative to autograft is tissue engineering. The tissue engineering approach combines principles and methods of engineering and life sciences in order to develop biological substitutes that restore, maintain and improve tissue function. This can be done by combining 3D scaffolds made of suitable biomaterials with cells and biological molecules, such as growth factors3,5–7. Using biodegradable and bioactive porous 3D scaffolds, the aim is to reproduce the complex morphology of bone tissue using bioactive, osteoinductive materials, which are able to integrate with the biological environment and promote the regeneration of bone. The use of synthetic scaffolds avoids the limitations of autografts, such as additional operating time, healing of both donor and implant sites, as well as the pain and increased risk of infection at the donor site4,8.

Since its development in late 1960s9,10, 45S5 Bioglass® is one of the most important biomaterials considered for bone tissue repair: it is a degradable and surface reactive silicate glass with a high content of calcium and is able to form a stable bond with both soft and hard tissues. In fact, once Bioglass® is in contact with biological fluids, a layer of carbonate hydroxyapatite (HCA) similar to the mineral phase of bone is formed on the glass surface and the release of Ca and Si ions can stimulate the expression of several osteoblasts genes, being thus both osteoconductive and osteoinductive11. In addition, Bioglass® has been also shown to induce angiogenesis in vitro and in vivo providing thus an important function in bone tissue regeneration, namely induction of vascularization12.

Bioglass® has been used for many clinical applications as particulates and bulk in non-load bearing site as bone grafting and for the prevention of dental hypersensitivity10,13,14. Porous 3D Bioglass® scaffolds were first produced in 2006 by the replica method by Chen et al.15. However they are not yet available on the market for clinical application, due to the

1 difficulty to match at the same time the required mechanical properties and biological functions of real bone matrix6,10.

1.2. Aim of the Thesis The aim of the present work was the development of mechanically robust Bioglass®-based scaffolds via replica method of natural marine sponges instead of using polyurethane foams as templates15. The approach followed in this work was to reproduce natural sponges’ interconnected porous structure, optimized over years of evolution for water filtration. The specific goal was to obtain Bioglass®-based scaffolds with improved mechanical properties due to a reduction of the total porosity but without affecting the pore interconnectivity of the 3D structure. Moreover, in order to improve the functionality of these scaffolds, ordered mesoporous silica based materials were developed as a coating. The presence of ordered porosity with high specific surface area can open the possibility to introduce drug delivery capability to the final system. These sol-gel based silicate materials can be also doped with metallic ions, which are biologically active and can play a key role as antibacterial agent or to enhance neovascularization and new bone formation.

The thesis is organized in the following manner: Chapter 2 contains the fundamentals introducing the topics of relevance for the research project, including bone tissue, a classification of biomaterials, bioceramics and ordered mesoporous silicate materials. In Chapter 3 the characterization techniques used within this project are summarized. Chapter 4 presents the fabrication and characterization of bioactive glass-based scaffolds production and characterization. Chapter 5 and 6 contain the development and characterization of ordered mesoporous materials to be used as functional coating of the scaffolds. Chapter 7 summarizes the main achievements of this thesis work and suggests future work. Figure 1.1 shows schematically the overview of research tasks carried out in this thesis to achieve the set goals.

Figure 1.1 Schematic overview of the research approach of the present work: preparation of BG-based scaffolds with improved mechanical properties and development of functional coatings with ordered mesoporous silica materials, i.e. silica mesoporous MCM-41 particles and ion-doped mesoporous bioactive glasses (MBGs), for the release of metallic ions and drugs once in contact with the biological fluids.

2. Fundamentals 2.1. Bone Tissue: Functions, Composition, Structure and Mechanical Properties Bone is a specialized connective tissue. It is highly dynamic and vascularized and it remodels during the entire lifetime of mammals1,16,17. It plays a key role in human physiology and it fulfils three main functions including: i) mechanical: locomotion, ensuring that the skeleton has adequate load-bearing capability, protection of vital organs and levers for muscle actions; ii) metabolic: homeostasis through its storage of calcium and phosphorous ions; iii) homeostatic: it is involved in haematopoiesis, the formation of blood cells by the bone marrow16,17.

The bone matrix is made up of a mineral or inorganic matrix (65-70%), mainly mineral crystals (90% of calcium phosphate and 10% of calcium carbonate), that confers to bone tissue hardness and mechanical resistance, and the rest (20-25%) is an organic material. The organic matrix, called osteoid, is composed (90-96%) of collagen type I and 4-10% of non- collagenous proteins, such as proteoglycan and glycoproteins (Table 2.1)16. Bone mineral phase has been identified as a non-stoichiometric, poorly crystallized, highly defective form of mineral apatite, which accounts for almost two thirds of the weight of bone and occupies 40% of the volume17. These small crystals of impure hydroxyapatite are in the shapes of needles, plates and rods and are located within and between collagen fibers17. Combining the relatively brittle mineral phase with the flexible collagen bone results in a composite material.

Table 2.1 Composition of bone tissue (table modified from Redaelli and Montevecchi16).

Components Amounts in Weight [%] Inorganic matrix 65-70 Organic matrix (ECM) 20-25 Collagen 90-96 of ECM Proteoglycan and Glycoproteins 4-10 of ECM Water 9

As function of the dimension and of the order of the collagen fibers it is possible to recognize two different types of bone tissue: fibrous bone tissue and lamellar bone tissue. The fibrous tissue is characterized by collagen fibers with a diameter of 5-10 µm, that are

5 deposited randomly in the space and are woven together. This is considered as immature bone, or primitive bone, because it is the first type of bone deposited both during the physiological development and during the fracture repair. Than the fibrous bone is substituted by the lamellar bone which is the mature one. The collagen fibers (~ 60 nm of diameter) are organized in a well-ordered structure called bone lamellae. The stress-oriented collagen of lamellar bone gives its typical anisotropic properties16.

Bone has several levels of organization17. On the macroscale it is possible to identify two different types of bone: trabecular (spongy or cancellous) bone and cortical (dense or compact) bone16. Trabecular bone is mostly found at the epiphysis and metaphysis of long bones and inside flat or small bones. Trabecular bone looks like a sponge: it has an extensive network of bone beams, called trabeculae, which are oriented in function of the external loading. The trabeculae delimit cavities, which contain hematopoietic bone marrow. The compact bone, that appears as a solid continuous mass17, forms the most superficial part of short, flat and long bones, as well the diaphysis of these last. These two types of bone are biologically identical; the difference is in how the microstructure is arranged. The relative quantities of each type of bone and the architecture vary in a manner which reflects its overall shape, position and functional role18. The porosity of the bone tissue varies from 5% of the compact bone to the 90% of the spongy bone16.

At the microscale bone consists mainly of a number of cylindrical units called Harvesian system, which ran parallel to the long axis of bone. Each Harvesian system has a central canal, which is surrounded by concentric lamellae. The number of lamellae varies from 4 to 20. The central cavity contains one or two blood vessels and nerves and is 15-20 µm in diametr16,17. The overall diameter of a Harvesian system is in the range 100-200 µm. In figure 1 a schematic representation of the different organization levels of bone.

Bone contains four different types of cells: osteoprogenitor cells, osteoblasts, osteocytes and osteoclasts, of which osteocytes are the most abundant. Osteoclasts come from the hematopoietic stem cells and are able to dissolve the bone matrix, osteoblasts come from the differentiation of mesenchymal stem cells and are able to produce new bone matrix. The osteoblasts differentiate in mature bone cells, called osteocytes, when they are surrounded by bone matrix they produced. Osteocytes, star-shaped cells, reside inside spaces called lacunae and they are networked to each other via long cytoplasmic extensions that occupy tiny canals called canaliculi, which are used for exchange of nutrients and waste through gap junctions. They are responsible for the maintenance of bone and they are able to start the bone remodeling.

To accomplish its functions, bone undergoes continuous remodeling which involves bone resorption, carried out by osteoclasts and bone formation which occurs by the action of osteoblasts. In the adult physiological skeleton the two processes are in balance, maintaining a constant, homeostatically controlled amount of bone. This fact, as well as the histological observation that osteoclastic bone resorption is followed by osteoblastic bone formation, led to the concept that the two processes are mechanistically coupled (Figure 2.1).

Figure 2.1 Schematic illustration of bone formation/bone remodeling process including osteoblast and osteoclast differentiation from mesenchymal and hematopoietic stem cell, respectively; matured osteoblast become osteocytes and are entrapped in the bone matrix.

Existing evidences suggest that multiple factors are most probably involved in the maintenance of bone homeostasis5. Growth factors found in bone, such as IGFs or TGFβs, were proposed to be released during resorption and initiate local bone formation. Moreover, the ability of bone to change its structure and adapt to different mechanical loads implies that mechanical forces can regulate bone resorption and formation. If bone is subjected to loads in a physiological range, deposition and resorption of bone matrix are in perfect balance. Increased loads should promote the formation of bone and impare resorption in order to increase the total resistant area; on the other hand decreased loads should have the opposite effect with the consequent bone resorption (Figure 2.2).

Figure 2.2 Bone mechanical homeostasis mediated by different loads. When loads are in a physiological range (εmin<ε<εmax) deposition and resorption of bone matrix are in perfect balance. When ε<εmin the resorption of bone is preponderant, when ε<εmax the deposition is faster than resorption.

Another factor able to influence the deposition and the resorption of bone matrix is the calcemia in the extracellular fluids (ECFs), which is the concentration of calcium in the blood. The calcemia values should be kept in a physiological range (adult 8.5-10.5 mg/100 mL of blood, child 10-12 mg/100 mL of blood), independently from the adsorption of calcium through food intake and its elimination through kidneys. In fact the skeleton contains the 90% of the whole calcium in the body. When there is a condition of ipercalcemia there is the formation of new bone tissue, instead when occurs a hypocalcemia there is the resorption of bone.

2.2. Bone Tissue Engineering (BTE) Tissue engineering (TE) represents an emerging multidisciplinary field, which involves the application of the principles and methods of engineering and life science towards the fundamental understanding of structure-function relationship in normal and pathological tissues and the development of biological substitutes that restore, maintain or improve tissue functions19. The driving force behind TE is to create biological substitutes capable of replacing the damaged tissues. This is done by combining scaffolds, cells and biological signals in order to create living, physiological, three-dimensional tissues6. Cells and signally molecules, i.e. growth factors, can be seeded into highly biodegradable scaffolds, cultured in vitro and subsequently implanted into bone defects to induce and direct the growth of new bone. The first function of a scaffold is its role as substratum that allows cells to attach,

8 proliferate, differentiate (i.e., transform from a non-specific or primitive state into cells exhibiting the bone specific functions), and organize into normal, healthy bone as the scaffold degrades. Figure 2.3 summarizes the most important factors involved in the optimized design of TE scaffolds: structural properties, materials, cells and signaling molecules.

Figure 2.3 Schematic diagram of key factors in the design of optimal scaffolds for bone tissue engineering (figure modified from Gerhardt and Boccaccini6).

Scaffolds for bone tissue engineering are subjected to many interlinked and often opposing biological and structural requirements, which are summarized in Table 2.2.

Table 2.2 Design criteria for bone tissue engineering scaffolds6,19 (table modified from Chen et al.19).

1. Fabrication The material should possess desired fabrication capability, e.g., being readily produced into irregular shape of scaffolds that match the defects in bone of individual patients. 2. Porous Structure The scaffold should have an interconnected porous structure (>90%) and diameters suitable for cell penetration, tissue ingrowth and neovascularization. 3. Mechanical properties The mechanical strength of the scaffold, which is determined by both the properties of the material and the porous structure, should be sufficient to provide mechanical stability prior to synthesis of new extracellular matrix. 4. Biodegradability The composition of the material, combined with the porous structure of the scaffold, should degrade in vivo at rates appropriate to tissue regeneration 5. Osteoconductivity The material should encourage osteoconduction in the host bone. In fact, osteoconductivity eliminate the formation of fibrous capsule, but it also enhances the formation of a strong bond between the scaffold and the host bone.

6. Cell Activation The material should not be only biocompatible, but it should also stimulate cell attachment, proliferation and differentiation. 7. Commercialization Potential The synthesis of the material and fabrication of the scaffold should be suitable for commercialization

A major limit in the design of engineered scaffolds is that most materials are not simultaneously mechanically competent and bioresorbable, i.e. mechanically strong materials are usually bioinert, while biodegradable materials tend to be mechanically weak. Some examples of materials used for the fabrication of scaffolds for bone tissue engineering (BTE) are:

• Bioactive glasses and glass-ceramics, due to their capability to develop a layer of hydroxycarbonate apatite (HCA) once in contact with the biological fluids, which resembles the bone mineral phase for chemical composition and structure20. Although bioactive glasses are mechanically weak, it has been shown that 45S5 Bioglass® can crystallize when heated to high temperature, keeping the feature of bioactivity20,21; • Natural polymers, which should not cause foreign material response when implanted in human body, but they are characterized by low mechanical properties6; • Synthetic polymers, which have numerous advantages, such as excellent processing characteristics, which can ensure the off-the-shelf availability as well as being biocompatible and biodegradable at rates that can be tailored for the intended application6. In addition, components of polymer and inorganic bioactive phases are also suitable biomaterials for BTE22. Although a number of materials and fabrication techniques have been developed, several issues need to be addressed prior to clinical application, e.g. mechanical behavior, induction of vascularization and tailored biodegradability.

2.3. Classification of Biomaterials

Nowadays, it has been accepted that no foreign material placed within the human body is complete biocompatible. The only substances that conform completely are those manufactured by the body itself; any other substance induces a host-tissue response. When a synthetic material is placed within the human body, tissue reacts in different ways depending on the material type. In general there are three terms in which a biomaterial could be classified considering the tissues responses: bioinert, bioactive and bioresorbable6,20,23.

The term bioinert refers to the first generation of biomaterials that once placed in the human body have minimal interaction with the surrounding tissue. This is due to the fact that for many years, it was thought that interactions between body and implants could cause only tissue irritation, damage and sometimes death. Examples of these materials are stainless steel, titanium, alumina, zirconia and UHMWPE. Generally a fibrous capsule might form around these implants after they remain in contact with biological environment. This type of interface cannot last for a long time and can cause the failure of the implant. For this reason, research switched to the development of materials that could interact with the body6.

The term bioactive refers to a material, which upon being placed within the human body interacts with the surrounding hard tissue and in some cases with soft tissue20. The level of bioactivity of a specific material can be related to the time taken for more than 50% of the 24 interface to bond to tissue (t0.5bb) :

Bioactivity Index Ib = 100/ t0.5bb (Eq.1)

Ib is used to identify the bioactivity level of material. Materials exhibiting an Ib value greater than 8 (class A of biomaterials) will bond to both soft and hard tissues. Materials with an Ib value lower than 8 (class B biomaterials), but greater than zero will bond only to hard tissue25,26.

The term bioresorbable refers to a material that upon placement within the human body starts to dissolve and it is slowly replaced by newly formed tissue. These materials are supposed to enhance and increase the body capability of self-repairing, for this reason they represent an alternative solution to the problem of long-term implant failure of the bioinert materials.

For this class of materials it is very important that the dissolution products are biocompatible and the resorption should occur at a similar rate to tissue regeneration rate.

2.4. Bioceramics

The nature of the atomic bonding in ceramics varies from primarily ionic to primarily covalent with many ceramics having a combination of the two types of bond. The strength of the bonds determines the characteristic properties of high hardness, thermal and electrical resistance of ceramics. The high melting point of ceramics means that they cannot be easily heated to their melting temperature to facilitate the processability. The processing of powders to form solid products necessarily involves the packing together of fine powders, their consolidation and heating to form bonds. This process of diffusion and bonding is

11 called sintering process17. The sintering process takes place in three steps: i) the powders are compacted and the particles are in contact with one other but are not physically bonded (Figure 4a). ii) The compacted powders are heated up to a temperature, which is usually 2/3 of the melting temperature (Tm). During this step “necks” are formed between the particles, bonding them together (Figure 4b). The contact areas between the particles expand and the density of the compact increases and the total void volume decreases. Small diameter particles are characterized by high surface area and high free surface energy. The high surface energy of the particles is the driving force of the sintering process: thermodynamic tendency to decrease the free surface energy by bonding the particles together. Increasing the number of bonds, the free surface energy of the system decreases. iii) The particles are fully bonded together leaving a small closed porosity between them (Figure 4c). The original powder particle size will control the final pore size and distribution: the smaller the particles size the smaller the pores and the better the mechanical properties will be.

Figure 2.4 a) powder particles compacted together, b) particles beginning to bond together, c) fully sintered ceramic (figure modified from Narayan et al.17).

Bioceramics are a class of ceramic materials developed to have a specific biological behavior in the human body. The potential of bioceramics as biomaterials is due to their main feature of biocompatibility with the physiological environment. For this reason, nowadays bioceramics are widely used for numerous applications inside the human body.

A classification of bioceramics, according to the interface formed with tissue, in: almost inert, bioactive and resorbable. Almost inert bioceramics are not subjected to chemical variations after a long exposition in the physiological environment. Advanced ceramics, such as aluminium oxide (Al2O3), silicon carbide (SiC), silicon nitride (Si3N4), belong to this category. The host tissue recognizes the almost inert ceramic material as foreign body and produces a thin fibrous capsule (about 10 µm) all around the device, which prevents further

12 interaction with the surrounding tissues. This type of interface cannot last for a long time and sometimes there is the mobilization of the implant. This leads usually to the need of surgical removal of the device. This kind of materials, used as component for orthopaedic implants subjected to high mechanical stresses and in the cardiovascular field, generally as coatings, includes:

• Al2O3 , bioceramics of alumina are designed as an alternative to metallic alloys for the production of orthopaedic devices. Compared with the conventional metallic implants, they offer a low friction coefficient and a low rate of deterioration, meaning a lower number of detritus and superior tolerability by the organism. • Turbostratic carbon, which has the best biocompatibility with blood, for this reason is widely used as coating of vascular valves.

The composition of the bioactive and bioresorbable bioceramics is designed in order to have a specific reaction with the physiological environment, enhancing the formation of a chemical bond between the tissue and the implant`s surface. The main features of bioactive ceramics are the bond to bone without the formation of a fibrous tissue at the interface and an increase of the resistance to movement of the device placed inside the body. These two classes of biomaterials include: bioceramics based on calcium phosphates and bioactive glasses. Both of them try to mimic the mineral phase of bone which has a complex structure made of calcium carbonates, calcium phosphates and specially carbonate hydroxyapatite. The surface bioactive materials are designed to form a stable bond with bone and to be chemically stable in the physiological environment. Bioresorbable materials degrade in physiological environment and have a life time in the body similar to the necessary time of bone regeneration.

2.4.1. Structure of Glass

Glass is an amorphous solid inorganic material, also defined as a non-crystalline solid. The most familiar type of glass, used for centuries in windows and drinking vessels, is soda-lime glass, composed of about 75% silica (SiO2) plus sodium oxide (Na2O) from soda ash, lime CaO, and several minor additives. Glasses are obtained by the progressive solidification of a liquid that does not crystallize during the cooling. They are very hard, brittle and transparent materials.

In a crystal solid, the transition from the liquid to the solid phase happens in correspondence of a characteristic temperature (solidification temperature). In an amorphous solid (glass) this passage happens through the progressive and continuous increase of the viscosity of the

13 material during cooling. This is possible if the cooling rate is higher than the crystallization rate of the material, when the temperature has a value lower than the solidification temperature. The glass thus obtained has a spatial organization characterized by disorder. Since the glassy state is not an equilibrium state, the glass can crystallize. This process, called devitrification, consists in the formation of crystal domains. Crystallization can also happen during sintering of amorphous glass particles. It is possible to describe the formation of a glass by observing the variation of a property of the vitrifying material, for example the specific volume, as function of the temperature. During heating a crystal, its specific volume increases. In correspondence of the melting temperature (Tm) it is possible to observe a sudden increase of the specific volume (Figure 2.5 line E-B) because of the phase transition from solid to liquid. From point B the volume increases again linearly (Figure 2.5 line A). Two different phenomena are possible when cooling the obtained liquid, namely:

• If the cooling rate is slow enough, the curve of the specific volume will follow the same curve of the heating. At the end crystal is reformed (Figure 2.5). • If the cooling rate is high enough, the liquid is cooled without abrupt changes of the specific volume, passing through the point B without crystallization. At the end, the specific volume of the glass is higher than the specific volume of the crystal. In fact the structure does not have time to organize in the most ordinate way in the space. It is frozen in a disordered configuration typical of liquids, despite having reached the viscosity and the consistency of a solid. From the curve it is possible to obtain a

particular temperature, called temperature of glass transition Tg, that delimits the glassy behavior from the liquid behavior (Figure 2.5)27.

Figure 2.5 Relation between glassy, solid and liquid state. Heating a crystal its specific volume increases (line F- E) and in correspondence of the melting temperature (Tm) there is high increase of the specific volume (line E-B) because the solid-liquid transition; later the volume increase linearly (line B-A). When the liquid is cooled, if the cooling rate is slow, the curve will follow the same curve of heating (line A-B-E-F), if the cooling rate is fast the liquid passes through point B without abrupt change of specific volume and without crystallizing (line A-B-C-D). 27 Tg is the temperature of glass transition .

The most common structure of glass is the one present in the silica glass (SiO2). The 4- repetitive unit is the tetrahedral SiO4 , where every silica atoms is in the middle of a tetrahedron, whose vertices are occupied by four oxygen atoms. Every vertex is shared with another tetrahedron, in this way every oxygen atom is bond with two silica atoms. In figure 8 is represented the ordinate structure of the quartz (crystal SiO2) and the amorphous structure of the silica glass (amorphous SiO2) where the SiO4 tetrahedra are still linked together, but the angle between them is not constant and resulting structure is not well organized (Figure 2.6).

Figure 2.6 Structure of the crystalline and amorphous silica.

The constituents of a glass are:

• Forming oxides: are the silica, germanium, boron and arsenic oxides able alone to

form a glass. These oxides generate tetrahedral structures such as silica SiO2, with X-O ionic or covalent bonds.

• Deforming oxides of lattices, e.g. alkali oxide like Na2O or like CaO. These oxides do not have the capability to form a glass, but they are incorporated in the glass network formed by the glass former (Figure 2.7). In pure silica glass, every oxygen atom is linked to a silica atom. These oxygen atoms are called bridging oxygen. When a modifier is added to glass composition, cations interrupt the silica network and some oxygen atoms are not linked to silica atoms. These oxygen atoms are called non-bridging oxygen. The bond between the not bridging oxygen atoms and cations is less strong than the Si-O bond.

+ + à ≡Si-O-Si≡ + Na2O à ≡Si-O-Na Na -O-Si≡

Figure 2.7 Structure of a glass SiO2-Na2O

• Intermediate oxide (e.g. Al2O3 and PbO) are not able to form a glass network, but they can be part of the lattice replacing some atoms of the former oxide. An example

is the glass SiO2-Na2O-Al2O3 (Albite).

2.4.2. Silicate Bioactive Glasses

“A bioactive material is defined as a material that elicits a specific biological response at the interface of the material, which results in the formation of a bond between the tissue and the material”9. This definition was given by L. Hench, who discovered that certain compositions of glass in the system SiO2, CaO, Na2O and P2O5 are able to form a bond with bone and soft tissues. Indeed, when a glass is in contact with biological fluids a layer of carbonated hydroxyapatite (HCA), similar to the mineral phase of bone, is deposited on the surface of the material and ionic dissolution products are released. This layer is responsible for the strong bond between bioactive glasses and human bone. Moreover, the controlled rates of release of ionic dissolution products, especially critical concentrations of soluble silica and calcium ions, stimulate osteoblast gene expression and angiogenesis in vitro and in vivo9,14,26,28. These materials have also demonstrated anti-microbial and anti-inflammatory properties, but the mechanism of these properties has not been demonstrated yet29. All the bioactive glass properties are summarized in Figure 2.8.

Figure 2.8 Overview of biological response to ionic dissolution products of bioactive glasses (figure modified from Hoppe et al.26).

The chemical response at the interface of a material is usually subjected to a variety of variables; one of them is the bulk composition of the glass (Figure 2.9).

Figure 2.9 Compositional diagram for bone bonding. Materials belonging to the region A are materials able to bond with bone; region S is a region of class A bioactivity where bioactive glasses bond to both bone and soft tissue and are gene activation, as 45S5 Bioglass® (figure modified from Hench9).

Bioactive glasses, which belong to the region A of the compositional diagram, are able to bond with hard tissues; inside the region A there is the sub-region S, that identified the composition of bioactive glasses with the capability to bond not only with bone, but also with soft tissues. Bioactive glasses, belonging to the region B of the diagram, are able to bond only with hard tissue. Materials belonging to the region C dissolve too fast, so there is no possibility to form a bond with the surrounding tissues. The region D identifies the glass compositions that are not able to bond, neither to hard nor to soft tissues. An important feature of materials belonging to the class A is that they have shown osteoinductive as well as osteoconductive properties8,23. In contrast, class B bioactive materials exhibits only osteoconductivity. Osteoinduction means that primitive, undifferentiated and pluripotent cells are somehow stimulated to differentiate into bone-forming cell lineage23. An osteoconductive material allows bone growth on its surface28. The components of the BGs are:

• SiO2 (silicon dioxide): the SiO2 content affects the bond rate formation between bone and BG. If this content is > 60%, as in soda-lime-silica glass, the material is nearly

inert and is encapsulated by a fibrous tissue. The SiO2 content decreases the solubility rate of the other ions and thus provides stability. When the glass

composition exceeds the 52% by weight of SiO2 it will bond to bone but not to soft tissues.

• CaO (calcium oxide) and Na2O (sodium oxide): these elements ensure easy aggression of the BGs in a physiological environment if they are 25 wt.% in the BG 2+ + + + composition. Ca ions exchange with H or H3O is slower compared to Na . For

this reason, an increase of Na2O content enhances the resulting dissolution rate of BGs.

• P2O5 (phosphorous oxide): this element converts the inert soda-lime-silica glass in BGs. This is possible only if its wt.% in the glass composition is < 6%. If its rate is > 6% the material is not able to bond to bone9,25,30–32.

2.4.3. The Mechanism of Bioactivity

Once a BG is in contact with an aqueous environment, chemical and structural changes occur within the material. The accumulation of dissolution products induces a variation of the solution pH. The release of Si and Ca2+ ions from BG and the formation of HCA layer are key factors to stimulate the tissue growth and regeneration6,8,9,24,28. There are eleven stages in the process that leads to the complete bonding of bioactive glasses to bone. The first five steps are chemical; the remaining ones are related to the biological response of the surrounding tissue26:

+ 2+ + + i. Rapid exchange of Na and Ca with H and H3O from solution, causing hydrolysis of the silica groups, which creates silanols:

+ + - - + - Si-O-Na + H + OH à Si-OH + Na (aq) + OH

The pH of the solution increases because the H+ are replaced by cations.

ii. Loss of soluble silica in the form of Si(OH)4 to the solution resulting from the breakage of Si-O-Si bonds and formation of Si-OH (silanols) at the glass solution interface:

Si-O-Si + H2O à SiOH + OH-Si

iii. Condensation and re-polymerization of a SiO2 rich layer on the surface that is depleted in Na and Ca2+.

2+ 3- iv. Migration of Ca and PO4 groups to the surface through the SiO2-rich layer

forming an amorphous calcium phosphate layer (CaO-P2O5-rich film) on the surface

of the SiO2-rich layer, followed by growth of the amorphous CaO-P2O5-rich layer by incorporation of soluble calcium and phosphate from solution.

- v. Crystallization of the amorphous CaO-P2O5-rich layer by incorporation of OH 2- and CO3 anions from solution to form a mixed hydroxy carbonate apatite (HCA) layer (Figure 2.10)8,25,26.

Figure 2.10 Schematic illustration of the surface stages (i-v) reactions on bioactive glasses, forming double SiO2- rich layer and Ca-P- rich layer (figure modified from Peitl et al.26).

vi. Adsorption and desorption of growth factors in the HCA layer to activate the stem cells differentiation.

vii. Action of macrophages to remove debris to the site allowing cells to occupy the space.

viii. Attachment of the stem cells to the bioactive surface.

ix. Differentiation of the stem cells to osteoblasts.

x. Generation of extracellular matrix by the osteoblasts to form bone.

xi. Crystallization of inorganic calcium phosphate matrix to enclose bone cells in a living composite structure (Figure 2.11)33.

Figure 2.11 Sequence of interfacial reactions involved in forming a bond between bone and bioactive glass (figure modified from Peitl et al.26).

The first stages of the HCA layer formation depend on the composition of the BG; the last stages depend on the response of the surrounding tissue to the released ions. For many years it was assumed that the formation of HCA layer was the critical requirement for the bioactive behavior of BGs9,25,29,34. Nowadays, the formation of HCA layer is considered to be useful but not critical for bone regeneration. It is known that the key mechanisms are related to the controlled release of ionic dissolution products from the degrading BG, especially critical concentrations of soluble silica and calcium ions, at the rate needed for cell proliferation and differentiation9,25,35–38. The osteoprogenitor cells must receive the correct chemical stimuli from the local environment to instruct them to enter the active segments of the cell cycle6,9,38,39. Sun et al.38 showed that 45S5 Bioglass®, belonging to class A bioactive materials, promotes human osteoblast proliferation by reducing the growth cycle: in the presence of the critical concentration of Si and Ca2+ ions, within 48 h, osteoblasts that are able to differentiate into a mature osteocyte phenotype begin to proliferate and generate new bone. Moreover osteoblasts that are not in the correct phase of cell cycle and unable to proceed towards differentiation are switched into apoptosis6,9,38–40. Bioinert materials or Class B bioactive materials do not produce the local environment necessary to stimulatethe osteoprogenitor cells to proliferate and to differentiate. Only Class A bioactive materials induce a rapid formation of new bone in vivo and in vitro9,40. After few hours from the exposure of human primary osteoblasts to soluble 45S5 Bioglass®, several families of genes are activated, including: genes encoding nuclear transcription factors and growth factors, especially Insulin like Growth Factor IGF-I and IGF-II, which regulate human osteoblast proliferation9,28,40. Activation of several immediate early response genes and synthesis of growth factors is likely to modulate the cell cycle response of the osteoblasts to BG. These

21 findings indicate that Class A BGs enhance new bone formation through a direct control over genes, that regulate cell cycle induction and progression.

2.4.4. Two Classes of Bioactive Glasses

The two classes of BGs are the melt-derived and sol-gel BGs. Until the late 1980s, BGs were generally melt-derived, with the majority of researches aimed at the 45S5 Bioglass® composition and apatite-wollastonite glass ceramics10,14. This process involved dry mixing of powders, melting the glass phase in a fused silica crucible at high temperature (1350°C), followed by casting of bulk or quenching the powders. The rapid rate of HCA formation exhibited by melt-derived BGs was attributed in the past to the presence of Na2O, which increases the solution pH at the implant-tissue interface. It is widely accepted that increasing silica content of melt-derived glass decreases dissolution rate by reducing the availability of ions such as Ca2+ and HPO4- to the solution and inhibiting development of silica-gel layer on the surface of the material. The result is a reduction, and in some cases the elimination, of the bioactivity of the melt-derived BGs when the silica content is > 60%33. Nowadays, it is demonstrated that the formation of the silica-gel layer on the surface of the BGs is the key factor of the nucleation and crystallization of HCA, dispelling the theory that Na2O is the active component of the material, as shown by Li, Clark and Hench41. The only way to obtain bioactivity for silica levels > 60% is to use the sol-gel process, which is a novel processing technique for the synthesis of bioactive glasses8,9,20. Rounan, Li, Clark and Hench41 showed that it was possible to produce stable BGs by sol-gel process and the most known is the 58S gel-glass (60%SiO2, 36%CaO, 4%P2O5), which was approved by the Food and Drug Administration (FDA)33.

The sol-gel process is a chemical synthesis method where an oxide network is formed by polymerization of chemical precursors dissolved in a liquid medium. The sol-gel process involves the generation of colloidal suspensions (sols) of solid particles in a liquid. During the process, the colloids are converted to viscous gel (gelation) and then to solid materials. The most common precursors for the formation of colloidal suspensions are metal alkoxides, which consist of metal elements, such as Si, surrounded by various alkoxide groups as ligands. Then hydrolysis and condensation of metal alkoxydes from large metal oxide molecules. The hydrolysis reaction, which can be either acid or based catalyzed, replaces alkoxyde groups of organosilanes with hydroxyl groups. Siloxane bonds (Si-O-Si) are formed during the subsequent condensation reaction. Alcohol and water are products of the condensation reaction and evaporate during drying. A typical precursor for SiO2 is TEOS

(tetraethyl orthosilicate), that hydrolyses as Si(HO)4 (hydrated silicon dioxide) and ethanol

(C2H5OH). In fact TEOS has the remarkable property of easily converting into silicon dioxide and the side product is ethanol. Rates of this conversion are sensitive to the presence of acids and bases, both of which serve as catalysts42.

Si(OC2H5)4 + 2H2O à Si(OH)4 + 4C2H5OH

In parallel with Si-OC2H5 bond hydrolysis, catalysed by nitric acid, a process of re- condensation occurs between the formed silanol groups. TEP (triethylphosphate) is a precursor for P2O5 and Ca(NO3)2 4H2O is a precursor for CaO. A polymerized network is obtained, which finally becomes a gel. After gelation, the sample is kept usually at room temperature (RT), and then dried at 120-150°C, in order to eliminate water excess and the alcohol obtained from the hydrolysis. Dried gels can then be stabilized at higher temperatures. Resulting materials are amorphous, with high specific surface area due to the formation of mesopores33,43.

There are several advantages of a sol-gel derived glass over a melt-derived one, that are important for making tissue engineering scaffolds. Sol-gel derived glasses have:

• Lower processing temperature, 600-700°C instead of 1350°C of the melt-derived process; • Improved homogeneity of the glass compostion;

• Wider composition can be used (up to 90% SiO2), maintaining the bioactivity; • Better control of bioactivity by changing composition or microstructure; • Structural variation can be produced without compositional changes; • A higher bioactivity in vitro; • Higher specific surface area.

The mechanism for dissolution and HCA formation of the sol-gel bioactive glasses follows the same eleven stages of the melt-derived glasses listed in the previous paragraph. In vitro studies show a more rapid dissolution and faster HCA formation for sol-gel derived 58S gel- glass compared with 45S5 Bioglass®: the 45S5 implant exhibits a loss of sodium ions, but has a lower rate of silica dissolution and a thicker HCA layer. During the first hours of implantation silica is lost at the similar rapid rate for both materials, after that time the silica dissolution rate increase for the 58S gel-glass. Similar results are found in vitro, but in vivo results show no difference in bone formation between the two glasses after twelve weeks. The increased rate of HCA formation and higher index of bioactivity for the sol-gel derived glasses is due to the more release of soluble silica that nucleates the HCA crystal810,33.

2.4.5. Clinical Applications of Bioactive Glasses

The 45S5 Bioglass®, since when it was approved by the FDA, was widely used in several clinical applications and in many different forms20,36,44,45. So far, 45S5 Bioglass® has been used in more than a million patients to repair bone defects10,14,46. E brief summary of clinical applications of BGs is reported below:

• Monolithic medical devices The first successful use of 45S5 Bioglass®, in 198410,14,34, was a device used to replace the bones of the middle ear to treat conductive hearing loss. The implant was designed to replace the bones and transmit sound from the eardrum to the cochlea, restoring hearing. Materials previously used to treat this disease were inert metals and polymers. However, they failed because thy engendered a fibrous tissue reaction which effectively holds the implant in place, but scar tissue around an implant will dampen, rather than transfer sound waves. Consequently these implants became gradually less efficient. These issues were overcome by using a prosthesis made of 45S5 Bioglass® (middle ear prosthesis, MEP®. The advantage of the MEP® over other devices in use at the time was its ability to bond with soft tissue (tympanic membrane) as well as bone tissue: it bonds with the remainder of the stapes and with the tympanic membrane. After 10-year follow-up studies, 4 out of 21 failed due to fracture, but the others retained function42. The second commercial 45S5 Bioglass® device was commercialized in 1988, the Endosseous Ridge Maintenance Implant (ERMI®)10. It was a simple cone shaped 45S5 Bioglass®, inserted into fresh tooth extraction sites to repair tooth roots and to provide a stable ridge for dentures10. These products are not extensively used by surgeons, because they need to cut the implant in shape rather than use cones of fixed sizes. Another successful application of monolithic 45S5 Bioglass®, was reported by Thompson et al.10. He performed clinical trials on trauma patients with orbital floors that were so severely damaged that their vision was affected. Using a computed axial tomography, the defect site was scanned and, using a rapid prototyping machine, mould for casting the bioactive glass implants was produced. At 5-years follow-up, the patients regained full movement of their eyes.

• Particles and granule Products based on BGs and used worldwide are those based on particles. In fact, dentist and orthopaedic surgeons often prefer to use particles or granules, because

they can be easily pressed into the defect site. The first 45S5 Bioglass® particulate product was PerioGlas®, commercialized in 199310,34. It is successfully used to regenerate bone around the root of a healthy tooth to save the tooth, or can be used to repair bone in the jaw so that the bone´s quality is sufficient for anchoring titanium implant. Following the success of BG particles for dental bone repair, NovaBone (NovaBone Bone Products LLC) was released in 1999 for orthopaedic bone grafting in non-load bearing sites10. It is usually mixed with patient´s blood until it reaches a putty-like consistency due to the blood´s coagulation, and finally pushed into the defect site. Clinical trials with NovaBone were reported for posterior spinal fusion operations for treatment of adolescent idiopathic scoliosis. In a 4-year follow-up, NovaBone performed as well as autograft but with fewer infections (2% vs 5%) and fewer mechanical failures (2% vs 7.5%)47. In 2006, BonAlive® (composition S53P4), received the European approval for orthopaedic use as bone graft substitute. This bone graft substitute has been applied in different clinical trials, some examples are: i) implantation of BonAlive® granules in the maxilla (upper jaw, consists of porous cancellous bone) with autologous bone allowing the implantation of titanium roots in the porous maxilla; ii) treatment of severe spondylishtesis (displacement of the vertebral column) using BonAlive® and auograft44,45,48; iii) treatment of subchondral bone defects49; iv) repairing of bone defects caused by benign bone-tumor surgery in hands, tibia and humerus50. More clinical data are available for BonAlive® compared to 45S5 Bioglass. The clinical results are good, but one drawback of BonAlive® is that its degradation rate is slower than ideal10. ® ® In 2004, a very fine 45S5 Bioglass particulates (D50 ~ 18 µm) called NovaMin has been added to toothpaste for treating tooth hypersensitivity. This toothpaste is now available in 20 different countries10. In vitro tests showed that 45S5 Bioglass® particles are able to attach to dentine, explaining how the glass particles can stimulate long-term repair even though may only be for few minutes51,52. The success of NovaMin® leaded to trial with sol-gel derived glasses.

Overall 45S5 Bioglass® is available in a range of compositions which are able to bond to soft tissues and/or bone. It has been successfully applied as solid and particulate form and may be combined with other materials, both natural and synthetic, to provide treatments for many disparate clinical conditions. What are still missing are commercialized highly porous scaffolds for load-bearing applications or suitable bioactive glass coatings to prevent the fibrous tissue encapsulation of metallic implants.

2.4.6. Production of Porous Bioactive Glass Scaffolds

As shown in the previous section, particulate BG systems have been successful in regenerating bone defect or to treat tooth hypersensitivity and different products have already been commercialized. The drawback of particulate systems is the lack of dimensional stability once placed into the defect site. A bone defect cavity can hold the particles in place until they are integrated with the new-formed tissue, but in some clinical applications there is no bony chamber and additional fixation material are needed6,10,19. For all these reason, an ideal synthetic bone graft is a porous structure that can act as a temporary template for bone ingrowth. An ideal BTE scaffold should have a highly interconnected porous structure, with pores in a range 200-500 µm, which are necessary for bone integration, neovascularization and nutrient delivery53,54. Moreover, pores in a range 0-200 µm play an important role in the complete colonization of the scaffolds by cells, enhancing the flow of biological fluids also in the inner core of the of the porous structure54. Numerous techniques have been reported for the production of 3D BG-based scaffolds:

• Foam Replica Method The foam replica method was one of the first manufacturing techniques developed for the production of ceramic materials with controlled porosity. The first patent was deposited in 196355 and, despite its age, it is still one of the most widely used techniques in industry. This technique was initially used to produce hydroxyapatite scaffolds and in 200615, it was proposed for the production of BG-based scaffolds. The produced scaffolds are a positive replica of an open cell porous template which is usually polyurethane (PU)15. After one coating with a BG slurry, the template is burned out and then the BG sintered at high temperature. The resulting 3D structures are characterized by a high open and interconnected porous network which mimics the architecture of the trabecular bone 15,56,57 however at the expense of a relatively compressive strenght6. Moreover, it is possible to obtain foams with graded porosity58,59 or, using different templates including marine sponges, scaffolds with increased mechanical properties60. It is also possible to produce scaffolds via foam replica starting from a sol-gel glass synthesis solution61. Cabana-Polo et al.62 applied the electrophoretic deposition of a sol-gel solution in combination with the foam replica method, accelerating the scaffold production time. BG-based scaffolds developed via replica method can also be coated with ordered mesoporous silica particles63,64, biodegradable polymers65 and polymer microspheres66 transforming them in local drug delivery systems and leading to enhanced mechanical properties.

• Direct Foaming Technique Highly porous BG-based scaffolds can be produced by directly foaming a colloidal sol or a powder suspension, followed by a solidification treatment67. The resulting foams exhibit a hierarchical structure with interconnected macropores (ranging from ∼20 µm to 1-2 mm24). To prevent the foam struts from collapsing, a surfactant is used to stabilize the bubbles. One of the most successful foaming method is gel- casting67, in which the in situ polymerization of monomers can be initiated foaming a 3D polymeric network, a gel, which produces stronger green bodies24. The samples are then sintered to provide mechanical strength. A further development of this technique is sol-gel foaming67, in which a sol-gel glass solution is prepared and a surfactant is added, the foam is generated by vigorous agitation and the sol is transferred to a mold for aging and finally heated treated67. A more recent technology involves the use of polymer-derived ceramics68,69. In this approach, metal oxide precursors are added to a polymeric precursor, e.g. a silicon resin, allowing the production of silicate bioceramics. The foaming is obtained by water release from specific hydrate fillers. Another approach to produce porous BG-based scaffolds involves a powder metallurgy technique70. The National Research Council of Canada NRC-CNRC has recently developed a powder metallurgy approach for the synthesis of titanium foam71. In this method, BG powder is dry mixed with a solid polymeric binder and a foaming agent. The mixture is then moulded and heat-treated72. The resulting scaffolds showed compressive strength in the range 5-40 MPa, i.e. in the upper range of values reported so far for these types of porous materials. The scaffolds exhibited highly interconnected pore structure and tunable porosity (55- 77%) and they retain good compressive strength also after one month in simulated body fluid73.

• Freeze-Drying Another possibility for the preparation of porous scaffolds generally involves freezing a ceramic slurry inducing the formation of ice crystals along the freezing direction and agglomeration of the BG particles between the crystals74. The obtained structure undergoes sublimation for the ice removal. The obtained green body is then thermally treated to consolidate the structure75. The technique has been mainly used for the production of porous polymeric composite scaffolds containing BG as filler22,76.

• Additive Manufacturing Technique Additive manufacturing technique is the standardized name given to all those processes where 3D structure are fabricated layer by layer without any specific tooling to obtain the targeted geometry77,78. These methods allow customizing the final porous scaffold based on the patient-specific defect using a CAD file10,78. The Stereolitography ceramic manufacturing78–80 starts with a layer-by-layer build-up of a photo-reactive polymer highly filled with BG powders. The 3D printing has mainly been used for the production of BG/ceramic composite scaffolds. A liquid binder is printed onto a powder bed layer-by-layer, gluing the powder together in the desired areas until the designed 3D structure is completely printed78,81,82. The dispense plotting technique is also known as direct ink writing or robocasting83. It is a method in which a paste-like material is extruded through a nozzle onto a building platform.

2.5. Ordered Mesoporous Silica Materials Based on IUPAC, porous materials according to the pore diameter can be classified into three categories: those with pores diameter less than 2 nm are microporous, pore size between 2 and 50 nm are mesoporous; and pore diameter greater than 50 nm are called macroporous materials. Ordered mesoporous materials with ordered arrays became a hot research topic in 1992, when Mobil Oil Corporation (Mobil) scientists developed for the first time the M41S series of ordered mesoporous silica materials42,43,84. Long-chain cationic surfactants were used as structure-directing agent to synthetize ordered mesoporous (alumina-) silicate materials. This approach was not a new method; twenty years earlier a group of French scientists developed a protocol to tune the density of silica gels by using long-chain cationic surfactants42. The production of these materials combine the sol-gel process of metal oxides, describe in a previous section, with cationic surfactants, as templating agents to form ordered structures. The resulting silica based materials are unique materials characterized by an ordered mesostructure of porous and disordered arrangement at atomic level (amorphous materials). They are produced using surfactants (amphiphilic organic molecules) as template to direct the assembly and subsequent condensation of the inorganic precursors, leading to a network of cavities arranged periodically. Two mechanisms have been proposed to describe mesoporous silica material formation. The first model describes the addiction of silicate to micelles formed using n- decyltrimethylammonium bromide (CTAB). In this way, the silica precursor polymerizes around the already formed micelles. The second proposed mechanism is that the addition of the silica precursor to an aqueous n-decyltrimethylammonium bromide solution induces the ordering of silica-encased surfactant micelles simultaneously. In this case, the micelles

28 formation requires the silica precursors to be present. The two mechanisms for the formation of silica ordered mesoporous materials are schematized in Figure 2.12.

Figure 2.12 Two different mechanisms proposed for the production of silica ordered mesoporous materials. Observe that the main difference occurs when adding the silica precursor (TEOS). After the surfactant has added, the supramicellar phase is formed (route 1), or at the same time as the micellar mesophase in being formed (route 2). This figure has been adapted from Vallet-Regí et al.43.

These silica ordered mesoporous materials are characterised by abundant silanol groups on their surface. These silanol groups offer an anchoring point for organic functions (chemical property) via weak interactions, such as van der Waals forces or hydrogen bonds43,85–88. All these properties are essential in the loading and release of functional molecules88. Nevertheless, silanol groups also give the possibility to functionalize the mesoporous walls with different organic groups via sol-gel chemistry89. To proceed with this functionalization, two main routes are possible. One is the post-synthesis, including sylation or grafting, the other is the co-condensation43,85. The post-synthesis functionalization process is performed by grafting the organic functions to previously formed pure inorganic silica matrix, i.e. by reaction of organosilanes (R´O)3SiR with the free silanol groups of the pore surface, under anhydrous conditions43,85 (Figure 2.13). In this way it is possible to retain the mesostructure of the starting material, although there might be a reduction in the porosity due to the presence of the organic moiety. The main disadvantage of this pathway is that some organosilanes preferentially react at the pore entrance, which may reduce the entrance of

29 further organosilanes inside the pore channels. This could lead a non-homogeneous distribution of the organic groups inside the pore walls43.

Figure 2.13 Schematic representation of the post-synthesis functionalization method of ordered mesoporous materials. This figure has been adapted from Vallet-Regí et al.43.

The co-condensation method has the simultaneous condensation of the corresponding silica and organosilica precursors, (R´O)3SiR, in the presence of the structure-directing agent during the mesoporous synthesis and all the functionalization process can be carried out in one step43,85 (Figure 2.14). Particular attention should be taken during surfactant removal to avoid destroying or degrading the organic functionalization. For this reason solvent extraction methods should be used instead of calcination. The main advantage of the co- condensation method is the more homogeneous distribution of the organic functionalities inside the pore walls, because the organic units are direct components of the ordered mesoporous silica matrix. At the same time an increase of the organosilanes amount in the mixture reaction could affect or totally destroy the mesoporous order. Also, by increasing the organic functionalities, an important reduction of the pore diameter, pore volume and specific surface area has been observed43.

Figure 2.14 Schematic representation of the co-condensation functionalization method of ordered mesoporous materials. This figure has been adapted from Vallet-Regí et al.43.

The maximum achievable degree of organic modification varies with these two functionalization routes. The co-condensation method, leading to disordered products, does not achieve more than 40 mol.% of organic functionalities content in the modified silica phase43,89. Contrarily, the post-synthesis method enables higher functionalization rate, but not homogeneously distributes. The functionalized mesoporous silica materials containing functional groups will interact with the guest molecules via attracting interactions (electrostatic attractive interactions, hydrophilic-hydrophobic forces or electronic interactions), allowing a fine tuning of the drug loading and release kinetics88.

2.5.1. Ordered Mesoporous Silica Nanoparticles After the development of MCM-41 materials by the Mobil Oil Corporation42, in 1997 Grün et al.90 described the first synthesis of spherical MCM-41 nanoparticles. The synthesis procedure was a modification of the Stöber´s reaction91 for the preparation of monodispersed silica spheres. The Stöber´s reaction involves the hydrolysis of tetraalkoxysilane (e.g. TEOS) in a mixture of low boiling alcohol and aqueous ammonia91. Grün et al.90 modified this procedure by adding a cationic surfactant to the synthesis solution, those providing a source of micelles, which play a crucial role during the formation of MCM-41. After the reaction the surfactant was removed by calcination at 550°C90. The pores are distributed in a 2- dimensional hexagonal symmetry. MCM-48 has a similar pore size as MCM-41, but the pores show cubic symmetry. Few years later, ordered mesoporous silica nanoparticles with larger pore size (4.6 – 30 nm) were produced at the University of California, Santa Barbara. The material was named Santa Barbara Amorphous type material, or SBA-1592. These

31 particles also have a hexagonal array of pores. A schematic representation of the most used and investigated ordered mesoporous silica particles is reported in Figure 2.15. The researchers, who developed these types of particles, planned to use them as molecular sieves. Today, mesoporous silica nanoparticles have many applications in medicine, biosensors, thermal energy storage and imaging.

Figure 2.15 Schematic representation of some of the most used and investigated ordered mesoporous silica particles with their respective pore sizes.

2.5.2. Ordered Mesoporous Bioactive Sol-gel Glasses Due to their good biocompatibility, osteostimulating properties, high degradation rate and angiogenic potential bioactive glasses play a central role as bone substituting materials93–97. In 2004, a new family of sol-gel glasses with tailored porosity at the nanometer scale was developed by Yan et al.42,98 who combined the supramolecular chemistry of surfactants and the sol-gel method. These materials present the same composition of glasses but with designed mesoporosity, opening a new direction for applying nanothechniques to regenerative medicine by coupling drug delivery with bioactive material85,93,99. These materials, based on a CaO-SiO2-P2O5 composition, are characterized by highly ordered mesopore channel structure with a pore size ranging from 2 to 50 nm, and they have attracted increasing attention due to their potential applications86,93100–103. Since the first synthesis of mesoporous bioactive glasses (MBGs), the research in tissue regeneration has been accelerated. Their large surface area, controlled mesostructure, well defined surface properties modifications, tunable pore size and volume enhance the bioactive behavior and allow them to be loaded with biologic or pharmaceutical agents to promote new bone

32 formation86,87,99,104–108. In addition, they can be used in biomedical applications e.g. as scaffolds for bone tissue engineering43,109,110 and drug delivery systems69,71,92,99–103. During the past years, the drug delivery systems of mesoporous materials have experienced a notable progress86,111,112. Vallet-Regí and co-workers evidenced that mesoporous silica is a particularly attractive material considering its ordered mesoporous structure, good biocompatibility, low cytotoxicity, tailored surface charge and wide range of organic functionalization. They introduced clearly the concept of having a new class of biomaterials, combining good bioactivity and efficient drug delivery properties43.

The preparation of MBGs is similar to that for mesoporous SiO2 in which the supramolecular chemistry has been incorporated into the sol-gel process42,43,93. In this strategy, the incorporation of a non-ionic block copolymers as structure-directing agents (e.g. Pluronic® P123 and F127) is essential for obtaining well-ordered structures. The formation mechanism of the hierarchical MBGs can be described as follow. The initial sol is obtained by a homogeneous solution of soluble bioactive glass precursors (e.g. tetraethyl orthosilicate (TEOS), triethyl phosphate (TEP)), a surfactant and a strong acid, mostly hydrochloric acid or nitric acid, in an ethanol-water solvent. These surfactants self-organize in micelles, which are able to link the hydrolysed precursors (TEOS, TEP) through the hydrophilic component and self-assemble to form an ordered mesophase. Then the mixture reaction system of sol- gel glasses and surfactants undergoes an evaporation-induced self-assembly (EISA) process. During this time the surfactant assemble into micelles that maintain the hydrophilic parts in contact with the hydrolysed silica precursor while shielding the hydrophobic parts within the micellar core. Once the mixture is completely dried and the surfactant removed by calcination, a well-ordered mesoporous structure will be obtained, exhibiting high surface area and porosity43. Since temperatures are typically in the range from 600 to 700 °C, most disadvantages of processing at high temperature thermal treatment of glass and glass- ceramic can be eliminated. Moreover, it also offers potential advantages for the powder production due to enhanced and better control of bioactivity by changing either the composition or microstructure through processing parameters43. A schematic representation of the synthesis of MBGs through the EISA method113 is reported in Figure 2.16.

Figure 2.16 Schematic representation of MBGs synthesis through the EISA method (picture modified from Hum et al.113).

In the past few years, different therapeutic ions have been incorporated into MBGs by in situ synthesis method. In this process, the structure-directing agent is firstly dissolved in ethanol solution, and then the therapeutic ions (e.g. Cu2+, Ag+, Li+ among others) and the glass precursors are added. Then the mixture reaction undergoes EISA process and calcination. The resulting material is therapeutic ion-doped MBGs. In this way the ions are part of the glass framework and the ordered porous structure could be maintained. These ions, once released, have a significantly functional effect on the osteogenesis, cementogenesis, angiogenesis and anti-bacterial activity.

2.5.3. Ordered Mesoporous Silica Materials as Drug Delivery Systems Nowadays, the most popular ways for drug intake are oral administration and injection. However, these methods show a lack of efficiency especially since the release of the drug is not targeted to the area that needs to be treated86. For all these reasons the development of local drug release systems which enable controlled release kinetics has increased considerably during the past few years and it represents currently an important market for the industrial sector114–117. In this context, the combination of bioactive scaffolds with local drug delivery carriers is gaining increasing research efforts in the bone tissue engineering field127. Several matrices have been tested so far, such as organic polymers, organic–inorganic hybrid materials, bioactive glasses and ceramics. One approach gaining increasing interest involves obtaining drug carriers that are structured at the nanoscale. Since 1992, when Mobil Oil

Corporation synthesized the silica-based MCM-41, highly ordered mesoporous materials have attracted the attention of many scientists and in 2001 they were proposed for the first time as drug delivery system92. The main factors controlling the molecule uptake release capability of these materials are:

• Pore Size The dimensions of both the molecule to be encapsulated and the pore diameter should be considered before to decide which matrix should be employed. When the molecule is larger than the pore diameter, the physical absorption will take place only on the external surface of the ordered mesoporous material. On the other side, if the molecule is smaller than the pore, the adsorption will take place in both the external surface and inner part of the ordered mesoporous material. The pore size is also affecting the release rate: a decrease on the pore size leads to a decrease of the release rate43.

• Specific Surface Area The surface area of the mesoporous materials is a key parameter since the physical absorption of a molecule is a surface phenomenon, in which only molecules that are in contact with the silica pore wall will be retained. The bigger is the surface area, the higher is the amount of loaded molecules43. The release of the molecules would also be affected by the surface area. When the surface area is very high, there is a slower drug release in comparison with materials with smaller surface area. This happens because on their way out, the molecules would find more available area that can interact with them, promoting extra host-guest interactions that slow down the kinetic release.

• Pore Volume Pore volume has an important influence on the drug-loading capabilities: higher is the pore volume, higher is the amount of biomolecules that can be loaded in the mesostructure.

• Functionalization Molecular adsorption as drug loading is a surface phenomenon. The retention of molecules on the surface and inside the pores is governed by the chemical interactions between Si-OH groups from the silica network and the functional groups from the guest molecules. To modify this interaction, a modification of the

silica walls is necessary. The high density of silanol groups within ordered mesoporous materials makes possible to modify them as described in a previous section. In this way the chemical properties of the host matrices can be tuned at convenience, so there would be a stronger host-guest interaction. This would mean a greater drug adsorption and more sustained release. In fact, the functionalization influence has been observed to be more important than the surface area effect, the mesopore diameter effect and volume influence, and even the mesostructure43.

3. Characterization Techniques 3.1. Light Microscope Scaffolds images were taken with light microscope M50 Mikrosysteme Vertrieb GmbH. The image acquisition was done with Leica Application Suite LAS V3.8 software.

3.2. Scanning Electron Microscopy (SEM) - Energy dispersive X- ray spectroscopy (EDS)

All the samples prepared during this research project have been analysed by mean of SEM- EDS. Scaffolds were on a sample holder with conductive silver paste (Plano GmbH, Wetzlar). Particles and powders were dispersed in ethanol and few drops were put on the sample holder and let them dry under the hood for a few minutes. Different SEM-EDS machines have been used in the past years.

During the first month SEM micrographs were taken using SEM (LEO 435 VP Electron Microscopy Ltd, Cambridge) at WW5 Institute in Erlangen.a

Most of the images reported in the present work have been taken with SEM-EDS (Auriga 0750 from ZEISS) equipped with EDS detector (X-MaxN Oxford Instruments, UK). The EDS analyses were done at a fix working distance of 6 mm and a voltage up to 30kV.

SEM images of scaffolds culture with cells have been taken with EVO 50 SEM from ZEISS) coupled with an energy dispersion x-ray spectroscope (EDS, Oxford INCA 200) at Politecnico di Milano in Italy at the Department of Chemistry, Materials and Chemical Engineering “G. Natta”.b

3.3. High Resolution Transmission Electron Microscopy (HRTEM)

For TEM observation, samples were dispersed with ethanol on a lacey carbon film. In order to observe the mesoporosity of the samples, a Phillips CM30 operating at an acceleration

a SEM micrographs at Institute of Polymer Materials (WW5) were done by Dr. Judith A. Roether and Dirk Dippold. b SEM micrographs at Politecnico di Milano were done by Dario Picenoni and Monica Moscatelli.

39 voltage of 300 kV was used and images were taken using EM4 software. This machine was at WW9 Institute in Erlangen.c

HRTEM in a FEI Tecnai G2F30 S-Twin microscope (0.2 nm point resolution) operating at 300 kV and equipped with a HAADF Fischione detector (0.16 nm point resolution) and an INCA X-Max 80 silicon drift detector (SDD) for Energy-Dispersive X-Ray (EDS) was also used. This machine belongs to Institute de Ciencia de Materiales de Sevilla (CSIC- Universidad de Sevilla) in Spain.d

3.4. Micro Computed Tomography (micro-CT)

Micro-CT analysis was used to determine the pore interconnectivity of the fabricated scaffolds. Two different micro-computed topographies were used within this work.

Micro-CT scanner (Skyscan 1147 Micro-CT, Bruker). The pore interconnectivity was evaluated using CTan image analysis software. The machine belongs to Politecnico di Torino in the Institute of Materials Physics and Engineering, Applied Science and Technology Department in Italy.e

The scaffolds were also scanned with a high resolution micro-CT scanner (Skyscan 1172 Micro-CT, Bruker, Belgium). Scanning was carried out at a resolution of 5.15 µm/pixel over 360° rotation with a step of 0.4°. Cell size was calculated using a volume-based approach with the software CT-Analyser (1.1.13, Skyscan B.V., Kontich, Belgium). The machine belongs to the Institute of Glass and Ceramics at the University of Erlangen-Nuremberg.f

3.5. X-Ray Diffraction (XRD)

Two different diffractometers were used in the present work in order to measure low- and wide-angle X-ray diffraction.

Bruker D8 Discover diffractometer (Bruker, Karlsruhe, Germany) with Cu Kα radiation at 40 kV and 30 mA. The acquisition was performed in a range 10 – 80 two-theta with a step c HRTEM images at Institute of Micro- and Nanostructure Research (WW9) were taken also by Dr. Ana M. Beltrán and Anahí Philippart. d HRETM images at the Institute de Ciencia de Materiales in Spain were taken by Dr. Ana M. Beltrán. e Micro-CT analyses at Politecnico di Torino were conducted by Dr. Giorgia Novajra f Micro-Ct analyses at the Department of Glass and Ceramics in Erlangen were done by Dr. Tobias Fey.

40 size of 0.014° s-1. The machine is property of LFG department at the University of Erlangen- Nuremberg.g

Small angle XRD analysis (SAXRD) were carried out using Philips Xpert Diffractometer at Politecnico di Torino in the Institute of Materials Physics and Engineering, Applied Science and Technology Department in Italy. Diffraction data were recorder between 1 and 10 two- theta at an interval of 0.02° s-1.h

3.6. Fourier Transform Infrared Spectroscopy (FT-IR)

FT-IR pellets were produced mixing 300 mg of potassium bromide (KBr, Spectroscopy Grade, Merck, Germany) with 3 mg of sample. The two powders were dry-mixed and compressed into pellets with an electro-hydraulic press applying a force of 70 kN. FTIR (Nicolet 6700, Thermo Scientific, Germany), using KBr pellets and 32 scans at a resolution of 4 cm-1, which were repeated over the wavenumber range of 4000 – 400 cm-1. The machine is property of the Institute of Polymer Materials at the University of Erlangen-Nuremberg.i

3.7. Nitrogen Sorption Analysis

The Nitrogen sorption analysis was done to evaluate the specific surface area (BET method) and the pore size distribution (BJH method) of ordered mesoporous materials. Three different machines have been used within this project.

Analyses were done with Quantachrome Autosorb Instrument at 77 K. Prior to the measurements, the samples were autogassed for 12 h at 300°C under vacuum conditions. The machine belongs to the Institute of Chemical Reaction Engineering at the University of Erlangen-Nuremberg.j

Analyses were also done with a Quantachrome Autosorb Instrument Micromeritics Tristar II at 77 K. Prior to the measurements, the samples were outgassed for 24 h at 300 °C under

g XRD analyses at the Institute of Particle Technology (LFG) in Erlangen department were conducted by Kai Zheng, Barbara Myszka and Nicoletta Toniolo. h SAXRD analyses at Politecnico di Torino were conducted by Dr. Giorgia Novajra and Lucia Pontiroli. i FT-IR analyses were done by Dr. Liliana Liverani at Institute of Polymer Materials (WW5) in Erlangen. j Nitrogen sorption analyses at the Institute of Chemical Reaction Engineering in Erlangen were done by Dr. Alexandra Inayat.

41 vacuum. The machine belongs to the Institute de Ciencia de Materiales de Sevilla (CSIC- Universidad de Sevilla) in Spain.k

3.8. Inductively Coupled Plasma Atomic Emission Spectroscopy (ICP- OES)

The ICP-OES analyses were done in order to evaluate the ion release form the bioactive glass based scaffolds and the ordered mesoporous materials prepared in this work. The analyses have been done in two different institutes. During the bioactivity test, unchanged SBF was kept to determine the release of different ions, Ca, P, Si, Sr, Cu, Co and Ag. The SBF alone was also considered as a sample.

Inductively coupled plasma optical emission spectroscopy Optima 8300 Perkin Elmer (USA) was used at at Politecnico di Milano in Italy at the Department of Chemistry, Materials and Chemical Engineering “G. Natta”. Absorption wavelengths used were: Ca (317.933, 315.887 and 393.366) nm, P (213.617, 214.914, 178.221) nm and Si (251.611, 212.412, 288.158) nm.l

Concentration determination was performed also with another Optima 8300 (Perkin Elmer, USA), property of the Institute of Particle Technology, Department of Chemical and Biological Engineering at the University of Erlangen-Nuremberg. A sample flow rate of 1.5 mL min-1, and an Argon-plasma power of 1500 W were applied. The flow rates of plasma

(Ar)-, nebulizer (Ar)-, and auxiliary (N2) gas were 10 L/min, 0.6 L/min and 0.2 L/min, respectively. At least five point calibrations were performed for quantification of the elements using solutions derived from 1 g/L standard solutions (for ICP-OES, 1000 mg/L, Carl Roth, Germany).m

3.9. Compression Test

Compression test were performed on cylindrical Bioglass scaffolds. The test were performed using a Zwick Z050 mechanical tester at a crosshead speed of 1 mm/min, with a perforce of 0.1 N with load cells of 50 N and 1 kN. The load was applied until densification of the

k Nitrogen sorption analyses at Institute de Ciencia de Materiales in Spain were taken by Dr. Ana M. Beltrán. l ICP-OES analyses were done at Politecnico di Milano in Italy by Dr. Simone Gelosa. m ICP-OES analyses were done at Institute of Particle Technology (LFG) in Erlangen by Dr. Jochen Schmidt.

42 samples started to occur. The compressive strength was defined as the maximum stress of the linear elastic part of the stress-strain curve. The machine was property of the Institute of Polymer Materials at the University of Erlangen-Nuremberg.

4. Scaffolds Production and Characterization

4.1. Introduction

Bioactive glass of 45S5 composition (45SiO–24.5Na2O–24.5CaO–6P2O5 in wt.%) was the first BG developed and today it is still the most frequently investigated BG for bone defect repair10. 45S5 BG has been used in several clinical applications in bulk form and as particulate for bone grafting and for the prevention of dental hypersensitivity. Jones10 has recently summarized the different traditional BG applications in orthopedic, dentistry and bone regeneration whilst emerging applications in soft tissue engineering have been reviewed elsewhere118. The first 3D BG-based scaffolds based on 45S5 BG composition were introduced in 2006 by Chen et al.15, who used a standard foam replica technique with polyurethane (PU) foam as template to fabricate scaffolds of porosity >90%. Despite their high bioactivity15 and vascularization potential119, these scaffolds do not have sufficient mechanical strength to be used in load bearing bone sites. Following the first study on fabrication of BG based scaffolds by the foam replica method; other techniques have been proposed to make such scaffolds, including sol-gel casting24,120, freeze drying121, 3D printing122 and powder metallurgy70,72,73. Porous 3D BG based scaffolds are not yet available for clinical applications although numerous techniques have been reported for the production of porous ceramics. The main difficulty is the fabrication of a scaffold that has, at the same time, high porosity, sufficient mechanical stability and fracture strength similar to that of natural bone123. One approach is the manufacture of BG based scaffolds with improved mechanical properties due to a reduction of the volumetric porosity without affecting the pore interconnectivity and maintaining adequate pore size for new tissue ingrowth and vascularization. More recently, 3D BG based scaffolds have been produced using a powder technology process70. Aguilar-Reyes et al.72 showed that foams with interconnected porosity (64-79%) and with open pores in the 100-800 µm range can be produced with this process. The foams produced with this technique were structurally robust showing compressive strength values comparable with those of cancellous bone (2–12 MPa) and exhibiting high pore connectivity72,73 and stable mechanical properties also after 28 days in SBF73. The achieved values72,73 are notable higher than those obtained by the polyurethane (PU) sacrificial template method15, which however exhibit higher porosity. The replication of PU sacrificial template produces a porous structure that closely resembles that of cancellous bone, highly porous and interconnected structure up to 95% porosity. However, the relatively low mechanical properties limit its suitability for load-bearing applications.

Cunningham et al.124,125 used for the first time marine natural sponges as sacrificial template for the production of hydroxyapatite based scaffolds in place of PU foams. They were able to obtain scaffolds with high open porosity and improved mechanical properties. Additionally, these scaffolds had pores in the range 0-200 µm, which have been reported as being necessary for the complete integration of a bone substitute in natural tissue53,54,126. However, to the authors´ knowledge, no further developments have been reported on this material. In the present chapter the production of BG based scaffolds using marine natural sponges, Spongia Agaricina (SA) and Spongia Lamella (SL), as template materials is reported60. These sponges, thanks to the millenarian evolution for water filtration127, could be potential precursors in the production of bone tissue engineered scaffolds due to their efficient interconnected porous architecture. The final aim was thus to obtain BG based scaffolds with reduced total porosity and, consequently, increased mechanical properties, without the loss of interconnectivity of the pores essential for successful bone ingrowth. An in depth study on this novel family of BG scaffolds based on marine natural sponges is reported. In particular, the effect of pore structure on the bioactivity128, mechanical stability128 and oxygen diffusivity129 in SBF was investigated. Results of first cell culture studies using Saos-2 osteoblast on the scaffolds are also presented128. A comparison between these novel scaffolds and those produced using PU foam as template is also presented60,128,129. The results presented in this chapter have formed part of previous publications60,128,129.

4.2. Materials and Methods

4.2.1. Scaffold Production

The starting material was melt-derived 45S5 Bioglass® powder (Vitrexx, Schott, Germany, nominal mean particle size < 2 µm). The sacrificial templates for the scaffold production were PU packaging foams (45 ppi) (Eurofoam Deutschland GmbH Schaumstoffe), marine sponges Spongia Agaricina (SA), Spongia Lamella (SL), belonging to the “Elephant Ears” family, harvested respectively from the Indo-Pacific Ocean (Pure Spnges, UK) and Mediterranean Sea (Hygan Products Limites, UK). Harvesting of all natural sponges has been performed in an environmentally friendly manner, as specified by the supplier companies. BG scaffolds were produced by the replica technique, according to the method originally developed by Chen et al.15. Briefly, BG slurry was prepared by dissolving polyvinyl alcohol (PVA, Merck KGaA, Germany) in deionized water at 80°C for 1h, the ratio being 0.01mol L-1. Then BG powder was added to 25 mL PVA-water solution with concentrations of 40 wt.%. Each procedure was carried out under vigorous stirring using a

46 magnetic stirrer for 1h. The sacrificial templates, cut into cylinders, were immersed into the slurry for 10 min. The natural sponges were previously immersed in deionized water until they were completely wet in order to facilitate the slurry impregnation. The foams were then retrieved from the glass suspension, and the extra slurry was manually squeezed out. The samples were then dried at room temperature for at least 12 h. This dip coating procedure was repeated two or three times to increase the coating thickness and, consequently, the mechanical properties of the scaffolds. After the second and third coating, the superfluous slurry was completely removed using compressed air as shown in Figure 4.1.

Figure 4.1 Removal of the extra slurry from the templates by compressed air in order to keep an open porosity of the resulting BG scaffolds.

The PU foams and SL sponges needed to be immersed in the BG slurry three times, while for SA only two impregnations were sufficient to obtain the scaffolds. After drying, the samples, called green bodies, were submitted to heat treatment to remove the sacrificial templates and to densify the structure. The sacrificial sponge burnout and the sintering conditions were the same for all the sacrificial templates, namely, 400°C/1h and 1050°C/1h respectively. The heating and the cooling down rates used were 2 and 5 °C min-1 respectively. The process is summarized in Figure 4.2.

Figure 4.2 Scheme for the preparation of Bioglass®-based scaffolds via replica method and heat treatment programme designed for burning out templates (400°C – 1h) and sintering BG (1050°C – 1h).

4.2.2. Bioactivity Study and Characterization

The resulting scaffolds were immersed in SBF for up to 28 days in order to evaluate their bioactivity and mechanical stability in a physiological like environment. SBF was prepared as reported by Kokubo and Takadama130. For SBF preparation, 8.035 g L-1 NaCl, 0.355 g L-1 -1 -1 -1 NaHCO3, 0.225 g L KCl, 0.231 g L K2HPO4 (3H2O), 0.311 g L MgCl2 (6H2O), 0.292 g -1 -1 L CaCl2 and 0.072 g L Na2SO4 were dissolved in deionised water and buffered at pH 7.4 -1 at 36.5 °C with 6.118 g L tris(hydroxymethyl) aminomethane ((CH2OH)3CNH2) and 1M HCl. Cylindrical BG foams were immersed in SBF with a 1.5 g L-1 ratio37,130. The solution was kept in a polystyrene container at 37 °C in an incubator for up to 28 days. The solution was renewed every 48 h in order to better mimic in vitro the expected in vivo conditions15. At each time point, foams were collected, rinsed three times with deionised water, dried at room temperature for three days. The microstructure changes of the resulting scaffolds were investigated by mean of a scanning electron microscopy SEM-EDS (Auriga 0750 from ZEISS) and a stereomicroscope (Leica M50 with camera Leica IC80 HD). The structure of the foams was observed using µ-CT scanner (Skyscan 1147 Micro-CT, Bruker). The pore interconnectivity was evaluated using CTan image analysis software. The foams were also scanned using a high resolution µ-CT scanner (Skyscan 1172 Micro-CT, Bruker). The density of the foams, ρfoams, was measured using the mass and dimensions of the sintered cylinders. The porosity, p, was calculated with the formula:

!!"#$ �% = 1 − ∗ 100 Eq. 4.1 !!"#$%

-3 where ρsolid = 2.7 g cm is the theoretical density of 45S5 Bioglass® not considering change of density due to the crystallisation of the material during the thermal treatment or after the

HCA precipitation131. The nucleation of HCA was assessed by FTIR (Nicolet 6700, Thermo scientific).

The unchanged SBF was kept for the analysis of the concentration of Ca, P and Si ions through inductively coupled plasma optical emission spectroscopy ICP-OES (Optima 8300 Perkin Elmer) and for pH analysis. The percentage of the glass dissolution was calculated based on ICP results according to the following formula:

!"# !"# ��% = 100 ∗ ( ) Eq. 4.2 !"# !"

Where [ion]ICP is the concentration measured in solution by ICP analysis at a given time- point and [ion]BG is the theoretical concentration in the BG foams.

The compression strength of BG foams as-sintered and soaked for 7, 14, 21 and 28 days in SBF were measured using a Zwick Z050 mechanical tester at a crosshead speed of 1 mm min-1, with load cell of 50 N and 1 kN. The load was applied until densification of the porous samples started to occur. The compressive strength was defined as the maximum stress of the linear elastic part of the stress-strain curve.

4.2.3. In Vitro Cell Culture Studies

i. Indirect Cell Culture Test

Indirect cell culture tests were performed following the general guidance set by the International Standardization Organization (ISO 10993-5:2009). The extracting medium was prepared by placing the not-preconditioned scaffolds, sterilized at 120°C for 1h, in cell culture medium for 1 day at 37°C under 5% CO2 in a culture incubator. The total amount of cell culture medium per single scaffolds was a function of the scaffold weight (because the specific surface area of scaffolds was not available). The mass/extraction volume recommended by the ISO standard is 0.1 g mL-1. After 1 day the cell culture medium in contact with the scaffolds was collected and used for the test. Saos-2 human osteoblast-like cells were seeded in fresh and conditioned media. The conditioned medium was also diluted by 50, 10 and 1% and used for cytotoxicity evaluation. The cell culture was conducted for 24 h has recommended by ISO standard. The cell culture medium was McCoy cell culture medium (Sigma M4892, NaHCO3 free) containing 1 % vol. sodium pyruvate (Sigma S8736), 1 % vol. penicillin-streptomycine solution (Sigma P0781) and 15 % vol. fetal bovine serum (Sigma F7524). The viability of the cells was tested with Alamar Blue assay, measuring the optical fluorescence at an excitation wavelength of 570 nm and emission

49 wavelength of 590 nm. The viability of the cells cultured with fresh and aged medium was used as control.

ii. Direct Cell Culture Test

In order to evaluate the cell behavior once in direct contact with the BG scaffolds, a direct cell culture test was carried out. Every scaffold was pre-conditioned in 2 mL of cell culture medium with 30 mM HEPES in an incubator (37 °C, 5% CO2) until the pH was lower than 8. Previous studies have shown that cells better proliferate on BG scaffolds after a preconditioning in order to avoid the rapid increase of pH over non-physiological values132. In the present study the medium was changed every day and the pH was continuously monitored. After 1 week the samples were washed with PBS to remove the remaining medium from the inner core of the scaffolds, immersed for 1 h in fresh cell culture medium and dried. Saos-2 cells were seeded at a density of 5 x 105 cell cm-2 in 2 mL of fresh culture medium and cultured under an atmosphere of 5 % CO2 at 37 °C for up to 14 days. The cell culture medium used for the direct cell culture was the same used for the indirect cell test but without HEPES in order not to affect the cell viability and proliferation. The viability of cells was evaluated by mean of the Alamar Blue assay. The culture medium was renewed every two days.

The starting SA sponges are characterized by an inhalant and an exhalant surface, referred to the directionality of water flow in the sponges in their natural habitat127. For the resulting BG-SA scaffolds it was possible to distinguish the inhalant and exhalant surfaces. For this reason cells were seeded on both surfaces in order to verify if the viability was affected by the different surface structure of the scaffolds.

4.2.4. Oxygen Diffusion Evaluation

The effective oxygen diffusivity in BG based scaffolds was determined using Lattice Monte Carlo method133,134. This study was carried out by Dr. T. Fiedler, University of Newcastle (Australia). This numerical method simulates random walks of probing particles within lattice models that represent the target geometry, in this case the interconnected porosity of BG scaffolds. For geometric accuracy, lattice models were derived directly from high resolution µ-CT scans. Scans were obtained from scaffolds in as-fabricated conditions and after immersion in SBF for 28 days. Different samples had to be used for the scan of as- sintered and resorbed scaffolds since the electron beam of the µ-CT scan could have affected

50 the structure of the BG scaffolds. In order to remove calculation errors due to uneven samples, surfaces cubical sub-volumes were extracted from the center of each µ-CT data set.

The used model approach, following a previous study on diffusivity in scaffolds135, assumes the worst-case scenario of no vascularization and no diffusion through the scaffold material. Thus, the interconnected pores (porosity �) were assumed to be completely occupied by regenerated tissue with a bulk diffusivity �!. At the same time, vascularization was assumed to be incomplete and not yet contributing to the oxygen transport. The diffusivity of the scaffold material itself was approximated by zero. Thus, scaffold material act as a local diffusion barrier and reduces the effective diffusivity �!"" of the new formed biological tissue. All the mathematical equations related to the models are reported in the Appendix of this thesis work.

4.3. Results

4.3.1. Templates and Scaffolds Architecture

Images of the three templates are shown in Figure 4.3, namely, PU packaging foam and two natural marine sponges (SA and SL). The mean PU foam pore size (Figure 4.3b) was 770 ± 60 µm and the thickness of the pore walls was 110 ± 15 µm. The SA and SL sponges were characterized by a vase or fan shaped (Figure 4.3a) growth and a surface composed by fine fibers. It was possible to recognize two different structures as reported by Pronzato and Manconi127: an inhalant (Figure 4.3c,d) and an exhalant (Figure 4.3e,f) surface, referred to the directionality of water flow in the sponges in their natural habitat.

Figure 4.3 Digital camera images of SA (a1) and SL (a2) sponges and PU foam (b) templates; inhalant surfaces of SA (c) and SL (d) sponges; exhalant SA (e) and SL (f) sponges; fibrous network of SA (g) and SL (h) sponges (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis)60.

The SL was composed of almost only primary fibers with a diameter of 35 ± 3 µm, while the SA was formed by a network of primary fibers (17 ± 3 µm) and an irregular network of secondary thicker fibers (63 ± 5 µm). The other main differences between these two marine sponges were (i) the wall pore thickness (540 ± 80 µm in SL and 420 ± 40 µm in SA) and (ii) the larger and more irregular pore structure of the SL sponges. The fibers formed a mesh of macro-pores in the inhalant surface (1.0 ± 0.1 mm in the SL and 590 ± 50 µm in the SA), and an almost regular cluster of five to seven pores on the exhalant surface (1.7 ± 0.2 mm in the SL and 920 – 90 µm in the SA). These results are summarized in Table 4.1.

Table 4.1 Summary of the natural marine sponge’s architecture properties, focusing on pore dimension and structural thickness60.

Mediterranean See Indian-Pacific Ocean Spongia Lamella Spongia Agaricina (SL) (SA) Pore Dimension inhalant surface 1.0 ± 0.1 mm 590 ± 50 µm Pore Dimension exhalant surface 1.7 ± 0.2 mm 920 ± 90 µm Pore wall thickness 540 ± 80µm 400 ± 40 µm

The resulting BG based scaffolds are shown Figure 4.4, compared to the starting sacrificial templates. The obtained PU scaffolds (BG-PU) were characterized by a porosity of 93.00 ± 0.25 %, while the SL and SA replica foams (BG-SL and BG-SA respectively) had a porosity of 76 ± 2 % and 68.0 ± 0.2 %, as determined by measurements of their mass and dimensions and applying Eq. 4.1.

Figure 4.4 Replicated scaffolds derived from Spongia Agaricina (BG-SA), Spongia Lamella (BG-SL) and polyurethane packaging foam (BG-PU): light microscopy and scanning electron microscopy (SEM) images. J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”129.

However, it is possible to observe from the SEM micrographs reported in Figure 4.5, that almost all the space between the fibrous structures of the marine sponges was filled with BG and the macro-pore patterns were perfectly replicated. In fact, it was still possible to distinguish the inhalant and exhalant surfaces of both types of sponge. Some small sized

53 pores connecting the larger pores were also witnessed, forming an overall, highly interconnected porous structure.

Figure 4.5 a)b) BG-PU scaffolds at different magnification, c) inhalant and d) exhalant surface of BG-SA, e) and f) inhalant and exhalant surface of BG-SL (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale- Brovarone, Boccaccini. Published by Taylor & Francis)60.

It is also possible to observe in SEM images that no leavings of the sacrificial templates, PU packaging foam and natural marine sponges were left, confirming the effectiveness of the burning out thermal treatment used to make the scaffolds. Moreover, the hollow structure of the BG-PU scaffolds is clearly revealed in Figure 4.6a, with a resulting strut thickness of only ~8 µm (Figure 4.6b). Also for BG-SA and BG-SL scaffolds it is possible to observe the presence hollow structure of the struts (Figure 4.6c,d), which is characterized by round shape, due to the starting shapes of the fibers forming he natural marine sponges.

Figure 4.6 SEM micrographs of cross sections of (a) BG-PU and (e-f) BG-SA scaffolds to observe the strut microstructure of these BG scaffolds; schematic cross section of (b) BG-PU and (c) BG-SA struts.

SEM micrographs of the surface of the BG scaffolds are shown in Figure 4.7a. It is possible to identify three different surface typologies, formed due to the thermal treatment at high temperature: the surface typologies are reported in higher magnification in Figure 4.7b-d

Figure 4.7 (a) SEM micrographs of the surface of the BG scaffolds after the sintering process, which induces a partial crystallization of the starting BG. The different phases formed during the thermal treatment are: (b) crystalline phases, combeite and rhenanite (see XRD results below), (c-d) amorphous glassy particles fused together.

From the micrographs it is only possible to hypnotize to what the three different surfaces refer to for this reason EDS analysis have been performed to try to identify them (Figure 4.8). In Figure 4.7a, cubic structures are shown, which can be referred to one of the crystalline phases formed during the thermal treatment. Combining the SEM results with the EDS analysis on the same sample it is possible to hypnotize that this phase is Combeite, because this area is reach in Si (Figure 4.8d) and no P (figure 4.8c) is detected. In fact, of the two crystalline phases formed during the sintering of 45S5 BG, Combeite does not contain any P but only Si, Na and Ca. In Figure 4.7d is reported the SEM high magnification image of glass particles which fused together due to the diffusion during the thermal treatment process. It is possible to consider this phase as the remaining glassy phase of the starting glass which is still amorphous. This particular structure is mainly composed by P (Figure 4.8c), Na and Ca but no Si (Figure 4.8d). In some area, over the fused particles, it is possible to recognize the third surface type (Figure 4.7c). From EDX analysis it is not possible to determine its composition.

Figure 4.8 EDX analysis of the BG scaffold surface in order to identify the different phase compositions.

4.3.2. Micro-CT Analysis Micro-CT analysis was employed to assess the interconnectivity of the pore network. It was also possible to calculate the strut thickness and the pore dimensions of the obtained 45S5 Bioglass®-based scaffolds. It could be observed from the pore size distribution analysis (Figure 4.9) that when using PU as the sacrificial template, the resulting Bioglass®-based foams were characterized by a wide range of pore sizes from 100 to 900 µm, an average pore size of 670 ± 70 µm and a high level of interconnectivity (99.95%). The natural marine sponge replicas were characterized by a lower total porosity but at the same time, by a very high pore interconnectivity (> 99.5%). They contained pores in the range 0-600 µm and 0- 900 µm and an average pore size of 215 ± 20 µm and 265 ± 120 µm for the BG-SA and BG- SL, respectively. These foams also possessed significant microporosity, which was almost completely absent in the PU foam replicas, as summarized in Table 4.2.

Table 4.2 Architectural properties of BG based scaffolds prepared with PU and natural marine sponges as sacrificial templates.

BG-PU BG-SL BG-SA Pore Interconnectivity % 99.95 99.66 99.96 Average Pore Dimension [µm] 670 ± 70 265 ± 20 215 ± 2.7 % pores ≤ 200 µm 2.7 38 56 % pores within 150-500 µm 25 93 40 % pores ≥ 500 µm 80 5 16

A 3D reconstruction of the micro-CT analysis in figure 4.8 provided images of their internal architecture, showing the highly interconnected pore network.

Figure 4.9 2D section and 3D reconstruction of a),d) BG-PU, b),e) of BG-SL and f,g BG-SA obtained by µ-CT analysis.n (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis)60.

n The µ-CT analysis has been done at Politecnico di Torino in collaboration with Dr. Giorgia Novajra.

4.3.3. XRD Analysis The XRD analysis assessed the crystalline phase formation induced by the thermal treatment process. It was possible to observe that the starting BG was completely amorphous, but after the sintering process, two crystalline phases formed. The main crystalline phase was combeite (Na2Ca2Si3O9) and the secondary phase, rhenanite (CaNaPO4), as reported in Figure 4.10. The same amorphous crystalline transformation was observed for all samples. The formation of combeite and rhenanite in sintered Bioglass® structures has been discussed in literature150. From Rietveld analysis it was possible to estimate that around 68% of the starting BG crystallized and that the 62% of the crystalline phase is combeite and the remaining 6% is rhenanite.

Figure 4.10 XRD spectra of BG powder, BG-SA and BG-SL scaffolds showing the crystalline phases formed after the thermal treatment. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis)60.

4.3.4. Mechanical Test

The maximum compressive strength values were recorded as being 1.8 ± 0.3 MPa for the BG-SL, 4.0 ± 0.4 MPa for BG-SA and < 0.05 MPa for the BG scaffolds prepared using PU foam as template. Typical stress-deformation curves are shown in Figure 4.11 for BG-SA and BG-SL. For BG-PU, the measurements were below the detection limit of the equipment.

Figure 4.11 Stress-displacement curves of BG foams manufactured with different sacrificial templates. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis)60.

4.3.5. Bioactivity and Mechanical Stability

i. Weight and Porosity Variation The as-sintered scaffolds BG-SA, BG-SL and BG-PU scaffolds were characterized, respectively, by porosities of 68 ± 0.2%, 76 ± 2.5% and 93 ± 0.25%. The total porosities were determined by measurement of the scaffold mass and dimensions and applying Eq. 4.1. After the immersion in SBF a decrease of the weight (Figure 4.12a) and an increase of the total porosity (Figure 4.12b) were observed for BG-SA and BG-SL samples. Both types of scaffolds prepared with natural marine sponges as the sacrificial template lost between 40 and 50% of their starting weight upon 28 days in SBF. The main weight change was observed in the first week of the test. The porosity of BG-SA samples, at the end of the test, increased up to 80 % and the porosity of BG-SL samples increased up to 88 %. For the BG- PU scaffolds it was not possible to analyse the weight and porosity variation because the samples were almost completely degraded due to their high dissolution in SBF and fragility.

Figure 4.12 (a) Weight and (b) porosity variation of BG-SA and BG-SL scaffolds after different immersion time in SBF for up to 28 days. The standard deviation values of the data for BG-SL in (b) were too low to be visible. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale- Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

ii. Mechanical Stability By testing ten samples in as-fabricated conditions the resulting maximum compressive strengths values were determined as follow: 1.8 ± 0.3 MPa for the BG-SL, 4.0 ± 0.4 MPa for BG-SA and < 0.05 MPa for BG-PU (the measurements were below the detection limit of the equipment) samples. BG-PU scaffolds were too fragile to be handled and it was not possible to perform the compression strength test on these samples. This was probably a consequence of the reduction of weight and increase of the volumetric porosity, after soaking in SBF. On the other hand, even after 28 days in SBF, it was still possible to handle BG-SA and BG-SL scaffolds without damaging them. For BG-SA and BG-SL samples, the highest loss of mechanical strength took place during the first 7 days of immersion when a reduction of 50% from the starting maximum compressive strength values occurred. At the end of the test, the resulting compressive strength values were 1.2 ± 0.2 MPa for BG-SA and 0.23 ± 0.01 MPa for BG-SL samples (Figure 4.13). In both cases the values were higher than the maximum compressive strength of as-sintered BG-PU samples.

Figure 4.13 Compressive strength variation of BG-SA and BG-SL scaffolds after immersion in SBF for up to 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

Figure 4.14 Stress-displacement curves of BG-SA scaffolds after immersion in SBF for up to 28 days.

iii. Ion Release and pH Variation The pH variation and the changes of ion concentration in SBF during the first hours of immersion are reported in Figure 4.15. Positive values of ion release imply an increase of the ion concentration in SBF, while negative values refer to a depletion of the ion concentration

62 in SBF due to the immersion of the samples. Scaffold dissolution took place immediately after the first hours of immersion, as shown by the rapid increase in pH and the continuous ion release. After 3 days of immersion in SBF around 40 mg L-1 of Si ions was released from both the BG-SA and BG-SL samples. The corresponding Si release from BG-PU was higher, at almost 61 mg L-1. The Ca concentration increased significantly after 6 h, reaching around 10 and 25 mg L-1, respectively for natural sponges (both BG-SA and BG-SL) and BG-PU samples. The P concentration in SBF decreased over time, indicating that a P-containing material was deposited on the surface of the scaffolds. No pH equilibrium was reached even after 2 days of testing, suggesting a continuous dissolution of the scaffolds (Figure 4.15d).

Figure 4.15 Concentration of (a) Si, (b) Ca and (c) P ions released from BG scaffolds during the first hours of immersion in SBF for up to 3 days; (d) pH variation in the first 48 h of immersion in SBF.o p (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

iv. Surface Modification Analysis: SEM and FT-IR SEM micrographs of foams, before and after immersion in SBF for 1, 3, 7, 14 and 28 days are reported in Figure 4.16. Already after 1 day of immersion a deposit was seen to form on the entire surface of the scaffolds. After 3 days of immersion (Figure 4.16c,i,o), the deposit increased in thickness and the cauliflower-like structure (typical of HCA)6,136,137 was identified. At 28 days of immersion in SBF, the deposit continued to grow and the needle- o The ICP analysis was carried out at Politecnico di Milano by Dr. Simone Gelosa. p The pH analysis was carried out at Politecnico di Milano under the supervision of Dr. Lina Altomare and Dr. Luigi De Nardo.

63 like structure of the HCA phase was clearly revealed. The same results were obtained for scaffolds prepared using PU foams and both natural marine sponges. No perceivable differences were identified in terms of HCA deposit rate and evolution of the deposit. It was observed that in all cases, the struts of the scaffolds were completely covered by HCA, as documented in Figure 4.16f,l,r.

Figure 4.16 SEM micrographs of BG-SA (a-f), BG-SL (g-l) and BG-PU (n-s) scaffolds at different immersion times in SBF (for up to 28 days) showing evolution of HCA) deposit. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

The HCA phase evolved not only on the external surface of the BG struts, but also inside the hollows structure of the struts, as shown in Figure 4.17.

Figure 4.17 SEM micrographs of cross sections of single struts of the BG-PU (a,c) and BG-SA (b,d) scaffolds after immersion in SBF for 14 days showing the evolution of the HCA phase both on the external surface and inside the hollow structure of the struts of the BG scaffolds.

Figure 4.18 summarizes FTIR spectra of the samples before and after immersion in SBF for up to 28 days. In the case of the reference 45S5 BG, the FTIR spectra at room temperature presented the characteristic peaks of the non-bridging oxygen stretching mode of Si–O located at 950 cm-1 26,138.After 14 days, characteristic peaks ascribed to the HCA layer formation appeared as a doublet at around 600 cm-1, corresponding to the bending mode of P–O (crystalline phosphate)26,138. Moreover, the observation of P–O stretching at around 1000 cm-1, where the band became narrow, suggested the presence of HCA26. The spectra also showed the narrowing of the band at around 800 cm-1 corresponding to the bending mode of C–O and finally the manifestation of the stretching mode of C–O at around 1400 cm-1 26. The SEM micrograph in Figure 4.18d shows a well-developed HCA layer on the BG- SA scaffold surface after 28 days in SBF.

Figure 4.18 FTIR analysis on BG-SA (a), BG-SL (b) and BG-PU (c) scaffolds before and after immersion in SBF for different number of days and SEM micrograph (d) showing the needle like structure of HCA formed on the BG-SA surface after immersion in SBF for 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

v. Micro-CT Analysis Analysis of the 3D structure of the present scaffolds by µ-CT was carried out in order to gain insight on how immersion in a physiological-like environment could influence the scaffold porosity. In Figure 4.19, typical 2D reconstruction images of the BG scaffolds immersed in SBF for up to 28 days are reported. For the samples obtained from natural marine sponges, it is observed that after 7 days the material became less dense at its edge (reduction in brightness) due to the dissolution process starting from the external surface of the foams. After 14 days, this less dense area expanded deeper into the specimens. The scans at 28 days revealed the formation of an external layer of a more dense material, likely related to the HCA layer observed by SEM (Figure 4.19). The white arrows in figure 18a indicate the thicker layer of HCA, grown after 28 days in SBF. For the BG-PU scaffolds it was possible to observe that almost all the material dissolved after only 7 days of immersion, but there was a progressive increase of the pore wall thickness due to the deposition of HCA (observed also through SEM, see Figures 4.16 and 4.17). All these findings assume that material dissolution in SBF caused density variations. In particular, a less dense area identified a portion of dissolved material and appeared darker than the non-dissolved zones73. The HCA phase grew over time on the surface of the foam which was clearly visible as a whiter layer (because of its higher X-ray absorbance compared to the degrading BG). A

66 broader distribution of pore sizes was observed for BG-SA and BG-SL scaffolds immersed in SBF for up to 28 days, as shown in Figure 4.19b,c. This was observed for up to 3 weeks of immersion in SBF. However, after this time-point the pore size distribution did not change further, most probably due to the simultaneous occurrence of BG dissolution and HCA precipitation. In addition, it was confirmed that the pore interconnectivity of the starting scaffolds did not change over time, both BG-SA and BG-SL scaffolds were characterized by a high interconnected pore structure (>99%).

Figure 4.19 µ-CT reconstructions of BG-SA, BG-SL and BG-PU scaffolds before and after immersion in SBF for up to 28 days (a), pore size variation of BG-SA (b), and BG-SL (c) scaffolds after immersion in SBF for up to 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

Figure 4.20 shows the spatial porosity variation within the produced BG scaffolds in the perpendicular x-, y- and z-directions. The porosity of BG-SA scaffolds shows a fluctuation around an average value of 68.2% (as sintered) and 72.3% (after 28 days in SBF). In contrast, BG-SL samples exhibits a symmetric decrease of porosity in the pore (x-) direction from 81.6 to 51.8%. This deviation is caused by the different geometries of inhalant and exhalant surfaces of the SL sponges. For the orthogonal y- and z-directions the porosity values oscillate around the average values of 61.2% of the as-sintered BG-SL scaffolds and 69.8% of the samples after immersion in SBF for up to one month. The analysis of BG-PU samples reveals an almost constant porosity with an average value of 93.5% and 94.7% respectively before and after the bioactivity test in SBF.

Figure 4.20 Porosity variation in BG-SA (a), BG-SL (b) and BG-PU (c) scaffolds in as-fabricated conditions and after the immersion in SBF for 28 days.q (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”)129.

vi. Cell Culture Study

Indirect cell tests were conducted in order to evaluate potential cytotoxicity of the dissolution products of the BG-based scaffolds. Saos-2 human osteoblast-like cells were used. The toxicity test showed that all the extraction volumes obtained from immersing the non-pre-treated BG-based scaffolds for 1 day had a deleterious effect on cell viability if not diluted (i.e. 100%) (Figure 4.21). However, the diluted conditioned medium (50%) which was in contact with the BG-SA samples showed high cell viability comparable with that of the control (no significant viability reduction, p < 0.05). Conditioned medium diluted at 10 and 1% showed the same cell viability compared to the control for all the samples. Saos-2 cells were also directly seeded onto BG-PU, BG-SA and BG-SL scaffolds in 24-well culture plates. Considering the cells adhered to the specimens, SEM observations revealed normal

q The evaluation of the porosity variation has been done by Dr. Thomas Fiedler at the University of Newcastle, Callaghan in Australia.

68 cell morphology after 14 days of culture with clear evidence of active cell attachment over the scaffold surfaces (Figure 4.21 b-d). The cell viability on the BG-PU and BG-SL scaffolds was significantly lower when compared to the control, while the viability of cells seeded on the BG-SA samples was comparable to the control. For BG-SA scaffolds, no significant differences were found in cell viability values obtained on the inhalant or the exhalant surfaces of the scaffolds.

Figure 4.21 (a) Cell viability results from indirect cell culture test with different dilutions of the extraction volumes, SEM micrographs of (b) BG-PU, (c) BG-SA and (d) BG-SL scaffolds after 14 days of direct cell culture study, (e) cell viability results from direct cell culture for up to 14 days (* statistically different).r (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale- Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”)128.

vii. Oxygen Diffusivity In Figure 4.22 normalized oxygen diffusivity data are plotted versus the scaffold porosity. As reported in the materials and method section, the simulation method assumes that the newly regenerated tissue completely occupied the interconnected porosity. For comparison, r The cell culture study has been done at the Department of Chemistry, Materials and Chemical Engineering “G. Natta” at Politecnico di Milano. SEM images were obtained by Dario Picenoni and Monica Moscatelli, cell viability study was done by Dr. Lina Altomare.

69 dimensionless diffusivities of the present study are shown along with data from a previous study135.

Figure 4.22 Normalised oxygen diffusivity data plotted versus scaffold porosity for BG-SA, BG-SL and BG-PU scaffolds.s (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”)129.

Partial scaffold resorption due to 28 days immersion in SBF increases both the BG-PU scaffold porosity and diffusivity. In addition, a distinct anisotropy is observed. Due to the strong porosity variation and in order to investigate the variation of diffusivity in the scaffolds, the geometry of original BG-SL and BG-SA scaffolds was subdivided into non- intersecting prisms and diffusivities determined for each sub-volume. Scaffolds of the same type with similar porosity exhibit comparable diffusivities indicating that diffusivity is governed by porosity. The data also shows that BG-SA scaffolds exhibit a higher diffusivity than BG-SL at the same porosity. The partially resorbed BG-SL exhibits a significantly higher porosity and diffusivity that is similar to BG-SA samples.

Table 4.3 correlated the effective diffusivity �!"" and structural parameters of tissue engineering scaffolds, i.e. the mean scaffold strut thickness t and the mean pore size s, which were determined using the plugin BoneJ of ImageJ free software. As expected, the scaffold thickness t generally decreases and the pores size s increases due to immersion in SBF. The apparently contradicting decrease of the average pore size for resorbed BG-SL scaffold is attributed to the use of a different sample for the resorption study.

s The diffusivity analysis was carried out at the University of Newcastle in Australia by Dr. Thomas Fiedler.

Table 4.3 Diffusivities and structural parameters of tissue engineering scaffolds. Table modified from Boccardi et al.129.

* D STDEV Φ t CVt s CVs Structure * [-] D [%] [µm] [-] [µm] [-] 0.39 0.052 BG-SA (original) 68.2 80 0.36 209 0.55 BG-SA (resorbed) 0.45 - 72.3 81 0.40 244 0.51 BG-SL (original) 0.25 0.107 61.2 103 0.55 335 0.74 BG-SL (resorbed) 0.43 - 69.8 86 0.38 268 0.64 0.87 0.017 BG-PU (original) 93.5 86 0.36 687 0.17 BG-PU (resorbed) 0.90 - 94.7 75 0.45 729 0.20

It has been observed previously that the diffusivity is affected by the coefficient of variation

CVs of the pore size. A low value of CVs indicates uniform pore size resulting in a high diffusivity. Furthermore, high values of CVs are consistent with the bimodal nature of the pore size distribution in these scaffolds. A close correlation between the variation of pore size and the standard deviation of diffusivity is found, i.e. a constant pore size yields a low variation of diffusivity. Figure 4.23 shows anisotropy plots of diffusion in scaffolds. The graphs indicate directional diffusivities in the perpendicular xy-, xz- and yz-planes. In these plots, isotropic behavior is indicated by circular graphs as obtained for BG-PU scaffolds (Figure 4.23c). The degree of anisotropy can be quantified using the ratio χ between minimum and maximum diffusivity where a maximum possible value of unity corresponds to perfect isotropy. In the case of BG-PU scaffolds (Figure 4.23c) the ratio is χ = 0.98 for both the original and the partially resorbed scaffold thus indicating isotropic diffusion properties. BG-SA scaffold is shown in Figure 4.23a and moderate anisotropy is found. The corresponding anisotropy ratio is χ = 0.78 (as-fabricated conditions) and χ = 0.85 (after 28 days in SBF) indicating that anisotropy decreases due to resorption. As expected, maximum values are found for the x-direction which is the predominant pore direction of BG-SA samples. BG-SL scaffold shows a similar behavior; however, anisotropy is significantly stronger. Two datasets for original scaffolds are shown that represent the inhalant and exhalant surfaces, respectively. The maximum anisotropy (χ = 0.48) is observed within the original exhalant surface. This can be explained by a large number of narrow channels formed by the tubular pores. Anisotropy is decreased in the inhalant area region (χ = 0.66) and for the partially resorbed BG-SL scaffold (χ = 0.80).

Figure 4.23 Anisotropy of normalized scaffold bioactivity showing the porosity variation inside the BG scaffolds along different directions. (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”)129.

4.4. Discussion 45S5 Bioglass®, the original formulation of bioactive glasses discovered by L. Hench in 19699,10,123, is the most used bioactive glass in clinical applications as powders or bulk material10. So far, BG based porous scaffolds have not been considered for load bearing applications due to their relatively low mechanical properties10,123. In fact, it is still a challenge to obtain a BG scaffolds presenting at the same time good mechanical properties and high porosity123. In the present work, the possibility to increase the mechanical behavior of BG based scaffolds by reducing the total porosity was considered using natural marine sponges, SA and SL, chosen as sacrificial templates for the production of the scaffolds123,124. These sponges were characterized by a high interconnected porous structure made by fine fibres127. The architectural properties were well reproduced for SA; however, high variability in pore dimension and strut thickness was observed for SL. As a result of the high porosity of these natural marine sponges, the resulting BG-based scaffolds were characterized by a very high interconnectivity of pores (> 99%). However, the total porosity was ~ 18–20% lower than that of samples obtained with PU foam (45 ppi) as sacrificial template. In addition, there was a significant amount of pores in the range 0–200 µm, which were impossible to obtain using the PU template. Pores in this size range play an important role in

72 the complete colonization of the scaffolds by cells, enhancing the flow of biological fluids also in the inner core of the porous structure and the complete integration with the surrounding tissue53,54. Moreover, pores in the range 200–500 µm were present, which are necessary for bone integration and neovascularisation53,54. As a main consequence of the reduction of the total porosity, there was an improvement of the mechanical properties up to 4 MPa for BG-SA and up to 1.8 MPa for SL scaffolds. Another key point for the better mechanical properties of the BG scaffolds obtained using natural marine sponges as sacrificial templates, is the presence of fewer and smaller holes in the struts of the scaffolds, such hollows are typical for BG-PU scaffolds139. This is due to the different thickness of the walls of pores in PU foam and in natural marine sponges. The synthetic foam was characterized by pore walls of 110 µm, compared to the 35 µm of the SL primary fibers, 63 µm and 17 µm of the primary and secondary fibers type of the SA sponges. In the case of BG-PU scaffolds, the resulting maximum compressive strength was lower than the average values found in literature (~0.4 MPa)123. This result was probably due to the compressed air used for the removal of the slurry excess after the second and third coating cycle during scaffold preparation. Following this technique, it was easier to maintain an open porosity, but the amount of slurry removed was higher possibly reducing the thickness of the struts. The XRD results confirmed that the thermal treatment caused a modification of the BG structure, inducing the formation of two crystalline phases: combeite (Na2Ca2Si3O9) as the main phase (~ 62% of the total crystalline phase), as widely reported in literature15, and rhenanite (CaNaPO4) as the secondary crystalline phase (~ 6% of the total formed crystalline phase)140–142. It was demonstrated that rhenanite can act as a heterogeneous nucleus for carbonated hydroxyapatite crystallization in contact with simulated body fluid, and thus, it should improve the bioactivity of the material140,141,143,144. A description of the structural transformations of BG occurring during thermal treatment was reported by Lefebvre et al.141, summarized in Figure 4.24. The first step is the glass transition around Tg1 = 550°C, followed by a glass in glass phase separation at Ts = 570°C, with the appearance of one silica-rich phase and one phosphorous-rich phase. This means that the glass is no more homogeneous but rather consists of two immiscible phases. This kind of separation is expected when two high valence ions, such as Si4+ and P5+, are present simultaneously in a glass. Each ion type tends to concentrate in a separate phase. The Si-phase leads to the major crystalline phase (combeite) and the P-phase leads to the secondary crystalline phase. The major crystalline phase appears at TC1 = 610°C and the secondary crystalline phase at TC2 =

800°C. At Tg2 = 850°C the second glass transition occurs the melting of the BG takes place at Tm = 1200°C.

Figure 4.24 Summary of 45S5 BG structural transformations identified by Lefebvre et al.141.

The achievement of improved mechanical behavior, combined with the high pore interconnectivity and wide pore size distribution, confirmed the feasibility of using natural marine sponges from the Elephant Ear family as porous precursors for production of bone tissue scaffolds. As a consequence of the immersion of the BG scaffolds in SBF for up to 28 days, a decrease in weight and an increase in total porosity of the BG-SA and BG-SL scaffolds were observed. At the end of the test, BG-SA and BG-SL scaffolds respectively lost 40 and 50% of their weight due to the degradation process of the bioactive glass. Interesting, almost the whole weight loss occurred during the first week of immersion in SBF, while in the subsequent 3 weeks the weight values remained almost constant. This behavior was more pronounced on the BG-SL samples and was most probably a consequence of the established equilibrium, being achieved after 7 days in SBF, due to the simultaneous degradation of the BG and the precipitation of HCA on the surface of the scaffolds. In addition, the porosity changed continuously during the entire period of the test and a plateau was reached only between three and 4 weeks of immersion in SBF. One possible explanation for this apparent mismatching between weight and porosity variations can be the different densities of BG and HCA. The material deposited on the surface of the scaffolds is characterized by a higher density compared to that of the BG struts. It should be noted that throughout the SBF test, no variation in the geometry of the scaffolds was observed over time. During the first 3 weeks of immersion in SBF, both degradation and precipitation occurred but degradation was the dominant phenomenon (Figure 4.25). After 3 weeks, the two mechanisms reached equilibrium and the results indicate that at higher immersion times in SBF, precipitation could be the dominating phenomenon until the HCA completely covered the structure of the scaffolds (Figure 4.25). This behavior is in agreement with previous modeling work reported in the literature145 and similar results were recently obtained for Bioglass®-based scaffolds produced via powder metallurgy-derived technology128. As a consequence of the porosity increase, the compression strength of the BG-SA and BG-SL scaffolds decreased. After 28

74 days in SBF, the resulting compressive strength values were 1.2 ± 0.1 MPa for the BG-SA and 0.23 ± 0.01 MPa for the BG-SL scaffolds, these values were still higher than those of as- sintered BG-PU scaffolds (< 0.05 MPa). In the present case, most of the compressive strength loss took place during the first week of immersion in SBF and afterwards the values were almost constant. This result is in agreement with the measured weight and porosity variations of the scaffolds. One possible explanation is that due to the higher density of the HCA deposit compared to the degraded BG struts, even if the total porosity was increased, strengthening of the scaffolds was induced. Moreover, since the HCA formation commenced at the areas of the scaffolds with higher concentrations of cracks, these sites probably acted as favorable nucleation points for HCA, when exposed to SBF. It is possible that the HCA crystals filled these cracks strengthening the overall scaffold structure.

Figure 4.25 Schematic representation of the two phenomena taking place after immersion of BG scaffolds in SBF: the BG starts to dissolve on the external surface and in the inner core of the struts, the HCA starts to form as a homogenous layer on the external surface of the scaffolds and inside the BG struts.

As shown by the ICP results (Figure 4.15), the dissolution of the glass took place immediately after immersion in SBF inducing a rapid increase in the pH, mainly due to the fast exchange of Na+ and Ca2+ with H+ and OH- ions from the solution. This effect caused the hydrolysis of the silica groups and consequently, the formation of silanol groups which are known to be the starting point for the nucleation of HCA25. Due to this very fast reactivity of BG, it was possible to identify the presence of a deposit on the surface of the scaffolds already after 1 day in SBF. This deposit evolved in thickness and after 14 days it was possible to identify the crystallization of the deposit (HCA), as confirmed by FTIR spectra. This deposit did not grow only on the external surface of the scaffolds but also on the surfaces of the inner core of the BG-foams, as confirmed by µ-CT analysis. These observations highlight the high level of interconnectivity of the porous network of all produced samples. Another interesting aspect is that the pore size distribution became

75 broader for both BG SA and BG-SL scaffolds. Both smaller and larger pores increased in number, especially at 14 days of immersion in SBF. Moreover, even though HCA precipitation was significant, the smaller porosity was not occluded suggesting homogeneous degradation of the scaffolds in the entire volume. The high surface reactivity of non-pre- treated BG based scaffolds also resulted in cell culture test showing a low viability of cells cultured for 24 h in the non-diluted extraction medium. In the case of the BG-SA scaffolds, already after a dilution of 50%, no significant viability reduction was observed (p < 0.05) compared to the control. Moreover, the direct cell culture test on BG-SA samples showed that cells grew similarly on both surfaces of the scaffold (inhalant and exhalant) and the cell viability was comparable to that on the control. In the case of BG-PU and BG-SL scaffolds, the cell viability was significantly lower, most probably due to the higher pH values around the scaffolds. This behavior indicated that in these scaffolds the preconditioning treatment was not sufficient to avoid the formation of an undesirable non-physiological pH.

Tissue engineering scaffolds act as diffusion barriers that decrease the intrinsic diffusivity of growing tissue. Results on the present BG-based scaffolds indicate a decrease to 39 % (BG- SA) and 25 % (BG-SL) of the corresponding bulk tissue diffusivity. Furthermore, the effect of scaffold resorption due to immersion in SBF for 28 days was addressed. The resulting increase in porosity caused increased effective diffusivities equal to 45 % (BG-SA) and 43 % (BG-SL) of the intrinsic value. In addition, the anisotropy of the diffusivity was investigated. The degree of anisotropy was quantified by calculating the ratio χ of minimum and maximum directional diffusivities. Strong anisotropy was found for BG-SL (χ = 0.48) and moderate anisotropy for BG-SA (χ = 0.78). This directional variation of diffusivity can be explained by the tubular geometry of pores in the marine sponges. In contrast, scaffolds replicated from PU foams showed isotropic diffusion behavior (χ = 0.98). In comparison to scaffolds manufactured using sacrificial PU foam templates (BG-PU) a distinct decrease of the diffusivity was found for BG-SA samples and BG-SL scaffolds. In the case of BG-SA samples the average diffusivity is only half the effective diffusivity of BG-PU and for BG- SL scaffolds only a third of this value is obtained. However, the mechanical strength of BG- SA and BG-SL scaffolds exceeds that of BG-PU scaffolds by a factor of at least 40 and the optimum scaffold geometry depends on the combination of strength requirements, diffusivity and biological compatibility.

4.5. Conclusions The use of natural marine sponges as sacrificial templates for porous 45S5 Bioglass®-based scaffold preparation has been demonstrated as promising alternative to PU foams. The obtained scaffolds are characterised not only by the high interconnectivity of their porous structure but also by sound mechanical properties. In this way, it is possible to combine the main properties (structural, mechanical and biological) that a 3D scaffold should have for bone regeneration. The bioactivity and mechanical stability of BG-based scaffolds derived from natural marine sponges was demonstrated. The obtained samples were characterized not only by improved mechanical properties (compressive strength up to 4 MPa) compared to the foams prepared using PU foam as a template, but also after immersion in SBF since the mechanical properties were stable for up to 7 days of testing. Moreover, the reduction of total porosity of the BG-SA and BG-SL scaffolds did not affect the bioactivity and already after 1 day of test in SBF, a deposit of HCA (the marker of bioactivity) was detected. Preliminary cell culture tests demonstrated that no toxic residues came from the natural marine sponges used as templates and cells proliferated extensively on the scaffolds. In conclusion, the best results in terms of mechanical properties, bioactivity and cell biocompatibility were obtained using “Spongia Agaricina” as template for BG scaffolds. These foams represent thus a new attractive family of BG-based scaffolds for bone tissue engineering warranting further in vitro and in vivo studies.

5. MCM-41 Coating of BG-based Scaffolds

5.1. Introduction One promising field of tissue engineering involves the development of porous 3D engineered scaffolds to enhance bone regeneration and neovascularization3. The main challenge is the design of materials able to match at the same time the biological and the mechanical properties of the natural bone tissue1,54,123. However, the design of the scaffolds is not the only challenge, in fact the first problem after implantation is the exposure to inflammatory and infection risks with further complications, e.g., septicemia and potential implant failure146. To avoid these consequences a large amount of antibiotics and anti-inflammatory drugs are administered to the patient, which can increase the healing time, the stay at the hospital, and costs147. Nowadays, the most popular ways for drug intake are oral administration and injection. However, these methods may be affected by a lack of efficiency especially since the release of the drug is not targeted specifically to the area that needs to be treated43. For all these reasons, the development of local drug release systems, which enable controlled release kinetics, has increased considerably during the past few years104,114,116,148. In this context, the combination of bioactive scaffolds with local drug delivery carriers is gaining increasing research efforts in the bone tissue engineering field123. Several matrices have been tested so far, such as organic polymers, organic–inorganic hybrid materials, bioactive glasses (BG), and ceramics148. One approach gaining increasing interest involves obtaining drug carriers that are structured at the nanoscale. Since 1992, when silica- based MCM-41 was developed (Mobil Composition of Matter No. 41)84, highly ordered mesoporous materials have attracted the attention of many scientists and in 2001 they were proposed as drug delivery system63. The most interesting features of these materials are the regular pore system, high specific surface area and high pore volume42,43,85,86,92. These silica- based mesoporous materials are able to incorporate relatively high content of drugs into the mesopores. Moreover, their silanol groups can be functionalized and the pore diameter can be modulated, allowing a better control of the drug-release kinetic86,92,148. Two mechanisms have been proposed to describe mesoporous silica material formation. The first model describes the addition of silicate to micelles formed using n-decyltrimethylammonium bromide (CTAB). In this way, the silica precursor polymerizes around the already formed micelles42. The second proposed mechanism is that the addition of the silica precursor to an aqueous n-decyltrimethylammonium bromide solution induces the ordering of silica-encased surfactant micelles simultaneously. In this case, the micelle formation requires the silica precursor to be present42,43. MCM-41 has become the most popular member of the

79 mesoporous silicate materials family, and it has been considered also as drug carrier85. Nowadays, it is possible to find in literature different approaches for the synthesis of spherical MCM-4190,149–151. Grün et al.90 proposed a novel pathway for the production of spherical MCM-41 applying a modification of the Stöber reaction91 for the synthesis of spherical non-porous silica particles. The approach involves introducing a low-boiling alcohol, such as ethanol or isopropanol, as co-solvent for the silica source in order to obtain a more homogeneous solution90. Starting from the work of Grün et al.90, it is possible to obtain well-shaped spherical particles; however, the mesoporosity is not homogeneously present. On the other hand, following a standard procedure reported by Zeleňák et al.150, it is possible to obtain well-ordered mesoporous structures; however, the particles are not spherical and the size distribution is usually broad. Combining these two synthesis pathways, a new solution for the synthesis of spherical mesoporous silica particles has been proposed in this study. Thus, the aim of the work reported in this chapter is the synthesis of spherical silica mesoporous particles (MCM-41) developing a novel synthesis procedure. The obtained particles will be evaluated for their features as ordered mesoporous materials and the stability of the ordered structure in SBF will also be studied. The possibility to functionalize these particles with Ag nanoparticles is also presented. The final goal is the use of the developed particles as a functional coating of BG-scaffolds in order to combine the drug up- take/release capability of MCM-41 particles with the bioactivity properties of BG. The concept of incorporating a silica drug carrier into bioactive silicate scaffolds has been previously explored63; however, the main advantage of the approach introduced in this paper is the possibility to obtain a highly homogeneous coating of the BG scaffold struts with highly ordered mesoporous silica particles enhancing the BG bioactivity and improving the resulting mechanical properties of the combined system. Moreover, the total amount of produced particles obtained per single batch increases by ~60% combining the two standard procedures reported in literature90,150, which represents another advantage of the present approach. The results presented in this chapter have formed part of previous publications64.

5.2. Materials and Methods

5.2.1. Synthesis of MCM-41 Particles The procedure adopted to prepare MCM-41 was a combination of the standard pathway for the production of mesoporous silica particles and a modification of the Stöber reaction91 for the preparation of non-porous silica spheres proposed by Grün et al.90. In this way, the reaction took place in a more homogeneous environment, resulting in the formation of submicron sized spherical MCM-41 particles and the total amount of the cationic surfactant, which is extremely toxic, can be reduced90. A low-boiling alcohol, e.g. ethanol, was added as co-solvent for the tetra-n-alkoxysilane to make it soluble. The reactants were ammonia (catalyst of the reaction), n-hexadecyltrimethylammonium bromide (CTAB, surfactant), pure ethanol (co-solvent of silica source), and tetraethyl ortho-silicate (TEOS), all purchased from Sigma-Aldrich (Germany). Pure ethanol and ammonia (28–30 wt.%) solution were mixed with deionized water. The cationic surfactant was added to the solution under continuous stirring for 20 min. Once the solution was clear, TEOS was added (0.25 mL min−1). All synthesis steps were carried out at room temperature (RT), which is the optimal temperature condition for the reaction with cationic surfactant in basic conditions as reported by Zhao et al.42. After 2 h of stirring, the resulting dispersion was centrifuged and washed once with deionized water and twice with ethanol in order to remove completely every trace of ammonia, collected in a ceramic crucible, dried, and calcined in air. The solutions used are reported in Table 5.1. For samples MCM-41_A150, MCM-41_B, and MCM-41_C, the thermal treatment was 60°C (2°C min−1) for 12 h and 550°C (2°C min−1) for 6 h; for sample MCM-41_D90, the thermal treatment was 90°C (2°C min−1) for 12 h and 550°C (1°C min−1) for 5 h.

Table 5.1 Composition of four different synthesis solutions used for the preparation of MCM-41 particles. MCM-41_A particles were prepared following the work of Zeleňák et al.150, MCM-41_D particles were prepared following the work of Grün et al.90, MCM-41_B and MCM-41_C particles were developed combining the two previous synthesis solutions. Table modified from Boccardi et al.64.

Sample H2O EtOH NH3 CTAB TEOS [mL] [mL] [mL] [g] [mL] MCM-41_A150 29 - 18.5 0.2 1 MCM-41_B 11 18 18.5 0.2 1 MCM-41_C 4 25 18.5 0.2 1 MCM-41_D90 11 19 3.3 0.62 1.25

The optimized particles, MCM-41_B, were then compared to one of the most used particles known in literature, which were synthesized for the first time by Cai et al.149. A solution of 480 mL of deionized water and 3.5 mL of NaOH (2M) were heated up to 80°C in an oil bath. Once the temperature was stable, 1 g of CTAB was added and the solution was stirred until complete dissolution of the surfactant. 5 mL of TEOS were added (0.25 mL min-1). After 2 hours of stirring, the solution was centrifuged and the precipitate was washed once with water and twice with ethanol. The surfactant was removed overnight using a surfactant extracting solution made with 25 mL deionized water, 475 mL of ethanol and 5 g of ammonium nitrate. These particles will be named MCM-41_Ref in the text.

5.2.2. Synthesis of Ag_MCM-41 Particles Silver nanoparticles were loaded to the optimized ordered mesoporous silica particles, MCM-41_B, by template ion exchange (TIE) method. 2 mg of as synthesized particles were immersed in a 50 mL water solution containing different concentration of silver nitrate (6, 3 -1 and 0.6 mg mL of AgNO3, purchased from Sigma-Aldrich, Germany), at RT for 1h and at 80°C for 20 h under gentle continuous stirring. The particles were then centrifuged (7000 rcf,

22°C) to separate them from the unreacted AgNO3, washed once with deionized water and twice with ethanol, dried at 60°C (2°C/min) for 12 h and calcinated at 550°C (2°C/min) for 6 h.

5.2.3. Composite System Preparation The scaffolds used for the preparation of the composite system were BG-PU and BG-SA foams. The scaffolds were prepared as reported in chapter 4, section 4.2.1. Two different methods were used for the preparation of the composite systems: direct coating (DCM) and post coating (PCM) methods (Figure 5.1). The DCM coating procedure consisted in four steps, i.e. hydrolysis of TEOS, dipping of the scaffolds in the particles synthesis solution, drying of the scaffolds and calcination63. After TEOS addition, the solution was stirred for 10 minutes to promote the hydrolysis of the silica precursor. The scaffolds were then immersed in the silica synthesis batch for 10 minutes and meanwhile the solution was kept under vigorous stirring in order to enhance the coating also of the inner core of the scaffold. The BG-PU scaffolds were closed in a high porous cage in order to avoid the direct contact with the stirring bar, which would destroy them. The BG-SA scaffolds were put in the beaker containing the particles synthesis solution without any protection and they were not damaged63. The PCM method was developed in two different ways, i.e. A and B. In case of PCM_A method the particles were synthesized, washed and dispersed in ethanol using

-1 -1 different ratios of wet particles weight in ethanol (0.1 gwet mL , 0.2 gwet mL and 0.3 gwet mL-1). The scaffolds were then impregnated with this solution, dried and calcinated. In order to better control the amount of particles used for the scaffold impregnation, the method PCM_B was also developed. Once the particles were synthetized and washed, they were dried at 60°C for 12 h before being dispersed in ethanol (10 mg mL-1, 50 mg mL-1and 100 mg mL-1) and used to impregnate the scaffolds. To facilitate the particles infiltration also in the inner core of the scaffolds, the samples were placed for 10 min at RT in a vacuum oven. The calcination temperature was 550°C (2°C min-1) for 6 h in air for all processing methods.

Figure 5.1 Flowchart of the different methods developed for the production of the composite system MCM-41 coated BG scaffolds.

MCM-41_A, MCM-41_B, and MCM-41_D samples were used for the preparation of BG_MCM-41 composite scaffolds following the CCM method, which was the one already developed in literature. MCM-41_C solution was not used because it did not show any

83 ordered mesoporosity. The PCM methods were then applied only to the optimized ordered mesoporous silica particles, MCM-41_B. In case of the functionalized ordered mesoporous silica particles, the CCM approach was not suitable. In fact, the functionalized materials are characterized by more complex synthesis procedure and for this reason one of the possible solutions is the use of the PCM method. In order to understand the interaction between the BG and the MCM-41, also BG pellets were coated with ordered mesoporous silica particles and heat treated using the similar schedule used for the MCM-41_BG scaffolds.

5.2.4. Drug Release Test To load silica particles with a drug, Ibuprofen (>98%, purchased from Sigma-Aldrich) as model drug was dissolved in hexane (33 mg mL−1) and MCM-41 particles were added to the drug solution (33 mg mL−1) at RT following the procedure reported in a previous work92. The samples with the drug solution were then placed in a vacuum hood at RT at 300 mbar for 10 min in order to enhance the drug infiltration inside the mesoporosity. After 12 h, this procedure was repeated, the drug solution was removed and the particles were dried in a vacuum hood at RT. All the particle synthesis solutions were tested for their drug-releasing capability. The scaffolds coated with MCM-41_B were in contact with the drug solution for 3 days (10 mg mL−1) before the starting of the drug-release test, in order to get a better infiltration of the drug. Only the scaffolds coated with MCM-41_B particles were considered for the test, because the resulting particles were still spherical and with ordered mesoporosity and the results were compared with those obtained on scaffolds not coated with MCM-41. The drug-release kinetics from all samples, particles (10 mg each sample), and scaffolds was assessed by soaking the samples in 4 mL of PBS, kept at 37°C until in a shaking incubator (90 rpm) the complete release of ibuprofen. At every time point, 1 mL of solution was uptake for the drug release analysis and substituted with 1 mL of fresh PBS. A UV-vis spectrophotometer was used to evaluate the amount of released drug. The calibration curve was calculated using a solution of ibuprofen in PBS with different known concentrations, on the basis of the absorption at 273 nm, typical of this molecule92.

5.2.5. Bioactivity and Stability of the Composite System Simulated body fluid (SBF) was prepared by dissolving reagent grade 8.035 g L−1 NaCl, −1 -1 −1 −1 0.355 g L NaHCO3, 0.225 g L KCl, 0.231 g L K2HPO4 (3H2O), 0.311 g L MgCl2 −1 −1 (6H2O), 0.292 g L CaCl2, and 0.072 g L Na2SO4 in deionized water and buffered at pH

−1 7.4 at 36.5°C with 6.118 g L tris(hydroxymethyl) aminometh-ane ((CH2OH)3CNH2) and 1M HCl, as previously reported by Kokubo and Takadama130. MCM-41_B and MCM-41_Ref particles were immersed in SBF at a 1.5 g L−1 ratio37,152. The specimens were kept in a polypropylene container at 37°C in incubator on an oscillating tray (90 rpm) for up to 28 days. The solution was not renewed152 and a falcon tube containing SBF as a control was also used for the entire period of the experiment, in order to control overtime the stability of the testing solution. Ag_MCM-41 particles were also tested in SBF with the same conditions. At the end of the incubator period, the particles were washed twice with deionized water, dried at 60°C overnight, and stored for further characterizations. Cylindrical BG-PU scaffolds coated with MCM-41_B by CCM method and not coated with mesoporous silica particles were first immersed in Tris-buffered solution (tris(hydroxymethyl) aminomethane, TRIS) in order to evaluate the stability of the MCM-41 coating. Only the scaffolds coated with MCM-41_B were tested, because the silica particles showed suitable features in terms of homogeneous coating, shape, and ordered mesoporosity. In both cases, the solution was kept in a polystyrene container at 37°C in a shaking incubator (90 rpm) for up to 1 week. The solution was renewed every 2 days in order to better mimic the in vivo behaviour, as carried out also in previous studies15. At the end of the incubator period, the foams were washed with deionized water, dried at RT for 3 days, and stored for further characterizations. After confirming the stability of the MCM-41 coating on the surface of the BG scaffolds, the samples were tested in SBF, using the same incubation parameters applied for the test with TRIS solution. As for the particles, also for the coated scaffolds a FalconTM tube containing SBF as a control was also used for the entire period of the experiment.

5.2.6. Characterization Techniques The shape and the surface structure of the resulting MCM-41 particles and MCM-41 coated BG scaffolds were evaluated by means of scanning electron microscopy (SEM-EDS). The porous structure of the particles was assessed with high resolution transmission electron microscopy (HRTEM). The pore diameter analyses were conducted on HRTEM images with ImageJ analysis software (Java). Small angle X-ray diffraction (SAXRD) was used to analyse the porous structure and the pore diameter of the silica particles. Nitrogen adsorption desorption analysis was conducted to assess the specific surface area and the pore size of the particles. The specific surface area and pore size of MCM-41 microspheres were evaluated, respectively, with BET method and BJH method. The compression strength of the BG scaffolds coated with the ordered mesoporous silica particles was also evaluated by compression test. In order to better distinguish between the effect of the coating and the

85 effect of the double thermal treatment (DTT) on the mechanical behavior of the resulting scaffolds, uncoated BG scaffolds underwent the same thermal treatment of coated scaffolds. All the details about the characterization techniques, brands and samples preparation were reported in Chapter 3.

5.3. Results

5.3.1. Optimization of MCM-41Particles Production Process The morphology and the microstructure of the obtained MCM 41 particles were assessed by HRTEM micrographs shown in Figure 5.2. HRTEM images of samples MCM-41_A and MCM-41_B showed the existence of highly ordered hexagonal array and streaks structural features (Figures 5.2a–d). The hexagonal array and the streaks are the view of the crystals whose axes are, respectively, parallel and perpendicular to the line of vision. Sample MCM- 41_C, which was prepared with a high concentration of ethanol in the synthesis solution, was porous however the porosity was not ordered (Figures 5.2e,f). Moreover MCM-41_D particles were porous but the porosity was not completely ordered, in contrast with the results reported in literature90 (Figure 5.2g,h).

Figure 5.2 HRTEM images of sample MCM-41_A (a,b) and sample MCM-41_B (c,d) which were characterized by high ordered mesoporosity; sample MCM-41 (e,f) and sample MCM-41_D (g,h), which were characterized by a disordered porosity. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

From the analysis of the HRTEM images with ImageJ analysis software, the dimension of the pores was evaluated, which was found to be around 3 nm for all samples (Figure 5.3). The analysis has been done applying the Fast Fourier Transform (FFT) and the inverse FFT

(Figure 5.3a) to the image, and the plug in plot (Figure 5.3b) was used to evaluate the distance between the pore channels.

Figure 5.3 High-resolution image of the ordered mesoporous structure of MCM-41_B after analysis with FFT and inverse FFT (a); plot of the distance between the parallel pore channels obtained with ImageJ plug in Plot Profile along the yellow line (b). (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale- Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

The pore size dimensions were confirmed also by SAXRD analysis. The spectra of sample MCM-41_A exhibited three sharp peaks, called Bragg peaks, indicating the long-range order present in the material, which is typical of MCM-41 materials90,92 (Figure 5.4a) in agreement with literature90. These peaks arise from the quasi-regular arrangement of the mesopores in the bulk material90. The Bragg peaks can be indexed assuming a hexagonal symmetry. 2 theta values of sample MCM-41_A namely 2.75, 4.65, and 5.10 can be indexed as (100), (110), and (200) reflections, respectively. These values were close to those reported by Grün 90 et al. . The repeating distance, a0, between two pore centers may be calculated by a0 =

(2/√3)d100. The pore diameter can be evaluated from a0 subtracting 1.0 nm, which is approximately the value of the pore wall thickness90. For MCM-41_B particles, it was possible to identify unequivocally only the main peak (100) (Figure 5.4b), meanwhile the 110 and 200 peaks were less pronounced but still visible. The SAXRD results combined with HRTEM results confirm thus the mesoporous ordered structures of MCM-41_A and MCM 41_B particles. For samples MCM-41_C and MCM-41_D, only the main peak (100) was identified, in agreement with the HRTEM analysis (Figures 5.4c,d). The first peak is in

88 fact an indicator of the presence of mesoporosity in the sample. Also in this case, it was possible to evaluate the pore diameter with the Bragg’s law and the resulting values were in agreement with the ImageJ analysis.

Figure 5.4 SAXRD analysis of samples MCM-41_A (a), MCM-41_B (b), MCM-41_C (c) and MCM-41_D (d). MCM-41_A particles were characterized by the three peaks, labelled 100, 110 and 200. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

The nitrogen isotherms of sample MCM-41_B are shown in Figure 5.5. The isotherms can be classified as type IV isotherms according to the IUPAC nomenclature for MCM-4142 which is typical of mesoporous materials with pore diameter in the range of 2–10 nm43. MCM 41_B particles were characterized by a specific surface area of 951 m2 g−1 and a pore volume of 0.24 cm3 g−1.

Figure 5.5 Nitrogen adsorption desorption isotherm of MCM-41_B particles. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

From Figure 5.6, it was possible to observe how the different amounts of solvent influenced the final shape and mesostructure of the resulting MCM-41 particles. The particles produced with only deionized water (Figures 5.6a,b) as solvent were characterized by hexagonal and not spherical geometry (MCM-41_A). Progressively increasing the amount of ethanol as co- solvent, it was possible to produce spherical particles, which exhibited a fairly homogeneous distribution of particle size but reduced ordered mesoporosity.

Figure 5.6 HRTEM images of the porous structure of samples prepared with different water/ethanol ratios: (a-b) sample MCM-41_A, no ethanol and 20 min of stirring following the standard synthesis procedure, (c-d) sample MCM-41_B 60% of ethanol and 20 min of stirring, (e-f) sample MCM-41_C, 90% of ethanol and 2 h of stirring. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

5.3.2. Drug Release of MCM-41 Particles The drug-release capability of the different mesoporous silica particles was evaluated and the released profiles are reported in Figure 5.7. Both samples MCM-41_A (Figure 5.7a), and MCM-41_D (Figure 5.7d) were characterized by a burst release and it was confirmed that after one hour of test already the 80% of the loaded drug was released. The rest 20% of the drug was released within the next 7 days. The particles prepared with solutions MCM-41_B and MCM-41_C showed the best drug release profile (Figure 5.7b-c). In particular, MCM- 41_B, which was characterized by high ordered mesoporosity, did not show any burst release during the first hours of the test. 80% of the loaded drug was in fact released only after 30 hours and the rest of the ibuprofen was released within the 7 following days.

Figure 5.7 Ibuprofen released profile form MCM-41_A (a), MCM-41_B (b), MCM-41_C (c) and MCM-41_D (d) particles. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

5.3.3. Bioactivity and Degradability of MCM-41 Particles in SBF

i. Surface Modification Particles in as-synthesized condition and after immersion in SBF were analysed by means of SEM in order to evaluate how the surface of the materials was affected by the interaction with simulated body fluid. The ordered mesoporous silica particles synthesized with the new method (MCM-41_B) were spherically shaped with an average diameter of 300 nm (Figure 5.8a) and a highly porous surface (pores in a range 10-30 nm). They were larger compared to the reference particles shown in figure 6d, but also the MCM-41_Ref were not homogeneously distributed. An interesting point is that the MCM-41_B particles were agglomerated; instead the MCM-41_Ref looked well dispersed. The particles were linked together by small “arms” visible in figure 8d. After 1 day of immersion in SBF, a clear modification of the particles surfaces was visible (Figure 5.8b,e). In case of MCM-41_B particles (Figure 5.8b) the surface porosity increased and some particles, especially the smaller ones, broke. MCM-41_Ref particles (Figure 5.8e), which were characterized by a smooth surface, after the immersion in SBF, they were also characterized by highly porous surface and they agglomerated. The links that were connecting the particles to each other were degraded. After 3 days, no further changes on the surfaces were detected.

Figure 5.8 SEM micrographs of MCM-41 particles as synthetized (a), after the immersion in SBF for 1 day (b) and 3 days (c); SEM micrographs of MCM-41_Ref particles as synthetized (d), after the immersion in SBF for 1 day (e) and 3 days (f).

SEM-EDS analysis (Figure 5.9) was performed in order to confirm that no precipitation of hydroxycarbonate apatite (HCA) occurred during the immersion in SBF. The analysis confirmed that no variation in the chemical composition of the surface of the particles was observed between MCM-41_B MCM-41_Ref particles, also after 28 days of test. It was possible to detect only Si and O, but no P or Ca ions related to the formation of Ca-P rich phase.

Figure 5.9 SEM-EDS analysis of MCM-41_B particles before (a) and after (b) the immersion in SBF for 28 days; SEM-EDS analysis of MCM-41_Ref particles before (c) and after (d) the immersion in SBF for 28 days.

ii. Ordered Mesostructure Modification The highly order of the mesoporous structure of the synthetized particles was assessed by HRTEM analysis (Figure 5.10). Both types of particles were characterized by high ordered mesoporosity (Figures 5.10a,e). The images showed the existence of ordered hexagonal array and streaks structural features. There are the view of the crystals whose axes are, respectively, parallel and perpendicular to the line of observation. After immersion in SBF, the ordered structure of the materials was still visible also after 28 days of test (Figures 5.10 b-d, f-h) although a clear degradation, especially of the external surfaces, was detected over time.

Figure 5.10 HRTEM images of the MCM-41_B particles as synthetized (a) and after the immersion in SBF for 3 days (b), 14 days (c) and 28 days (d); HRTEM images of MCM-41_Ref as synthetized (e) and after the immersion in SBF for 3 days (f), 14 days (g) and 28 days (h).t u

t HRTEM images b-d and f-h were taken by Dr. Ana M. Beltrán at the Institute de Ciencia de Materiales in Spain. u HRTEM image e was taken by Anahí Philippart, at the Institute of Biomaterials, Erlangen

The plot profiles of the distance between the pore channels of the particles before and after the immersion in SBF, obtained by the analysis of the HRTEM images with ImageJ plug-in, are reported in Figure 5.11. The particles were characterized by a constant distance between the pores. After immersion in SBF, the distance increased due to the degradation of the material and also some order was lost. The MCM-41_B particles were characterized by a pore size of ~ 3.2 nm, which increased up to 3.7 nm already after 3 days of test and it stabilized to that value. The pore size of the MCM-41_Ref particles was more affected by the degradation. The initial pore size of the MCM-41_Ref was 3.6 nm and after immersion in SBF for up to 28 days the pore diameter increased up to 7 nm.

Figure 5.11 Plot profile of the distance between the pore channels characteristic of the MCM-41_B and MCM- 41_Ref particles before and after immersion in SBF for up to 28 days.

iii. FT-IR Analysis In Figure 5.12, the spectra of the FTIR analysis of the MCM-41_B particles before and after the immersion in SBF for up to 28 days was reported. All the peaks related to the Si-O bonds were detected and they did not change overtime. Typical bands centered at 560 and 600 cm-1, ascribable to the P-O asymmetric bending related to the formation of a Ca-P rich layer (amorphous or already developed HCA), were not detected in the analyzed samples at any time point111,153.

Figure 5.12 FTIR spectra of MCM-41_B particles before and after the immersion in SBF for up to 28 days. In the figure the peaks associated to the Si-O stretching mode are marked.

iv. Ion Release and pH Variation The pH of the SBF solution used as a control and of all samples was monitored during the entire test (Figure 5.13a). The starting pH of the SBF was around 7.4; due to the degradation of the silica particles the pH increased up to a maximum value of 7.65 for the MCM-41_B particles and 7.6 for the MCM-41_Ref materials. Also the pH of the SBF control increased overtime almost up to 7.65, most probably due to the aging of the solution, which has the property to be stable at 4°C for ~ 1 month. The Si ions release is reported in Figure 5.13b. The MCM-41_B particles showed a burst release of Si ions during the first day of test. The MCM-41_Ref showed a slower kinetic: the highest release of Si ions was reached after 2 days. Both the materials, although characterized by different particles size, pore size and surface porosity, released the same amount of soluble silica 400 mg mL-1, which correspond to ∼ 260 mg SiO2/g material.

Figure 5.13 pH variation of the SBF containing the silica particles and the control for up to 28 days of test (a); ICP analysis of the Si ion release for up to 7 days (b) for MCM-41_Ref and MCM-41_B samples.v

5.3.4. Ag Nanoparticles-doped MCM-41 particles The morphology and the surface of the Ag_MCM-41_B particles prepared with different concentration of silver nitrate were assessed by SEM and compared to the non- functionalized MCM-41_B particles (Figure 5.14). The starting MCM-41_B particles were characterized by a well-developed surface mesoporosity, which was not affected by the functionalization with silver. Using a higher concentration of AgNO3, the number of silver nanoparticles on the surface of the MCM-41 particles increased visibly. In case of -1 Ag_MCM-41 prepared using 6 mg mL of AgNO3, the silver nanoparticles are seen to be placed inside the surface mesoporosity of the silica particles (Figure 5.14d). The silver nanoparticles were homogeneously distributed on the MCM-41_B surface and not agglomerated, as shown in the backscattering image reported in Figure 5.14e. The FTIR analysis, reported in Figure 5.14f, showed that the functionalization with silver does not affect the silica network.

v The ICP-OES analysis was done at the Institute of Particle Technology (LFG) by Dr. Jochen Schmidt.

Figure 5.14 SEM micrographs of MCM-41_B particles (a) and Ag_MCM-41_B particles obtained with different concentrations of silver nitrate: 0.6 mg mL-1 (b), 3 mg mL-1 (c) and 6 mg mL-1 (d). Backscattering SEM micrographs of Ag_MCM-41_B particles obtained using the highest concentration of silver nitrate. FTIR spectra (f) of the non-functionalized and the Ag-functionalized MCM-41_B particles obtained with different silver nitrate concentrations.

The morphology of the non-functionalized and of the Ag-functionalized MCM-41_B particles was assessed by TEM images (Figure 5.15a,c). HRTEM images showed the existence of highly ordered hexagonal array structures, as shown in Figure 5.15b,d. The incorporation of Ag-nanoparticles in the MCM-41 material did not affect the ordered mesoporous structure. Moreover, the presence of black spots, most probably Ag- nanoparticles, dispersed in the silica network was clear in Figure 14c. These nanoparticles looked well dispersed and not agglomerated. From the analysis of the HRTEM images with ImageJ analysis software (NIH, USA), the dimension of the pores was evaluated, which was found to be ~ 3.3 nm for both samples, functionalized and non-functionalized. As before, the analysis was done applying the Fast Fourier Transform (FFT) and the inverse FFT (Figure 5.15b,d) to the images, and the plug in plot profile (Figure 5.15e,f) was used to evaluate the distance between the pore channels.

Figure 5.15 HRTEM images of MCM-41_B (a) and Ag_MCM-41_B (c) particles, which were characterized by high ordered mesoporous structure. High-resolution images of the ordered mesoporous structures of MCM-41_B (b) and Ag_MCM-41_B (d) particles after analysis with FFT and inverse FFT. Plot of the distance between the pore channels obtained with ImageJ plug in plot of MCM-41_B (e) and Ag_MCM-41_B (f) particles.w

The black spot on the silica particles surface (Figure 5.16a) were analysed with EDS (Figure 5.16b), which confirmed that were silver nanoparticles. The oxygen signal was very low, confirming that most of the formed silver nanoparticles were pure silver and not silver oxide. In Figure 16c the silver nanoparticles size distribution is reported. The 50% of the silver nanoparticles formed during the TIE process was characterized by a diameter of 5-10 nm and the other 40% of them were characterized by a diameter of 10-20 nm. The presence of silver nanoparticles with a diameter higher than 5 nm was also confirmed by means of XRD analysis (Figure 5.16d). After the TIE process, 4 peaks were visible on the XRD spectra, which were related to silver.

w HRTEM images were taken by Dr. Ana M. Beltrán at the Institute de Ciencia de Materiales in Spain.

Figure 5.16 HRTEM image of Ag-MCM-41_B particles (a); EDX analysis of a single silver nanoparticle on the surface of a MCM-41_B particle (b); silver nanoparticles size distribution (c); XRD analysis of functionalized Ag_MCM-41_B and not functionalized MCM-41_B mesoporous silica nanoparticles.x

The specific surface area of the Ag_MCM-41_B particles prepared with the higher concentration of silver nitrate during the TIE process was still characterized by high specific surface area (BET specific surface area 1021 m2 g-1) compared to the non-functionalized MCM-41_B ones (BET specific surface area 1102 m2 g-1).y

i. Degradability of Ag_MCM-41 Particles In Figure 5.17, SEM micrographs of Ag_MCM-41_B particles before and after the immersion in SBF for 1, 3 and 7 days are reported. Already after 1 day of immersion, the particles showed a marked modification of the surface which can be associated with the degradation of the silica particles. After 3 and 7 days it was not possible to observe a significant change of the silica particles surface compared to 1 day of immersion. Also after 7 days of test, no hydroxyapatite deposition was observed. This was also confirmed by EDS analysis, reported in Figure 5.17e,f. No signals coming from Ca or P are detected, only the

x HRTEM image and EDS analysis taken by Dr. Ana M. Beltrán at the Institute de Ciencia de Materiales in Spain. y The BET measurements were done at the Institute de Ciencia de Materiales in Spain by Dr. Ana M. Beltrán.

100 peaks of Si and O2 were visible. Moreover, the silver nanoparticles were not visible on the surface of the MCM-41 already after 1 day in SBF.

Figure 5.17 SEM images of Ag_MCM-41_B particles before and after immersion in SBF solution for up to 7 days (a-d); EDX analysis of Ag-MCM-41_B particles before and after 7 days of immersion in SBF.

No presence of hydroxycarbonate apatite (HCA) formation was also confirmed by mean of FTIR analysis as reported in Figure 5.18. No peaks related to the P-O and C-O bending and stretching vibrations were identified, confirming that no HCA was formed also after 7 days in SBF.

101

Figure 5.18 FTIR spectra of Ag_MCM-41_B particles before and after 7 days of immersion in SBF.

The Si and Ag ions release is reported in Figure 5.19a,b. The Ag_MCM-41_B particles showed a burst release of Si ions during the first 24 h of test. Also all the Ag incorporated in the particles was release immediately after the immersion in SBF. The pH of the SBF solution used as a control and of Ag_MCM-41_B particles was monitored overtime (Figure 5.19c). The starting pH of the SBF was around 7.4; due to the degradation of the silica particles and the release of silver, the pH increased up to a maximum of 7.65. Also the pH of the SBF control increased overtime, most probably due to the aging of the solution.

102

Figure 5.19 Concentration of (a) Si and (b) Ag ions released from Ag_MCM-41_B particles for up to 7 days of immersion in SBF; (c) pH variation after the immersion of functionalized and non-functionalized MCM-41_B in SBF.z

The Ag_MCM-41_B particles were loaded with ibuprofen in order to evaluate the capability of the system to incorporate a drug and to evaluate if the Ag presence on the particles could affect the up-take. The Ag_MCM-41_B particles showed a slower release kinetic compared to the non-loaded MCM-41, but the total amount of loaded drug was almost the same for both types of particles.

z ICP analysis was done at LFG Institute in Erlangen by Dr. Jochen Schmidt.

103

Figure 5.20 Ibuprofen release profile from MCM-41_B and Ag_MCM-41_B particles for up to 14 days.

5.3.5. MCM-41 Particles Coating of BG Scaffolds The synthesis of the MCM-41 particles directly inside the BG scaffolds (CCM method) is a technique already developed in literature63. For this reason, the first experiments on the coating of the scaffolds with these ordered mesoporous silica particles were done following this procedure in order to evaluate the quality and stability of the coatings and to access the effect of the MCM-41 particles on the bioactivity of the composite scaffolds. From SEM analysis it was possible to observe that the surface of the scaffolds was completely coated after immersion in the MCM-41 synthesis batch (CCM method) maintaining an open porosity. In the case of the synthesis solution of MCM-41_A (Figure 5.21a-b), the one without ethanol as co-solvent, the shape of the resulting MCM-41 particles was seen to have changed. The presence of the scaffold affected the formation of the particles, probably due to a reduction in the homogeneity of the solution. Moreover, using this solution, it was not possible to obtain a homogenous coverage of the BG scaffold surface. With the synthesis solution of samples MCM-41_B (Figure 5.21c-d) and MCM- 41_D (Figure 5.21e-f) the resulting particles on the surface of BG scaffolds were spherical and they covered completely the surface of the scaffold struts. By means of HRTEM analysis it was also possible to confirm that the MCM-41_B particles still exhibited ordered mesoporosity, as shown in Figure 5.22.

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Figure 5.21 BG-PU scaffolds coated with MCM-41_A (a-b), MCM-41_B (c-d) and MCM-41_D (e-f) particles by CCM procedure. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

Figure 5.22 HRTEM images of MCM_41_B particles coating the BG-PU scaffolds at different magnifications in order to show that the ordered mesoporous structure of the material was not affected by the CCM method. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

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In order to investigate the stability of the coatings and to assess whether or not the presence of the MCM-41 particles on the scaffolds would affect the bioactivity of the BG, the BG-PU scaffolds coated with MCM_41_B by CCM method were immersed in TRIS buffered solution for 1 week. After one week of immersion it was possible to observe that the presence of the silica particles has not affected the bioactivity of the scaffolds. In fact on both samples, coated and not coated, a surface hydroxycarbonate apatite (HCA) layer was seen to from (Figure 5.23c-d). Moreover it was possible to confirm the stability of the MCM-41 coating: after 10 days in TRIS buffered solution it was possible to identify the layer of MCM-41 particles covered with the HCA deposit (Figure 5.23e). The HCA layer was well developed and the surface of the particles was more porous (Figure 5.23f-g, black circle).

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Figure 5.23 SEM micrographs of BG-PU scaffolds uncoated (a) and coated with MCM-41_B (b-d) after 1 week in SBF; and of MCM-41_B_BG-PU scaffolds, after immersion in TRIS buffered solution for 10 days (e-g). (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

In order to prove the stability of the coating and the absence of a negative effect on the scaffold bioactivity of the starting BG, further experiments were conducted. Also BG-SA scaffolds were coated with MCM-41_B particles by the CCM procedure in order to evaluate the possibility to extend the process also to scaffolds with lower total porosity compared to standard foams prepared using PU foams as sacrificial template. From SEM analysis of CCM_MCM-41_BG-PU (Figure 5.24c,d) and CCM_MCM-41_BG-SA (Figure 5.24a,b) scaffolds, it was possible to observe that the external surfaces of the samples were homogeneously coated with the ordered mesoporous silica particles, maintaining an open

107 porosity. In case of the scaffolds prepared with PU foam as template, also the inner core of the samples looked completely covered with MCM-41 particles (Figure 5.24d). This was not the case for CCM_MCM-41_BG-SA samples; in fact it was not possible to identify a layer of particles on the inner core surface of the scaffolds (Figure 5.24b). In both cases, the particles forming the coatings were spherical.

Figure 5.24 SEM micrographs of CCM MCM-41_B_BG-SA external (a) and internal (b) surfaces and of CCM MCM-41_B_BG-PU external (c) and internal (d) surfaces.

A first attempt to modify the coating procedure was to use the as synthetized particles to impregnate the BG-based scaffolds. The SEM micrographs of the scaffolds obtained using different particles concentrations (wet weight) are reported in Figure 5.25. No significant -1 differences are observed for the different coating concentrations, only if the 0.3 gwet mL was used the porosity of the scaffolds was blocked (Figure 5.25e). The main problem of the system PCM_A was the low reproducibility.

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Figure 5.25 SEM micrographs of BG-PU scaffolds coated with MCM-41_B using the Post Coating Method A (PCM_A method, wet particles) using different concentration of wet particles in ethanol.

For this reason, for the PCM_B method, the dried particles were used to prepare the impregnation solution. Different particles concentrations were tested in a trial-and-error approach in order to determine the right one, which should result in homogeneous external and internal coatings but without affecting the open porosity characteristic of these samples. In Figure 5.26, the SEM micrographs of the PCM_B_MCM-41_B_BG-PU scaffolds are reported. Using 10 mg mL-1 the surface was not coated; increasing up to 100 mg mL-1 the pores were almost completely blocked. A good compromise between homogeneity of the coating and open porosity was reached using a concentration of 50 mg mL-1 for the coating suspension.

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Figure 5.26 SEM micrographs of BG-PU scaffolds coated with MCM-41_B particles PCM_B method obtained using different concentrations for scaffolds impregnation.

With this concentration also PCM_B_MCM-41_B_BG-SA scaffolds were prepared (Figure 5.27a,b) and a homogeneous coating inside the scaffolds and on the external surface of the samples was obtained.

Figure 5.27 SEM micrographs of external (a) and internal (b) surface of BG-SA scaffolds coated with MCM- 41_B using PCM_B method.

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i. Mechanical Properties The mechanical properties discriminate the mechanical properties changes due to the functional coating from the effect of the double thermal treatment (DTT) on the starting BG scaffold. The DTT on BG-PU scaffolds had a positive effect on the mechanical properties of these scaffolds and the maximum compressive strength increased up to 0.027 ± 0.004 MPa. Also the effect of the MCM-41_B particles coating on the BG-PU scaffolds compressive strength obtained with CCM and PCM_B methods was considered. If the coating is applied using the CCM, the maximum compressive strength increased up to 0.04 ± 0.01 MPa, compared to the 0.16 ± 0.03 MPa reached using the PCM_B method. The DDT on the BG- SA scaffolds did not produce any significant change in the compressive strength of the samples, which was ~ 4 MPa before and after the DTT. The coating with the MCM-41 particles, using both CCM and PCM_B methods, increased 3 times the maximum compressive strength of the scaffolds. The obtained values were 11 ± 1 MPa and 12 ± 5 for CCM_MCM-41_BG-SA scaffolds and PCM_B_MCM-41_BG-SA, respectively. All the values are summarized in Table 5.2. of the coated scaffolds were compared with those of the non-coated scaffolds, which were heat treated at 550 °C for 6 h. The idea was to evaluate the effect of the coating on the mechanical properties.

Table 5.2 Summary of the maximum compressive strength of the DTT and coated scaffolds.

Maximum compressive Strength [MPa] DTT CCM PCM_B BG-PU 0.027 ± 0.004 0.04 ± 0.01 0.16 ± 0.03 BG-SA 4.1 ± 0.2 11 ± 1 12 ± 5

ii. Bioactivity in SBF The scaffolds were tested in SBF in order to evaluate the bioactivity in the most common used testing solution130. In Figure 5.28, SEM images of the four scaffolds types, i.e. CCM MCM-41_B BG-SA and BG-PU scaffolds and PCM_B MCM-41_B BG-SA and BG-PU scaffolds, after immersion in SBF for 1 day are reported. Already after 24 hours the surface of all scaffolds was covered in already fully developed HCA. The typical cauliflower structure composed by needle-like structures of the HCA phase is visible on all samples. It was also possible to observe that the MCM-41 particles were not only still on the surface of the scaffolds, but they were incorporated in the developing HCA.

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Figure 5.28 SEM micrographs of CCM MCM-41_B BG-SA (a) BG-PU (b) scaffolds and PCM_B MCM-41_B BG-SA (c) and BG-PU (d) scaffolds after immersion in SBF for 1day.

In Figure 5.29, SEM images of a PCM_B MCM-41_B BG-SA scaffold cross section strut after 3 days in SBF are reported. The strut is seen to be completely covered in HCA (Figure 5.29a). Figure 5.29b,c shows two SEM images at high magnification of the HCA layer cross section (white circle). These images confirm that the MCM-41 particles have been incorporated in the developing HCA layer during SBF immersion.

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Figure 5.29 SEM micrographs of a cross section of a PCM_B MCM-41_B BG-SA scaffold after the immersion in SBF for 3 days at different magnifications showing a) that the strut is seen to be covered in HCA and that b-c) the MCM-41_B particles have been incorporated in the developing HCA.

In Figure 5.30 the FTIR spectra of CCM MCM-41_B BG-SA and BG-PU scaffolds and PCM_B MCM-41_B BG-SA and BG-PU scaffolds before and after the immersion in SBF for up to 7 days are summarized. All the samples, before the immersion in SBF, showed the characteristic peaks related to the Si bonds, i.e. Si-O-Si asymmetric stretching mode at 1000- 1200 cm-1 26, Si-OH stretching mode at ~ 950 cm-1 26, Si-O symmetric stretching at ~700 cm1 26 and Si-O-Si bending mode at ~ 475 cm-1 26. For all the samples, the main peaks related to the Si-O bonds decreased overtime and some of them disappeared completely. It is also possible to observe, that already after 1 day of test all the characteristic peaks ascribed to the HCA layer formation change at around 600 cm-1 26, corresponding to the bending mode of P- O (crystalline phosphate). The spectra also showed the narrowing of the band at around 800 cm-1 corresponding to the bending mode of C-O and finally the manifestation of the stretching mode of C-O at around 1400 cm-1 26. Only the CCM MCM-41_B BG-PU scaffolds showed the characteristic peaks associated with the formation of HCA only after 3 days of test.

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Figure 5.30 FTIR spectra of CCM MCM-41_B BG-PU and BG-SA scaffolds and PCM_B MCM-41_B BG-PU and BG-SA scaffolds before and after the immersion in SBF for up to 7days. The relevant peaks are discussed in the text.

iii. pH Variation The pH variations in SBF during the 7 days of bioactivity test are reported in Figure 5.31. The scaffolds dissolution took place immediately after the immersion in SBF, as shown by the pH increase and in agreement with the SEM analysis. CCM_MCM-41_B BG-PU samples led to a pH increase up to 8 in 2 days. For all the other samples, the highest value of pH ∼ 8 was reached only after 7 days of immersion in SBF.

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Figure 5.31 pH variation of after immersion in SBF for up to 7 days: CCM MCM-41_B BG-SA, CCM MCM- 41_B BG-PU, PCM_B MCM-41_B BG-SA, and PCM_B MCM-41_B BG-PU scaffolds.

iv. Drug Release Capability The amounts of released ibuprofen from the BG and BG_MCM-41_B scaffolds are shown in Figure 5.32. The presence of the mesoporous silica particles increased the drug incorporation capability, but in both cases most of the drug was released during the first hours of the test.

Figure 5.32 Drug-release profile from BG-PU scaffolds coated with ordered mesoporous silica particles compared to non-coated scaffolds. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale- Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology)64.

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5.4. Discussion One of the most investigated areas in the bone tissue engineering field is related to the development and characterization of mechanically robust and porous 3D scaffolds. The main challenge is the design of a material able to match at the same time the biological and mechanical properties of the natural bone and also release biologically active ions or drugs able to reduce the risk of inflammation and infections after the implantation. In a previous work of Mortera et al.63, the possibility to increase the functionality of BG-based porous scaffolds was considered using a coating with MCM-41 particles as drug delivery system. In this way, it was possible to combine in a single system the drug uptake and release capability of mesoporous particles with the bioactivity of BG. In the present work, a further development of this idea was presented, improving the homogeneity of the coating, assessing the bioactivity and stability of the composite system MCM 41_BG and investigation of the possible positive effect on the mechanical properties of the starting BG scaffolds. Four different solutions were evaluated for the preparation of mesoporous silica particles, and an optimal synthesis procedure was found (MCM-41_B). In fact, by combining two different synthesis pathways, both well known in literature90,150, it was possible to obtain particles characterized by spherical shape and high-ordered mesoporosity, as confirmed by HRTEM, SAXRD, and nitrogen sorption measurements (specific surface area 951 m2 g−1, pore volume 0.24 cm3 g−1). Moreover, the efficiency of the synthesis was increased and the total amount of produced particles obtained per single batch increased up to 60% compared to the previous synthesis procedure: every 500 mL of solution, 4 g of MCM 41_B were produced. MCM-41_B showed the best drug-release profile not exhibiting any burst release during the first hours of the test. The 80 % of the loaded drug was in fact released only after 30 h and the rest of the ibuprofen was released within the seven following days. The drug-release times obtained during this work are in agreement with previous studies on drug-release capability of mesoporous silica particles92.

The optimized MCM-41 particles (MCM-41_B) were characterized after the immersion in SBF for up to 28 days and compared to MCM-41_Ref particles which were produced following one of the most used synthesis routes known in literature149. After the immersion in SBF, the already porous surface of these MCM-41 particles increased the roughness even more and some of the particles broke. The main effects on the surface were observed during the first day of test and after that no other significant changes were identified. The same behavior was observed for the MCM-41_Ref particles used as a reference. The smooth surface that characterized the MCM-41_Ref particles was affected by the SBF and the surface became porous and extremely rough. Moreover, the MCM-41_Ref particles have

116 seen to have lost their feature of being homogeneously dispersed immediately after 1 day in SBF, i.e. the connection that linked them one to each other dissolved and the particles agglomerated. The ordered mesoporous structure of the MCM-41_B particles was not significantly affected by the immersion in SBF also after 28 days. In fact, from the HRTEM images and the plot of the pore channel distances, it was possible to observe that the ordered mesostructure was preserved and that the pore size increased only from 3.2 nm to 3.6 nm. A stronger impact of the SBF was observed for the MCM-41_Ref particles. The particles used as a reference maintained the ordered structure, but the pore size increased up to 7 nm after 28 days in SBF. No changes in the chemical composition were observed for both types of particles and no precipitation of HCA was identified, as expected from literature111. The highest Si ion release took place during the first 24 hours for the MCM-41_B particles and after 48 hours for the reference particles. In both cases the amount of released Si ions was the same (~ 260 mg SiO2/g material). This value was slightly higher than the one reported by Izquierbo-Barba et al.154 for SBA-15 particles in SBF.

The first problem after an implantation is the exposure to inflammatory and infection risk which can cause implant failure or tissue necrosis146. A potential alternative is the incorporation of antiseptic ions into the MCM-41 particles. Ag has a broad-spectrum of antibacterial activity; therefore silver release is a wide investigated approach to impact the antibacterial properties155–157. In the present work, Ag nanoparticles-doped MCM-41 submicron particles (Ag_MCM-41_B) were developed and characterised. In order to introduce the Ag in the ordered mesoporous material, the template ion exchange method (TIE) was selected. TIE is not the only method available; also the direct hydrothermal (DHT) could be used to directly introduce inorganic species during the silica network formation. However, during DHT, the Ag can interact with the Br of the surfactant forming an AgBr precipitate, whose biocompatibility is not known158. For this reason the TIE method was preferred to the DHT. In fact in case of the TIE technique used in the present work, the silica particles were carefully washed before the introduction of the silver precursor, hence the unreacted bromide ions were removed from the material.

Different silver precursor concentrations were used for the functionalization of the silica 159 particles . Increasing the amount of AgNO3 used for the TIE, higher was the number of Ag nanoparticles decorating the surface of the MCM-41_B particles. The Ag nanoparticles looked well dispersed on the surface of the ordered mesoporous material and the presence of silver did not have a negative effect on the order of the silica network. This was confirmed by the HRTEM images and by the specific surface area analysis: the silver doped silica

117 particles exhibited a reduction of only 100 m2 g-1 in specific surface area compared to non- functionalized particles. This is also a confirmation that most of the Ag nanoparticles are located on the surface of the silica particles, and not inside the ordered porous channels of the MCM-41. The MCM-41 particles are characterized by a high surface roughness, with pores in a range between 10 and 30 nm, where the Ag nanoparticles were located after the TIE process. In fact, most of the Ag nanoparticles are characterized by a diameter in the range 5-50 nm. The formation and the size of the silver nanoparticles were also confirmed by the XRD analysis. The detection of Ag in the XRD analysis is possible only if it is present in the form of nanoparticles bigger than 5 nm, according to the Scherrer equation160. In order to assess the degradation of the silica particles and the release of the Ag nanoparticles, Ag_MCM-41_B particles were immersed in SBF for up to 7 days. From the SEM micrographs it was observed that after 1 day in SBF it was not possible to identify the Ag nanoparticles on the surface of the silica particles. The fast release of the Ag was confirmed also by mean of ICP results. Almost 100% of the detected Ag was released during the first 24 h of the test. ICP-OES results allowed the quantification of the amount of Si released from the silica particles during the first week in SBF. Most of the silica release took place during the first 6 h of immersion in SBF, and the maximum released was reached after 24 h. This result is in agreement with the surface modification detected by SEM: the roughness of the MCM-41 increased visibly. The Si ions release will be a key element for the application of the silica particles as a functional coating of BG scaffolds. In fact, Si ions play an essential role in the osteoblast gene activation enhancing the cell proliferation and differentiation9. The Ag_MCM-41_B particles were also loaded with ibuprofen in order to assess their drug up-take and release capability. The functionalized particles were able to incorporate almost the same drug amount of the non-functionalized particles, confirming that incorporation of Ag particles did not have a negative effect on the silica particles as a drug carrier. From the preliminary test, it was observed that Ag_MCM.41 showed a better drug release kinetic and sustained released was obtained for up to 6 days. This novel synthesis solution for the production of MCM-41_B was used for the coating of 3D BG scaffolds. Initially, the adopted coating procedure involved the synthesis of the silica particles directly inside the BG-PU scaffolds (CCM method). Due to the high amount of particles produced during the synthesis, a highly homogeneous coating of the scaffolds was obtained. After the coating procedure, the particles were still spherical in shape and also the ordered mesoporosity was not affected. This was an improvement compared to previous works, in which the coatings were not homogeneously distributed on the surface of the scaffolds and the particles were not characterized by ordered mesoporosity63. The system BG_MCM-41 was assessed to be bioactive. After 1 week of immersion in TRIS buffered

118 solution on the surface of both coated and not coated scaffolds, a layer of HCA was observed. This behavior confirmed that MCM-41 particles did not have a negative effect on the bioactivity of the BG scaffolds; on the contrary, combined with BG, they seem to enhance the bioactivity. Moreover, most of the MCM-41 particles were still on the surface of the scaffolds indicating that MCM-41 particles were in fact adhered to the glass surface due to the thermal treatment: the calcination at 550 °C is likely to induce softening of the glass141 which should facilitate adhesion of the MCM-41 spheres. For this reason, the MCM-41 particles coating was stable on the surface of the BG scaffold also after immersion in SBF.

Confirmed that the ordered mesoporous silica particles coating was possible and did not have a negative effect on the bioactivity of the BG scaffolds, also the possibility to coat the foams after the particles synthesis was explored (method PCM). Two different procedures for the PCM method were developed: the PCM_A did not give reproducible results, because the wet weight of the particles was considered to prepare the coating solution. In order to get a better reproducibility, the PCM_B technique was developed using dried MCM-41_B particles for the preparation of the coating solution. The concentration of the coating solution which better combined the homogeneity of the coating with the open porosity of the resulting system was 50 mg mL-1. In this way, it was possible to obtain a homogeneous coating on both the external and internal surfaces of the BG-PU and BG-SA scaffolds without interfering with the high open porosity characteristic of the BG foams. The presence of the particles on the scaffolds had an extremely positive effect on the mechanical properties of the BG-SA scaffolds, resulting in a clear increase of the maximum compressive strength of the samples, up to 12 MPa for the CCM and PCM_B MCM-41_B BG-SA scaffolds. These values are 3 times higher than the maximum compressive strength that is possible to obtain without the coating, as reported in Chapter 4. In case of the BG-SA scaffolds, the DTT did not affect the mechanical behavior of the samples and for this reason the increase of the mechanical properties of the coated samples could be entirely attributed to the MCM-41 particles coating. In case of the BG-PU scaffolds, the PCM_B method allowed to increase the maximum compressive strength up to 0.16 ± 0.03 MPa, which is almost 6 times higher than the values obtained for the non-coated DTT BG-PU samples. The MCM- 41 particles coating applied on the BG-PU scaffolds using the CCM method did not give such good results as for the PCM_B method. In fact, the maximum compressive strength reached only 0.04 ± 0.01 MPa. A possible explanation is related to how the CCM method is performed and to the intrinsic fragility of the BG-PU scaffolds due to their high porosity (>93%, as reported in Chapter 4). The coating thus did not increase the mechanical properties because during the CCM procedure the scaffolds can become in contact with each

119 other leading to surface damage. The general improvement of the mechanical properties of the coated scaffolds with the ordered mesoporous silica particles indicated that, coupled with SEM observations, the coating was highly homogeneous. The reason why the MCM-41 particles were able to increase significantly the maximum compressive strength of the scaffolds is likely related to the interaction between the particles and the BG. The thermal treatment at 550 °C for 6 h led to softening of the BG, the particles adhered to the scaffolds surface filling the crack present in this type of scaffolds.

All the developed scaffolds were then tested in SBF for up to 7 days. The bioactivity of the combined system was already assessed in TRIS buffered solution for the CCM_MCM-41_B BG-PU samples, but the extremely high bioactivity observed in SBF was not predicted. In fact, already after a single day in SBF, all the samples showed the formation of a well- developed HCA layer on the surface of the scaffolds. For the non-coated BG-PU and BG-SA scaffolds immersed in SBF with the same testing conditions, it was possible to form such a well-developed layer of HCA only after 3 days in SBF, as reported in Chapter 4, Section 4.3.5. It is important to underline that during the entire period of the test, a falcon tube containing SBF from the same batch used for the scaffolds test was also incubated and no precipitation of the testing solution was observed. The high bioactivity of these combined systems was confirmed also by FTIR analysis. All the peaks related to the HCA formation and the progressive degradation of BG were present already after 1 day of immersion in SBF and it was possible to observe a progressive evolution of these two phenomena over time. It has been reported in literature that MCM-41 particles are not bioactive and formation of HCA on their surface was not observed after 2 months of immersion in SBF due to the small pore size and the lower concentration of silanol groups compared to other silica particles such as SBA-15 and MCM 48111. Similar results were obtained in the present work for both MCM-41_B and Ag_MCM-41_B particles, as described in previous sessions. However the combination of MCM-41 particles with BG scaffolds resulted in an increase of the bioactivity of the system. A possible explanation can be found considering the surface reactions occurring once BG is in contact with SBF: the first step is the rapid exchange of + 2+ + + Na and Ca of the BG with H and H3O ions from the SBF solution and the second step is the loss of soluble silica in the form of Si(OH)4 to the solution, resulting from the breaking of Si-O-Si bonds of the silica network and formation of Si-OH (silanols) on the glass surface (Chapter 2, Section 2.4.3). One of the features of the ordered mesoporous silica materials, as MCM-41 particles, is the presence of abundant silanol groups. For this reason the MCM-41 particles can act as nucleation sites for the HCA. This hypothesis was corroborated by the SEM images in which the silica particles were clearly included in the developing HCA

120 structure. Another interesting aspect was the pH variation over time. All the samples, except for the CCM MCM-41_B BG-PU samples, exhibited reduced increment of pH compared to the non-coated scaffolds (Chapter 4, Section 4.3.5). Also after 3 days in SBF, the pH values were < 8. This behavior could be due to the DTT, which stabilized the BG structure, and the presence of the silica particles acting as a barrier counteracting the fast degradation of the BG.

5.5. Conclusions Ordered mesoporous silica particles, MCM-41_B and Ag_MCM-41_B, were successfully developed combining two synthesis procedures reported in literature. The resulting particles were characterized by ordered mesoporous structure, highly porous surface and stability of the porous channels also after 28 days in SBF. These particles were used as a functional coating of BG-based scaffolds with two different procedures. The starting idea was to combine the drug up-take/ release capability of the silica mesoporous particles with the bioactivity of the BG scaffolds, however new findings were obtained. The mechanical properties of the coated scaffolds increased significantly and, surprisingly, also the bioactivity of the system was enhanced. In conclusion, it was possible to obtain a functional system with drug delivery capability, increased mechanical properties and enhanced bioactivity confirming the possibility to use these MCM-41_BG scaffolds for bone tissue engineering applications.

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6. Sol-Gel Glass Coating of BG-Scaffolds

6.1. Introduction Bioactive glasses (BGs) produced by melting have been used, in the past decades, in bone regeneration and tissue engineering applications as particulates and in monolithic form due to their degradation behavior, their excellent ability to form strong bonding with bone and soft tissues and their osteoproductive properties9,10,13,14,118. In fact, they are able to form a layer of carbonated hydroxyapatite (HCA) similar to the mineral phase of bone tissues in contact with biological fluids or when immersed in simulated body fluid (SBF)1. HCA, the main inorganic bone component, confers the specific biological response of bonding at the interface between tissue and BGs43. Ordered mesoporous sol-gel glasses (MBGs) are a new family of sol-gel glasses with tailored porosity at the nanometric scale98,113. They are characterized by the same composition of glasses but with ordered mesoporosity, opening the possibility to combine the bioactivity of bioactive glasses with the drug uptake/ release capability of mesoporous materials104,161. These nanostructured materials have a chemical composition similar to conventional sol-gel glasses and present a highly ordered mesoporous structure, a high surface area (up to 1000 m2 g-1), large pore volume (up to 1 cm3 g-1), regular and tunable mesopore size (2-50 nm)113. MBGs, thanks to the presence of this highly ordered structure, can play an important role as biomaterials; the inorganic matrix of mesoporous bioactive glass can be doped with therapeutic ions or drugs to limit bacterial infections, one of the main complications in bone surgeries148,162,163. In the last years MBGs with multifunctional properties, incorporating therapeutic ions (e.g. Ag+, Li+, Sr2+, Zn2+, Cu2+ and Ce3+) and various drugs have been investigated113. These studies have shown how MBGs can be efficient carries for the release of therapeutic ions and drugs at the same time. Such doped glasses possess antibacterial activity and positively stimulate angiogenesis, osteogenesis and cementogenesis148,162–170. 3D scaffolds based on MBGs have been developed but the resulting structures were characterized by low mechanical properties62,171. In the past years, the use of ordered mesoporous materials as functional coating for BG- scaffolds has been proposed, including MCM-41 particles obtained by direct synthesis method63,64 and mesoporous bioactive glass powder via electrophoretic deposition (EPD)172.

In the present work, an innovative coating method of BG scaffolds with ion doped MBGs is presented. Two different MBG compositions, Co-MBGs and SrCu-MBGs, which were developed by Philippart in 2016173, were synthetized. Co2+ ions have been reported to have osteogenic and angiogenic properties113, Cu2+ ions showed also antibacterial properties113,

123 instead Sr2+ can enhance osteogenesis and cementogenesis113. The range of substituent´s concentrations was chosen to not inhibit or decrease the bioactivity of the starting glass composition174,175. The metallic ion concentration has been chosen to be below toxic level in blood plasma148,176,177. The sol-gel glass synthesis solutions were used to coat BG-based scaffolds, by dip coating and impregnation, in order to improve the mechanical properties of the combined system and to introduce biologically active ions which will be released overtime. The resulting MBG_BG scaffolds were characterized; the mechanical properties and the bioactivity of the combined system are presented in this chapter. The results presented in this chapter have formed out of previous publication64.

6.2. Materials and methods

6.2.1. Co-MBG Synthesis Co doped ordered mesoporous sol-gel glass was prepared by an ethanol based synthesis solution. The Co-MBG composition (in mol.%) was 78SiO2 20CaO 1.2P2O5 0.8CoO. In a typical synthesis procedure 4.5 g Pluronic® F127, the surfactant directing agent, were added to a solution of 85 mL ethanol (96%) and 1.2 mL nitric acid (HNO3, 1M) under continuous stirring. Once the F127 was completely dissolved and the solution was clear, the following reactants were added in this order with an interval of 3h between one to the other: 9.597 mL of tetraethyl orthosilicate (TEOS), 0.112 mL of triethyl phosphate (TEP) and 2.576 mL calcium nitrate tetrahydrated (CaNO3 4H2O). After the addiction of the calcium source, the solution was stirred for another hour and 0.128 g of cobalt hexahydrated (Co(NO3)2 6H2O) were added and the solution was stirred overnight (12 h). The solution was then transferred to petri dishes for the evaporation induces self-assembly (EISA) procedure and placed under the hood in a protected environment. After 5-7 days, depending on the humidity, the obtained material was thermally treated at 700 °C (5 °C/min) for 3 h.

6.2.2. SrCu-MBG Synthesis SrCu doped ordered mesoporous sol-gel glass was prepared by an ethanol based synthesis solution. The SrCu-MBG composition (in mol.%) was 78SiO2 20CaO 1.2P2O5 0.4CuO 0.4SrO. 4.5g Pluronic® F127, the surfactant directing agent, were added to a solution of 85 mL ethanol (96%) and 1.2 mL nitric acid (HNO3, 1M) under continuous stirring. Once the F127 was completely dissolved and the solution was clear, the following reactants were added in this order with an interval of 3h between one to the other: 9.597 mL of tetraethyl orthosilicate (TEOS), 0.112 mL of triethyl phosphate (TEP) and 2.576 mL calcium nitrate

124 tetrahydrated (CaNO3 4H2O). After the addiction of the calcium source, the solution was stirred for another hour and 1.2 mL of strontium nitrate (Sr(NO3) and 0.0503 g of cupper nitrate (Cu(NO3) 2.5H2O) were added with a 1 hour interval between each addition. The solution was then stirred overnight (12 h). The solution was then transferred to petri dishes for the evaporation induces self-assembly (EISA) procedure and placed under the hood in a protected environment. After 5-7 days, depending on the humidity, the obtained material was thermally treated at 700 °C (5 °C/min) for 3 h.

6.2.3. MBG coated BG-based Scaffold Preparation The BG-PU and BG-SA scaffolds were prepared as reported in Chapter 4, Section 4.2.1. The ion doped sol-gel glasses solutions were prepared as described in the previous section. Once the solutions were stirred overnight (12h) they were used to coat the BG scaffolds following two different methods, namely dip coating (DC) and impregnation (IMP). For the DC procedure, the BG scaffolds were completely immersed in the sol-gel glass synthesis solution for 5 min, removed and dried under the hood at RT for 12 h. The process was repeated up to 6 times in order to optimize the coating. For the IMP method, 500 µL of sol-gel glass synthesis solution was dropped on the BG scaffolds, the extra solution was removed with blotting paper and they were dried at RT under the hood for 12 h. The procedure was repeated up to 5 times in order to optimize the coating. In both cases, DC and IMP procedures, the scaffolds were heat treated at 700 °C (5 °C/min) for 3 h.

6.2.4. Bioactivity Test in SBF Simulated body fluid (SBF) was prepared by dissolving reagent grade 8.035 g L−1 NaCl, −1 -1 −1 −1 0.355 g L NaHCO3, 0.225 g L KCl, 0.231 g L K2HPO4 (3H2O), 0.311 g L MgCl2 −1 −1 (6H2O), 0.292 g L CaCl2, and 0.072 g L Na2SO4 in deionized water and buffered at pH −1 7.4 at 36.5°C with 6.118 g L tris(hydroxymethyl) aminomethane ((CH2OH)3CNH2) and 1M HCl, as previously reported by Kokubo and Takadama130.

Co_MBG and SrCu_MBG powders were immersed in SBF at a 1.5 g L−1 ratio37,152. The specimens were kept in a polypropylene container at 37°C in an incubator on an oscillating tray (90 rpm) for up to 3 days. The solution was not renewed and a falcon tube containing SBF as a control was also used for the entire period of the experiment, in order to control overtime the stability of the testing solution. At the end of the incubator period, the particles

125 were washed twice with deionized water, dried at RT for 7 days, and stored for further characterizations. The unchanged SBF was kept for the analysis of the Ca, P, Si, Co, Sr and Cu ions concentrations through ICP-OES and for pH analysis.

BG pellets, double heat treated at 400°C-1h (2°C/min), 1050°C-1h (2°C/min), cooled at RT, 700°C-6h (5°C/min), were immersed in TRIS buffered solution in order to evaluate if the double treatment affected the bioactivity of the starting BG. The pellets were tested for up to 14 days in a shaking incubator (90rpm) at 37°C.

Cylindrical BG-PU and BG-SA scaffolds coated with sol-gel glasses were tested in SBF for up to 28 days at 37°C in a shaking incubator (90 rpm). The solution was renewed every 2-3 days in order to better mimic the in vivo behavior, as carried out also in previous studies32. At the end of the incubator period, the foams were washed with deionized water, dried at RT for 3 days, and stored for further characterizations.

For all the experiments, a falcon tube containing SBF as a control was also incubated in order to control that no precipitation occurred in the testing solution. In this way in case HCA is formed, it is possible to correlate it directly to the bioactivity of the tested system, making sure that is not due to the precipitation of the SBF solution. The shape and the surface structure of the resulting MBG coated BG scaffolds were evaluated by means of scanning electron microscopy SEM coupled with EDS. The porous structure of the MBG glasses was assessed with high resolution transmission electron microscopy (HRTEM). The pore diameter analyses were conducted on HRTEM images with ImageJ analysis software. Nitrogen adsorption desorption analysis was conducted to assess the specific surface area and the pore size of the particles. The compression strength of the BG scaffolds coated with the MBG glasses was also evaluated by compression strength test. In order to better distinguish between the effect of the coating and the effect of the double thermal treatment (DTT) on the mechanical behaviour of the resulting scaffolds, uncoated BG scaffolds underwent the same thermal treatment of coated scaffolds. All the details on the characterization techniques, brands and samples preparation were reported in Chapter 3.

6.2.5. Drug Up-take and Release Capability of MBGs To load MBGs with a drug, Ibuprofen (> 98%, purchased from Sigma-Aldrich) as model drug was dissolved in hexane (33 mg mL−1) and MBG powders were added to the drug solution (33 mg mL−1) at RT following the procedure reported in a previous work. Three different up-take times were used, i.e. 6h, 12h and 24 h, to optimize the up-take. The drug

126 solution was then removed and the MBGs were dried in a vacuum hood at RT. The drug- release kinetics (10 mg each sample) was assessed by soaking the samples in 4 mL of PBS, kept at 37°C in a shaking incubator (90 rpm) until the complete release of ibuprofen. At every time point, 1 mL of solution was uptake for the drug release analysis and substituted with 1 mL of fresh PBS. A UV-vis spectrophotometer was used to evaluate the amount of released drug. The calibration curve was calculated using a solution of ibuprofen in PBS with different known concentrations, on the basis of the absorption at 273 nm, typical of this molecule.

6.2.6. Capillarity Test The permeability of the scaffolds was qualitatively checked through capillarity tests using a red ink solution to better observe the fluid infiltration within the scaffold. A face of the scaffold was put into contact with the fluid and the infiltration driven by capillarity forces was observed178.

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6.3. Results

6.3.1. Co-MBG and SrCu-MBG: Preparation and Characterization The resulting SrCu-MBG and Co-MBG glasses, after the thermal treatment, were characterized by light blue and violet colours respectively due to the presence of the ions. Optical images of Co-MBG and SrCu-MBG before and after EISA process and thermal treatment at 700 °C for 6 h are reported in Figure 6.1.

Figure 6.1 Optical images of Co-MBG and SrCu-MBG before EISA (a,b), after EISA (c,d) and after the thermal treatment at 700°C for 6 h (e,f).

The resulting glasses exhibited high ordered mesoporosity, as shown in HRTEM images in Figure 6.2. The plots of the distance between the pore channels obtained by the analysis of the TEM images with ImageJ plug in Plot Profile were reported in Figure 6.2c. The plot profile was quite regular for both glasses and an average pore size of 7.5 nm was found. The glasses were also characterized by high specific surface area, 254 m2 g-1 and 346 m2 g-1, respectively, for SrCu-MBG and Co-MBG samples.

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Figure 6.2 HRTEM images of SrCu-MBG (a) and Co-MBG (b) glasses; plot of the distance between the parallel pore channels obtained with ImageJ plug in Plot Profile.

The resulting materials, after the thermal treatment, were still completely amorphous as it was possible to determine by XRD analysis shown in Figure 6.3. The amorphous background associated to amorphous silica at 2θ 15-30 degrees was visible on the spectrum.

Figure 6.3 XRD analysis of SrCu-MBGs and Co-MBGs after the thermal treatment.

i. Bioactivity Evaluation The two glasses were tested in SBF for up to 7 days in order to evaluate their bioactivity. The SEM micrographs of the samples before and after the immersion in SBF for 1 h are shown in Figure 6.4a-d. Already after 1 h of immersion in SBF the surface of both glass

129 powders was covered with a deposit (Figure 6.4 c,d) which evolved over time in fully- developed HCA, as shown in Figure 5.

Figure 6.4 SEM images of Co-MBG and SrCu-MBG before (a,b) and after the immersion in SBF for 1 h (c,d).

Figure 6.5 SEM micrographs of Co-MBG and SrCu-MBG powders immersed in SBF for 1d (a,b) and 3d (c,d).

The high bioactivity of these glasses was also confirmed by ICP analysis, evaluating the ions release form the Co-MBGs and SrCu-MBGs over time in SBF (Figure 6.6). Immediately after the immersion of the samples in SBF, a clear changed in the SBF ion concentration was observed, in agreement with the steps reported for the HCA formation.

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Figure 6.6 ICP analysis of Si, P, Ca, Co, Cu and Sr ions release for up to 7 days in SBF.

ii. Drug Test Three different up-take times were considered, 6, 12 and 24h, in order to evaluate the optimal drug up-take condition. The drug release profile is shown if Figure 6.7. For both Co- MBG and SrCu-MBG 12 and 24h showed the best drug up-take results, in fact a higher amount of drug was loaded inside the glass pores. Not a significant difference was observed between 12 and 24 h. Co-MBG loaded a significant higher amount of drug compare to SrCu- MBG. For both materials a burst release is observed during the first hours of test. During this time almost the 50% of the total loaded drug was released. The remaining 50% of drug was released within the next 7 days, after this time point a plateau is reached for all tested samples.

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Figure 6.7 Drug release profile of Co-MBG (a) and SrCu-MBG loaded for with drug using three different up- take times, i.e. 6, 12 and 24 h.

6.3.2. Double Thermal Treated BG Pellets: Bioactivity Evaluation A series of BG pellets was double thermal treated in order to evaluate if their bioactivity was affected by such treatment. Firstly, the pellets were heat treated up to 1050 °C for 1h, as for the preparation of BG-based scaffolds (Chapter 4, Section 4.2.1), next at 700 °C for 3 h as for the preparation of MBGs (Chapter 6, Sections 6.2.1 and 6.2.2). They were immersed in TRIS buffered solution for 14 days and the surface of the pellets was characterized by SEM analysis (Figure 6.8). They were compared to BG pellets that underwent the normal thermal cycle for the sintering of BG scaffolds, as reported in Chapter 4, Section 4.2.1. The double thermal treated BG pellets were bioactive and a HCA deposit was formed on their surface after 14 days of test (Figure 6.8d). The results obtained for the double thermal treated samples are comparable with results on BG pellets heat treated only at the sintering temperature of the BG scaffolds.

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Figure 6.8 SEM micrographs of BG pellets heat treated at 1050 °C, similarly to the sintering of the scaffold before (a) and after (c) the immersion in TRIS buffered solution for 14 days, BG pellets double thermal treated, first at 1050 °C and after at 700 °C before (b) and after (d) immersion in TRIS buffered solution; SEM-EDS analysis of the BG pellets after 14 days in SBF (e,f).

6.3.3. MBG Coating by Dip Coating (DC) Method

i. Coating Optimization The BG-PU scaffolds were dip coated in Co-MBG synthesis solution, increasing the number of coatings in order to optimize the process. From digital camera images reported in Figure 6.9, it was possible to observe that increasing the number of coating the blue color was more intense due to the increase of the coating thickness.

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Figure 6.9 Digital camera images of BG-PU scaffolds coated up to 6 times by dip coating procedure with Co- MBG synthesis solution after the thermal treatment at 700 °C for 6 h.

SEM images of obtained Co-MBG BG-PU are shown in Figure 6.10. After the second coating, the surface of the scaffold was almost completely covered with the sol-gel glass, but the pattern of the BG-PU scaffolds surface was still visible. Increasing the number of coatings, also the thickness of the sol-gel glass coating increased but with 5 repetitions an overflow of material was observed (Figure 6.10d).

Figure 6.10 SEM micrographs of BG BG-PU scaffolds dip coated in sol-gel glass synthesis solution increasing the number of coating cycles, from 2 up to 5.

ii. Mechanical Properties Evaluation Compression tests were performed on DTT BG-PU scaffolds and on the BG-PU foams coated with Co-MBG. The best compression strength values, ~ 0.1 MPa were obtained for the BG-PU scaffolds dip coated 4 times with the sol-gel glass synthesis solution, which is 4 times higher than values obtained for the DTT non-coated BG-PU scaffolds. Increasing the number of coating up to 5 or 6 did not enhance more the resulting mechanical properties of the coated scaffolds. In fact, the maximum compressive strength values obtained are similar to the Co-MBG BG-PU scaffolds dip coated 3 times. All the maximum compressive strengths obtained are summarized in Table 6.1.

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Table 6.1 Maximum compressive strength values of BG-PU scaffolds coated with Co-MBG increasing the dip coating cycles up to 6.

Maximum compressive Strength [MPa] BG-PU No Coating 2X 3X 4X 5X 6X scaffolds < 0.05 0.07 ± 0.02 0.10 ± 0.03 0.12 ± 0.03 0.10 ± 0.03 0.08 ± 0.03

iii. Bioactivity Test The scaffolds dip coated for times with Co-MBG were tested in SBF in order to assess the bioactivity of the system. Contrary to the expectation, the bioactivity of the Co-MBG BG- PU was very slow, how it was possible to observe form SEM images reported in Figure 6.11. After 1 day of immersion, the surface of the scaffold has some deposit (Figure 6.11a), but the development of this deposit was slow and after 7 days it was still no able to cover the entire surface of the scaffold (Figure 6.11b).

Figure 6.11 SEM micrographs of Co-MBG BG-PU dip coated 4 times with sol-gel glass solution and immersed in SBF for 1 day (a) and 7 days (b).

6.3.4. MBG coating by Impregnation (IMP) Method

i. Coating Optimization The BG scaffolds were coated with Co-MBG and SrCu-MBG synthesis solutions an increasing number of cycles, in order to optimize the coating procedure. In Figure 6.12 are reported the weight variation of the BG-SA scaffolds after the coating with the sol-gel glasses synthesis solution. The weight was increasing linearly, increasing the number of coating cycles.

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Figure 6.12 Weight variations of BG-SA scaffolds impregnated up to 5 times with Co-MBG and SrCu-MBG synthesis solutions.

In figure 6.13, SEM images of the BG-PU scaffolds impregnated 1, 2 and 3 times with Co- MBG sol-gel glass synthesis solution are shown. It is possible to observe that, already after 2 impregnations an overflow of coating material has occurred (Figure 6.13b) and after 3 coatings the sol-gel glass blocked almost completely the open porosity of the starting BG scaffold (Figure 6.13c). A single impregnation was considered as the optimized condition, because the surface of the BG scaffold was homogeneously coated and no overflow of material was observed.

Figure 6.13 SEM micrographs of BG-PU scaffolds impregnated 1, 2 and 3 times with Co-MBG sol-gel glass synthesis solution.

Independently of the type of sol-gel glass solution used for the coating or the type of BG scaffold impregnated, the single impregnation was shown to cover every single part of the scaffolds surfaces without blocking the pores of the samples. In addition, no overflow of MBG was observed from the SEM images reported in Figure 6.14.

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Figure 6.14 SEM micrographs of Co-MBG BG-SA (a), SrCu- MBG BG-SA (b), Co-MBG BG-PU (c) and SrCu- MBG BG-PU (d) scaffolds after one impregnation with sol-gel glass synthesis solutions.

The MBG coating was seen to cover every single part of the BG scaffold surface forming a very thin layer (Figure 6.15a). It was not possible to determine the thickness of this layer from cross sections of the scaffolds (Figure 6.15e) most probably due to the very similar nature of the scaffold and coating materials. The sol-gel glass synthesis solutions, which were ethanol based, were not viscous but liquid and for this reason it was easy to impregnate and cover every single part of the scaffold. In some areas, the coating was seen to cover completely the structure of the sintered BG (Figure 6.15b), in some parts the BG structure was still visible and the MBG coating was a thin layer following every change of its morphology (Figure 6.15d). In Figure 6.15c an SEM image of a BG-SA scaffold is reported. In this image the main phases composing the sintered BG are visible, which were described in Chapter 4, Section 4.3.1. In Figure 6.15d, it is still possible to recognize these phases, but the space between them has been filled with sol-gel glass.

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Figure 6.15 SEM micrographs of BG-SA scaffolds impregnated with Co-MBG sol-gel glass synthesis solution at different magnifications (a, b and d) compared to the non-coated BG-SA scaffold (c). SEM image of a cross section of a Co-MBG BG-SA scaffold strut (e).

ii. Mechanical Properties Evaluation The effect on the mechanical properties of the MBG coating was assessed my compression test. The maximum compressive strength of the BG-PU scaffolds increased significantly thanks to the coating with the sol-gel glasses. In case of BG-PU scaffolds coated with Co- MBG and SrCu-MBG, the maxim compressive strength increased up to ~ 0.1 MPa. In case of BG-SA scaffolds, the coating with sol-gel glasses did not have any effect on the mechanical properties of the samples.

iii. Bioactivity Test of MBG coated BG Scaffolds in SBF BG-SA scaffolds coated with Co-MBG and SrCu-MBG were tested in SBF for up to 28 days. After the immersion in SBF, a decrease of the weight (Figure 6.16a) and an increase of

138 the total porosity (Figure 6.16b) were observed. The samples lost around 15% of their starting weight during the first 7 days of immersion and in the following 3 weeks the loss increased only by 5%. The porosity of the samples increased by 10% in 2 weeks, but at the end of the test the total porosity started to decrease. The behavior of Co-MBG BG-SA and SrCu-MBG BG-SA scaffolds was comparable. For the BG-PU scaffolds coated with MBGs, it was not possible to analyze the weight and porosity variation because the samples were almost completely degraded due to their high dissolution in SBF and fragility with repeat handling.

Figure 6.16 Co-MBG BG-SA and SrCu-MBG BG-SA scaffolds weight (a) and porosity (b) variation after immersion in SBF for up to 28 days.

The stability of the mechanical properties of the prepared scaffolds after immersion in SBF was evaluated. Also in this case, the values reported referred only to the natural sponge BG scaffolds, because the MBG BG-PU scaffolds were too fragile to be handled and it was not possible to perform the compression strength test on them. In case of MBG BG-SA scaffolds, also after 28 days in SBF, it was still possible to handle them without damaging. The Co-MBG BG-SA and SrCu-MBG BG-SA samples exhibited different results in term of maximum compressive strength (Figure 6.17). The BG-SA scaffolds coated with Co-MBG showed a loss of almost the 65% of the starting maximum compressive strength during the first 7 days of immersion in SBF (Figure 6.17a). After 14 days, the 30% of the lost strength was seen to recover. At the end of the test at 28 days, the variation in the tested samples was very high. Some samples showed a maximum compressive strength higher than the before the immersion in SBF, other samples showed a maximum compressive strength similar or lower than the one obtained at 14 days. The average values of maximum compressive strength obtained for the Co-MBG BG-SA scaffolds at the end of the immersion in SBF was 2.4 ± 0.8 MPa. Also for SrCu-MBG BG-SA scaffolds during the first week in SBF, the maximum compressive strength decreased of almost the 65% from the starting value (Figure 6.17b) but already after two weeks it was not possible to identify a specific trend for the

139 mechanical properties. In fact, the maximum compressive strength values were characterized by an extremely high standard deviation. At the end of the test, the maximum compressive strength of SrCu-MBG BG-SA scaffolds was 1.58 ± 0.42 MPa.

Figure 6.17 Maximum compression strength variations of Co-MBG and SrCu-MBG BG-SA scaffolds after immersion in SBF for up to 28 days. The SBF was refreshed every 2-3 days.

The pH variation and the changes of ion concentrations in SBF during the first week of immersion of Co-MBG and SrCu-MBG BG-SA scaffolds are reported in Figure 6.18. Positive values of ion release imply an increase of the ion concentration in SBF, while negative values refer to a depletion of the ion concentration in the testing solution due to the immersion of the scaffolds. Scaffolds dissolution took place immediately after the first hours of immersion in SBF, indicated by the pH trends shown in Figure 6.18d. The pH of the scaffolds coated with SrCu-MBG was characterized by a faster increase, up to a maximum value of 7.9 which was already reached after 3 days of test. In both cases, more clearly for SrCu-MBG BG-SA scaffolds, pH equilibrium was almost reached. The scaffolds coated with Co-MBG exhibited a slower pH increase. After 3 days, the pH reached a value of 7.7 and only after 7 days it went up at 7.85. After 7 days of immersion in SBF, around 250 mg L-1 of Si species was released from both Co-MBG and SrCu-MBG BG-SA samples (Figure 6.18a) but the release profile did not increase linearly. The Ca ion concentration increased significantly up to maximum value of 400 mg L-1 at the end of the test (Figure 6.18c). The P concentration decreased overtime (Figure 6.18b), indicating that P-containing phase was deposited on the surface of the scaffolds during immersion in SBF. It was not possible to detect the functional ions, i.e. Co, Sr and Cu, by ICP analysis because the MBG layer was extremely thin and the concentrations of the released ions was too low to be measured.

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Figure 5.18 ICP Concentration of (a) Si, (b) P, (c) Ca ions released from Co-MBG BG-SA and SrCu-MBG BG- SA scaffolds up to 7 days in SBF, without SBF refresh every 2-3 days; pH variation in 7 days of immersion in SBF, without SBF refresh.aa

SEM micrographs of the BG-PU and BG-SA scaffolds coated with Co-MBG and SrCu- MBG after the immersion in SBF for 1d are shown in Figure 6.19. It is possible to observe that, BG-SA scaffolds surfaces are characterized by a degradation of the coating layer on their external surface, which reveals the BG structure underneath it (Figure 6.19a,b). BG-PU scaffolds showed not only degradation of the coating, but also the formation of a deposit on the surface (Figure 6.19c,d).

aa ICP-OES analyses were done at Institute of Particle Technology in Erlangen by Dr. Jochen Schmidt

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Figure 5.19 SEM micrographs of the external surfaces of Co-MBG BG-SA (a), SrCu-MBG BG-SA (b), Co- MBG BG-PU (c) and SrCu-MBG BG-PU (d) scaffolds after immersion in SBF for 1 day.

The situation of the inner core of the scaffolds was completely different compared to the external side. In fact already after 1 day of immersion in SBF, the inner core of all tested samples was covered with a layer of well-developed HCA (Figure 6.20). The layer of HCA deposit was observed to cover all inner parts of the scaffolds.

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Figure 6.20 SEM micrographs of the inner core of Co-MBG BG-SA (a), SrCu-MBG BG-SA (b), Co-MBG BG- PU (c) and SrCu-MBG BG-PU (d) scaffolds after the immersion in SBF for 1 day.

Also after 3 days in SBF, it was not possible to observe the formation of a homogeneous coating of HCA on the external surface of the samples (Figure 6.21b). In fact, some parts of of it were still characterized by a progressive degradation of the MBG coating (Figure 6.21c) and other parts were coated with well-developed HCA (Figure 6.21d,e). In figure 6.21, a comparison of the MBG coating before (Figure 6.21a) and after the immersion in SBF (Figure 6.21c) was possible. The MBG glass started to degrade where the thickness of the coating was lower, revealing the BG surface which was decorated with a Ca-P deposit.

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Figure 6.21 SEM micrographs of the external surface of Co-MBG BG-SA scaffolds before the immersion in SBF (a) and after the immersion for 3 days at different magnifications (b-dc) and high magnification images of the HCA layer formed (e).

Only after 7 days of immersion in SBF, the scaffolds external surfaces were covered with a continuous layer of HCA (Figure 6.22). SEM images of the strut cross sections of BG-SA (Figure 6.23) and BG-PU (Figure 6.24) scaffolds coated with Co-MBG and SrCu-MBG before and after the immersion in SBF are shown.

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Figure 6.22 SEM micrographs of Co-MBG BG-SA (a), SrCu-MBG BG-SA, Co-MBG BG-PU (c) and (b) and SrCu-MBG BG-PU (d) scaffolds after immersion in SBF for 7 days.

As already seen before, after 1 day in SBF the inner struts were completely covered with HCA layer which evolved overtime in thickness (Figure 6.23). The HCA layer was seen to grow on the external surface of the BG scaffold struts and on the wall of the hollow part of the struts (Figure 6.24). Progressively, the diameter of the holes decreased because of the continuous growing of HCA. Due to the HCA deposition, the opposite phenomenon occurred on the diameter of the strut, which increased progressively. In addition, increasing the time of immersion in SBF, the remaining BG was seen to become more porous.

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Figure 6.23 SEM micrographs of inner strut cross sections of Co-MBG BG-SA scaffolds (a-e) and SrCu-MBG BG-SA scaffolds (f-l) before and after immersion in SBF for 1, 7, 14 and 28 days.

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Figure 6.24 SEM micrographs of inner strut cross sections of Co-MBG BG-PU scaffolds (a-d) and SrCu-MBG BG-PU scaffolds (e-h) after immersion in SBF for 1, 7, 14 and 28 days.

Figure 6.25 summarizes FTIR spectra of the samples before and after immersion in SBF for up to 28 days. In the case of the coated scaffolds before immersion, the FTIR spectra at room temperature presented the characteristic peaks of the non-bridging oxygen stretching mode of Si–O located at 950 cm-1. After 7 days, characteristic peaks ascribed to the HCA layer formation appeared as a doublet at around 600 cm-1, corresponding to the bending mode of P–O (crystalline phosphate). Moreover, the observation of P–O stretching at around 1000

147 cm-1, where the band became narrow, suggested the presence of HCA. The spectra also showed the narrowing of the band at around 800 cm-1 corresponding to the bending mode of C–O and finally the manifestation of the stretching mode of C–O at around 1400 cm-1.

Figure 6.25 FTIR spectra of (a) Co-MBG BGPU, (b) SrCu-MBG BG-PU, (c) Co-MBG BG-SA and (d) SrCu- MBG BG-SA scaffolds before and after the immersion in SBF for up to 28 days. The relevant peaks are discussed in the text.

iv. Capillarity Test

The permeability of the scaffolds was qualitatively checked out through capillarity tests using a red solution to better observe the fluid infiltration within the scaffold (Figure 6.26). A face of the scaffold was put into contact with the fluid and the infiltration driven by the capillarity force was imediately abserved for all the BG-based scaffolds, non-coated and coated with MBGs. This confirm the interconnected porosity of BG-PU and BG-SA scaffolds and that the MBG coatings do not affect this inteconnectivity.

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Figure 6.26 Optical images of BG-SA scaffolds (a) non-coated (b) coated with Co-MBG and (c) coated with SrCu-MBG synthesis solutions; BG-PU scaffolds (d) non-coated, (e) coated with Co-MBG and (f) coated with SrCu-MBG synthesis solutions.

6.4. Discussion Ordered mesoporous sol-gel glasses, developed for the first time in 2004 by Yan et al.98, are nanostructured materials with a chemical composition similar to conventional sol-gel glasses and they can be doped with functional ions with different therapeutic purposes85,99,179. 3D bioactive glass based scaffolds based on MBG have been developed but the resulting mechanical properties were rather low62,171.

In the present work, an innovative coating method of 45S5 Bioglass®-based scaffolds with Co-MBG and SrCu-MBG was proposed. The idea was to combine the BG scaffolds with improved mechanical properties developed using natural sponges as sacrificial template (Chapter 4)60 with MBG glasses, which can incorporate functional molecules in their ordered structure and release functional ions overtime100–103,111,112,179,180.

The two ion doped sol-gel glasses, Co-MBG and SrCu-MBG, used for the functional coating of the BG scaffolds have been developed by Philippart173. The range of substituent´s concentrations was chosen to not inhibit or decrease the bioactivity of the starting glass composition174,175. The metallic ion concentration has been chosen to be below toxic level in blood plasma148,176,177. Co2+ and Cu2+ ions are both known for their capability to enhance angiogenesis and osteogenesis162,163,181. Wu et al.181 demonstrated that these two ions are able to induce an hypoxia-mimicking environment163,181. Hypoxia is a condition of low oxygen pressure and it plays a key role in coupling angiogenesis with osteogenesis. Hypoxia

149 activates a series of angiogenic processes mediated by the hypoxia inducing factor-1a (HIF- 1α, transcription factor)163. This HIF-1α molecule initiates the expression of different genes associated with tissue regeneration, skeletal tissue development and enhances fracture repair. Co2+ and Cu2+ ions can chemically induce HIF-1α to promote hypoxia-like response. Cu2+ ions can also significantly inhibit bacterial viability. Sr2+ ions, as other divalent ions, e.g. Zn2+ and Mg2+, can stimulate the osteogenic/cementogenic differentiation of human bone marrow stromal cells164,166–168,182,183. Zhang et al.166 showed that Sr2+ can stimulate new bone formation in osteoporotic fractures.

Both ion doped sol-gel glasses used, Co-MBG and SrCu-MBG, were characterized by ordered mesoporosity and high specific surface area and a pore size of ~ 7.5 nm. Both materials showed high bioactivity and after 1 h of immersion in SBF they were seen to develop HCA on their surfaces. The high bioactivity of these materials, even higher than that of standard sol-gel glasses, is related to the high specific surface area, which characterizes them. In fact, the surface in direct contact with the SBF is very high and a high number of silianol groups is available and can act as nucleation sites for HCA. Their fast reactivity in SBF was confirmed not only by SEM analysis, but also by the ICP results. After few hours of immersion in SBF, the glasses released relatively high quantity of Si and Ca ions and the P concentration in the SBF decreased overtime confirming that a Ca-P phase deposition on the scaffold surfaces. These glasses were used to prepare functional coating on BG scaffolds. The sol-gel glass synthesis solution was used to dip coat or to impregnate the samples and only after the coating procedure the thermal treatment at 700 °C was carried out to remove the nitrates and surfactant directing agent. This means, that the BG scaffolds underwent a double thermal treatment, i.e. a first thermal treatment at 1050 °C for 1 h during the sintering of the scaffolds and a second thermal treatment at 700 °C for 6 h after the coating with the MBG glasses. At 610 °C, 45S5 Bioglass® undergoes crystallization of the primary crystalline phase 141. Most probably, this means that the percentage of crystalline phase of the BG would increase and the amorphous phase would be most probably reduced as result of the second thermal treatment. For this reason, it was tested if the BG was still bioactive after this double thermal treatment. The worst case of bulk samples was considered, testing BG pellets in TRIS buffered solution. After 14 days of test the double thermal treated pellet showed still bioactivity and it was comparable to that of BG pellet heat treated only once (at the sintering temperature). This was a very important to confirm experimentally that the resulting BG would have not been affected by the thermal treatment and that the approach to coat the scaffolds with MBG was feasible.

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The first method developed for coating BG scaffolds with sol-gel glasses was the dip coating procedure, which was optimized on the BG-PU scaffolds. Different dip coating cycles were tried and it was found that 4 coatings gave the best results in terms of quality of the coating and mechanical properties. Repeating 4 times the coating procedure allowed obtaining a homogeneous coating on the entire surface of the scaffolds without an overflow of material. Moreover the mechanical properties were enhanced due to the functional coating. The maximum compressive strength achieved was ~ 0.1 MPa, which is 4 times higher compared to that of non-coated scaffolds heat-treated two times. The system looked very promising, but the bioactivity was unexpectedly low. After 1 day of immersion in SBF only few spots of HCA were formed and also after 7 days most of the surface of the scaffold was not coated with HCA. The BG scaffolds were left in the sol-gel glass synthesis solution for 5 min every time the dip coating procedure was repeated. The solution, due to the strong acid condition used to catalyze the reaction clearly affected the BG.

For all these reasons, the impregnation method was implemented. The base idea was to coat the BG scaffolds by a rapid procedure leading to the shortest interaction time between the BG scaffolds and the sol-gel glass synthesis solution. Also in this case, the coating procedure was optimized on the BG-PU scaffolds. It was possible to observe that with one single impregnation all the surfaces of the scaffold were covered with the sol-gel glass and the same was the case for the inner core of the samples. Increasing the number of impregnation cycles, after 2 repetitions there was an overflow of sol-gel glass. A single impregnation was considered sufficient and also BG-SA scaffolds were coated with SrCu-MBG and Co-MBG. The coating was very homogeneous and it covered efficiently the scaffold structure. Due to the similarity of materials used for the preparation of the BG scaffolds and for the coating, it was not possible to determine the thickness of the MBG layer by SEM observations. Also in this case, the mechanical properties of the resulting Co-MBG and SrCu-MBG BG-PU scaffolds increased thanks to the functional coating with the sol-gel glass. Independently from the type of sol-gel glass solution used, the maximum compressive strength was ~0.1 MPa. In case of BG-SA scaffolds, no improvement of the mechanical properties was observed after the coating with the sol-gel glasses. This is most probably related to the fact that the layer of MBG glass was too thin to affect the mechanical properties of scaffolds, which are already characterized by relatively high compressive strength.

The scaffolds impregnated with Co-MBG and SrCu-MBG were also tested in SBF and the composite system resulted highly bioactive. Immediately after the immersion in SBF, both

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Co-MBG BG-SA and SrCu-MBG BG-SA samples were characterized by weight loss and increased porosity. During a single week in SBF, the samples lost 15% of their starting weight. The weight loss stabilized and between 14 days and 28 days no further weight decrease was observed. The porosity of the samples increased over time up to a value of 10%. After 14 days, the porosity variation reversed the tendency and it start to decrease. The weight loss and the porosity variation are two direct consequences of the immersion in SBF of the BG scaffolds. In fact, immediately after the immersion the scaffolds start to degrade inducing a decrease of the weight and an increase of the total porosity. The effect of the degradation was much less severe for the MBG BG scaffolds compared to the non-coated one that in the same testing condition lost almost the 50% of the starting weight. This could be explained in two ways. The MBG BG scaffolds were heat treated two times, once at 1050 °C and the second time at 700 °C inducing an increase of the crystalline phase in the BG. An increase of the crystalline phase does not induce a complete absence of bioactivity, but just an increase of the necessary time in SBF to form HCA. This explanation is not the most presumable, because the MBG BG scaffolds showed high bioactivity immediately after 1 h in SBF. A second explanation is the increased bioactivity of the system, due to the presence of the MBG glasses on the surface of the BG scaffolds. Once the scaffolds are immersed in SBF two phenomena occur, i.e. the degradation of the glass and the precipitation of HCA on their surface. Initially the degradation of the glass is the main phenomenon, after there is equilibrium between degradation of BG and precipitation HCA and at the end the deposition of HCA is predominant. Due to the high specific surface area and the high number of silanol groups available on the MBG surfaces, the two phenomena of degradation and deposition can reach faster the equilibrium and the HCA deposition start to become the main phenomenon before 28 days. The SEM images confirmed that already after 1day of immersion in SBF, the inner core of the scaffolds were completely covered with full- developed HCA. From the porosity variation plot, it is possible to suggest that this change in the equilibrium between desorption and degradation took place after 2 weeks, but the mechanical properties results of Co-MBG BG-SA scaffolds indicate that this can occur already after 1 week in SBF. In fact, the mechanical property of Co-MBG BG-SA samples decreased by 65% during the first week of immersion in SBF, due to the progressive BG strut density decrease caused by the degradation of the BG. After 14 days, the mechanical properties started to increase and the loss of mechanical stress was reduced to 30%. This effect can be related to an increase of the HCA deposition not only on the external surface of the scaffolds, but also in the central hole of the struts. It is not possible to discuss the behaviour after 28 days, because the standard deviations of the measured values were too high. In case of SrCu-MBG BG-SA samples, the maximum compressive strength decreased

152 also to 65% during the first 7 days, but the values measured at 14 and 28 days cannot be used for further consideration. The test should be repeated, increasing the number of scaffolds used in order to assess the scaffold´s behaviours more accurately.

6.5. Conclusions Ordered mesoporous ion doped sol-gel glasses were successfully used to develop functional coatings of BG scaffolds. The Co-MBG and SrCu-MBG were characterized by high specific surface area and enhanced bioactivity compared to standard sol-gel glass or melt derived glasses. The presence of the MBG glass coating on the surface of BG-PU samples, obtained by repeating dip coatings, did not yield the expected results. In contrast, MBG BG-PU scaffolds obtained by impregnation were characterized by improved mechanical properties and bioactivity. MBG BG-SA samples exhibited improved bioactivity but no effects were observed on the mechanical properties. In conclusion, it was possible to obtain four different functional systems with improved bioactivity, which have also the intrinsic capability to release functional ions and drugs.

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7. Conclusions and Future Directions 7.1. Conclusions The present work aimed to synthetize 45S5 Bioglass®-based porous scaffolds by the standard replica method15 but utilizing natural marine sponges as sacrificial templates to achieve superior mechanical properties. In addition, increasing the functionality of the obtained scaffolds was investigated incorporating silicate coatings with ordered mesoporous structure obtained by sol-gel technique. The resulting systems have the capability of controlled bioactive ion and drug release to enhance bone formation, neo-angiogenesis and to prevent bacterial infection.

Within this work a novel family of 45S5 Bioglass®-based scaffolds was developed and characterized. The foams were manufactured via replica method, a well-established production process for BG scaffolds15. Instead of the standard polyurethane packaging foams, which were used in the first works for making Bioglass®-based scaffolds, as templates15,124 natural marine sponges were used as templates. These sponges are characterized by a high interconnected porous structure, the result of their evolution for 1000 of years in water filtration127. The resulting scaffolds were characterized by a reduced porosity (68%) compared to the standard BG-PU samples (93%) but the pore interconnectivity was still higher than 99%. As direct consequence of the reduced porosity, in comparison to scaffolds made from PU foams, a significant increase of the mechanical properties was observed. The BG-PU scaffolds were characterized by a maximum compressive strength < 0.05 MPa (although values of 0.4 MPa have been reported15) and BG-SA scaffolds reached 4.0 ± 0.4 MPa. Similar results in terms of mechanical performance for 45S5 Bioglass® scaffolds with 68% porosity were obtained, so far, using a powder metallurgy approach developed in 201070. Thus the scaffolds developed in this project, showed good mechanical strength comparable with the compressive strength of cancellous bone (2-12 MPa)72.

Due to the high interconnected porosity of scaffolds obtained with natural marine sponges, the bioactivity of the resulting foams was comparable to that of BG-PU samples. Moreover, BG-SA scaffolds showed suitable mechanical properties also after immersion in simulated body fluid (1.2 ± 0.2 MPa). After an initial decrease of the maximum compressive strength due to the degradation of the BG material, a stabilization of the samples mechanical strength was observed after 4 weeks in SBF. This result is due to the deposition of carbonate hydroxyapatite (HCA) on the surface of the scaffolds and in the internal hollow structure of

155 the sample struts. The deposition of the HCA layer counteracts the decrease of the mechanical properties associated with the degradation of the BG and contributes to the increase of stiffness of the samples, which is also formed by its higher density compared to that of BG. Moreover HCA crystals firstly formed on the scaffold surfaces where micro cracks were located, because in such areas the pH increase was faster, thus enhancing the HCA formation37,184. In this way, HCA crystals filled the cracks in the struts (Chapter 4, Figure 4.17); it is well-known that such micro-cracks are one of the main causes for the relatively low mechanical properties of BG scaffolds6,139.

A relative drawback of scaffolds synthetized with natural marine sponges, due to the lower total porosity, was a lower oxygen diffusivity compared to BG-PU scaffolds. In fact, results of µ-Ct analysis (Chapter 4, Section 4.3.5) of both types of scaffolds suggested a decrease of 39% of the corresponding bulk diffusivity of BG-SA foams compared to BG-PU samples. The difference between BG-SA and BG-PU scaffolds in terms of oxygen diffusivity decreased with increasing time of immersion in SBF due to the progressive degradation of BG which leads to an increase of the total porosity of the samples. It is important to underline that the simulation analysis (Chapter 4, Section4.3.5) was conducted considering the worst case scenario for the oxygen diffusivity, i.e. the newly regenerated tissue completely occupying the porosity and no diffusion possible through the solid phase (BG material). The potential effect of this reduced oxygen diffusivity on the angiogenic potential of the scaffolds remains to be investigated54.

A preliminary cell culture study confirmed that no toxic residues of the natural marine sponges were left during the production of the BG scaffolds. The cell culture study demonstrated that cells can grow on the BG-SA samples and the cell viability was higher than on BG-PU samples. This can be explained once again by the lower porosity of the BG- SA foams, i.e. the local pH increase due to BG dissolution products was lower than for BG- PU samples (Chapter 4, Section 4.3.5).

In order to further enhance the mechanical properties of the Bioglass®-based scaffolds made from natural marine sponges and to introduce the capability to release biologically active ions and drug molecules, different functional coatings with ordered mesoporous materials have been developed. The goal was to combine the bioactivity of the Bioglass®-based scaffolds with the drug uptake and release capability of ordered mesoporous silica based materials.

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The first attempt followed a previous study of Mortera et al.63, which involved coating the BG-based scaffolds with ordered mesoporous silica particles, MCM-41. MCM-41 particles were synthetized by a novel synthesis solution. The resulting particles were characterized by high specific surface area, an ordered mesoporous structure and by a novel feature, namely an extremely porous surface (10-30 nm). As expected from literature154, this material was not bioactive and after 28 days in SBF no HCA was formed. The obtained particles showed a better stability of the ordered mesostructure after immersion in SBF, compared to reference MCM-41 particles produced following one of the well-known synthesis solution reported in literature90. The stability of the ordered mesostructure is a key requirement to have a functional and reliable drug carrier system154. The BG scaffolds were coated with MCM-41 particles following the procedure reported by Mortera et al.63 but also developing a post coating method. As expected, the coating of BG-PU and BG-SA scaffolds with MCM-41 particles enhanced the mechanical properties of the resulting combined system, notably up to 12 MPa for MCM-41_BG-SA foams. This was possible because calcination of the particles at 550 °C for 6 h induced softening of the BG which facilitated the adhesion of MCM-41 particles to the BG struts. In this way, the particles in direct contact with the BG surface could have filled the cracks present in the scaffolds struts. Although an increase of the mechanical properties was expected, the achieved enhancement of the bioactivity of the composite system compared to the non-coated scaffolds had not been anticipated. As mentioned above, the produced MCM-41 particles did not show bioactivity in SBF, in agreement with other works available in literature154. The behaviour of MCM-41 particles in SBF changed completely when they were combined with BG. In fact one of the features of the ordered mesoporous silica materials, as MCM-41 particles, is the presence of abundant silanol groups. For this reason MCM-41 particles can act as nucleation sites for HCA formation and the source of Ca and P ions is the underlying BG struts, which, upon dissolution, increase the pH allowing the precipitation of HCA26,184.

A second strategy followed to increase the functionality of BG-based scaffolds was to develop a functional coating applying ordered mesoporous sol-gel glasses doped with functional ions. Two glass compositions were used, i.e. 78SiO2-20CaO-1.2P2O5-0.4CuO- 2+ 2+ 0.4SrO (mol.%) and 78SiO2-20CaO-1.2P2O5-0.8CoO (mol.%). Co and Cu are both known for their capability to enhance angiogenesis due to their hypoxia-mimicking effect113. Moreover, Cu2+ ions can also significantly inhibit bacterial viability. Sr2+ ions can stimulate the osteogenic/cementogenic differentiation of human bone marrow stromal cells113 and can lead to suppression of osteoclast activity166. The sol-gel glass synthesis solution was used to impregnate the BG scaffolds. The coating increased the mechanical properties only of the

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BG-PU scaffolds but no effects were detected on BG-SA foams. A possible explanation is that the thickness of the coating was too low to have an effect of the maximum compressive strength of the samples prepared with natural marine sponges as template. The MBG coated BG scaffolds were tested in SBF. The effect of the degradation was much less severe for the MBG coated BG scaffolds (Chapter 6, Section6.3.4) compared to the non-coated ones (Chapter 4, Section 4.3.5), which in the same testing condition, lost almost 50% of the starting weight. This could be explained by the increased bioactivity observed for the MBG coated BG scaffolds compared to the non-coated ones. Due to the high specific surface area and the high number of silanol groups available on the MBG, the degradation of BG and deposition of HCA can reach faster the equilibrium and HCA deposition was confirmed to dominate the behavior in SBF up to 28 days of immersion.

Summarizing, 45S5 Bioglass®-based scaffolds with improved mechanical properties and biological activity were successfully manufactured. The use of natural marine sponges as sacrificial templates for porous BG-based scaffold preparation was demonstrated as promising alternative to PU foams. The obtained scaffolds are characterized not only by the high interconnectivity of their porous structure but also by sound mechanical properties. Ordered mesoporous silica particles MCM-41_B and Ag_MCM-41_B were successfully developed combining two synthesis procedures. These particles were used as a functional coating for BG-based scaffolds. The starting approach was to combine the drug up-take/ release capability of the silica particles with the bioactivity of the BG scaffolds. By this combination, extra properties were obtained. For example, the mechanical properties of the coated scaffolds increased significantly and, surprisingly, also the bioactivity of the system was enhanced. In conclusion, it was possible to obtain a novel functional scaffold system with possible drug delivery capability, increased mechanical properties and enhanced bioactivity confirming that MCM-41_BG scaffolds exhibit attractive properties for bone tissue engineering applications. BG scaffolds were also coated with MBG glasses. The MBG BG-SA samples exhibited improved bioactivity properties but no effects were observed on the mechanical properties. In conclusion, it was possible to obtain functional systems based on MBG coated BG scaffolds with improved bioactivity and with the intrinsic capability to release both functional ions and drugs.

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7.2. Future Work Based on the results obtained within this work, different research areas appear as interesting for future investigations and several questions are identified, which must be addressed for the advancement of the field.

The development of BG-based scaffolds with improved mechanical properties was addressed thanks to the combination of the well-established replica method15 with natural marine sponges as sacrificial template. The improvement of mechanical compressive strength was obtained by a reduction of the total porosity of the resulting 3D scaffold. Although preliminary studies suggest that these scaffolds are characterized by interconnected porosity and bioactivity comparable to those of BG-PU foams, more detailed in vitro analyses need to be done to assess the cell attachment and proliferation on these scaffolds. In fact, scaffolds porosity, pore size and interconnected porosity must be sufficient to allow cell attachment, cell ingrowth and neovascularization of 3D scaffolds54, but this need to be confirmed by dedicated cell culture studies and by in-vivo investigations in suitable models (e.g. the AV- loop model of vascularization119).

Another key aspect is the improvement of the production process of the scaffolds via replica method. So far, all the production steps are manually done by a technician and the properties of the resulting BG-based foams, especially total porosity, are affected by the operator’s manual skills. Therefore, in order to make the process more reproducible and to reduce the processing time, an automatization of the process should be attempted. The automatization of the production process will also be required for the up-scaling production in case of future clinical applications.

Another key question that needs to be answered is related to the effect of ordered mesoporous silica particles, used as functional coating of scaffolds, on cells. So far, silica particles have been mainly used to address cancer therapy86 and they were never investigated with a solid 3D porous structure in regenerative applications. In the present work, the positive effect of these particles on the bioactivity and mechanical properties of the BG- based foams was proved, but in vitro cell culture test must be done to prove the absence of any cytotoxicity effects of the combined system. In fact, MCM-41 particles showed a fast Si ion release during the first ours in SBF, amounts that need to be confirmed to not negatively affect cells. Moreover, it was not evaluated how the improved bioactivity of the MCM- 41_BG scaffolds could affect the mechanical properties of the foams overtime.

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The same evaluations should be done in case of BG-based scaffolds coated with sol-gel glasses. In fact, the use of biologically active ions in the sol-gel glass compositions needs further investigations. Indirect cell culture tests need to be done to confirm that the different compositions do not show any cytotoxic effect and to prove that the selected ions have the predicted positive effect on osteogenesis, angiogenesis and antibacterial efficiency.

The aim of the present work was to develop a functional coating of BG-based scaffolds with ordered mesoporous silica-based materials. Both MCM-41 particles and sol-gel glasses were successfully used as coating of BG-scaffolds. Theoretically, both MCM-41 and MBGs should release biologically active ions and drugs overtime. A challenge for the design of such scaffolds with dual delivery (therapeutic ions and drugs) is to be able to avoid the release of the functional molecules and ions during the preconditioning step of the BG-based scaffolds. In fact, as shown in many works reported in literature143, in vitro test were not successful when BG-based scaffolds were not preconditioned for many days in PBS, SBF or cell culture medium until the increase of pH was reduced to < 7.8. In case of coated scaffolds, during the preconditioning step, all functional ions and molecules could be released. In this work a reduction of the fast pH increase of coated samples was detected compared to that of non-coated BG-scaffolds. MCM-41 particles and MBGs showed a passivating effect on the BG-based structure in SBF, but no in vitro test have been carried out to prove that this effect was strong enough to avoid a pH increase over physiological values. In the worst case scenario that in vitro tests prove that this passivation effect is not sufficient, the functional coating should be applied on already preconditioned BG-based scaffolds. Clearly, the use of other bioactive glasses with reduced bioreactivity, like the composition 13-93137, is another alternative.

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Appendix - Oxygen Diffusion Evaluationbb

The effective oxygen diffusivity in BG based scaffolds was determined using Lattice Monte Carlo method133,134. This numerical method simulates random walks of probing particles within lattice models that represent the target geometry, in this case the interconnected porosity of BG scaffolds. For geometric accuracy, lattice models were derived directly from high resolution µ-CT scans. Scans were obtained from scaffolds in as-fabricated conditions and after immersion in SBF for 28 days. Different samples had to be used for the scan of as- sintered and resorbed scaffolds since the electron beam of the µ-CT scan could have affected the structure of the BG scaffolds. In order to remove calculation errors due to uneven samples, surfaces cubical sub-volumes were extracted from the center of each µ-CT data set. The used model approach, following a previous study on diffusivity in scaffolds135, assumes the worst case scenario of no vascularization and no diffusion through the scaffold material. Thus, the interconnected pores (porosity �) were assumed to be completely occupied by regenerated tissue with a bulk diffusivity �!. At the same time, vascularization was assumed to be incomplete and not yet contributing to the oxygen transport. The diffusivity of the scaffold material itself was approximated by zero. Thus, scaffold material act as a local diffusion barrier and reduces the effective diffusivity �!"" of the new formed biological ∗ tissue. The dimensionless diffusivity � is calculated as �!""/�! and allows the exploration of results to different types of tissues. Since the geometry of scaffolds obtained with PU as sacrificial template and scaffolds obtained using natural marine sponges are characterized by different structures, a different theoretical approach was used. For the marine natural sponges based scaffolds, due to their bimodal nature of the pore size distribution, the matrix was considered to be consisted of BG with smaller sized interconnecting pores, and large pores as inclusions into matrix phase. Next, the amount of porosity in the matrix phase was considered as a fraction g of the total porosity �. Then the total volume fraction �! of the matrix phase is:

�! = 1 − � + �� Eq.3

And the normalized effective diffusivity �! of the matrix phase is:

�! = ��/(1 − � + ��) Eq.4 For the total normalized effective diffusion coefficient there is a Maxwell-type185,186 expression available:

!"" ! (!!!!)(!!!!) ! ! (!!!!) �! = �! 1 − = ! Eq.5 !!!!!!(!!!!)(!!!!) !!!!!! ! !!! !!

bb The diffusivity analysis was carried out at the University of Newcastle in Australia by Dr. Thomas Fiedler.

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The fraction � is not known, for this reason it was treated as a fitting parameter using the results of Lattice Monte Carlo simulations of the effective normalized diffusion coefficient. The scaffolds produced using PU as template are high-porosity open cell structures and the Dulnev model187 provides best results. The model should describe well both diffusion along the struts134 and diffusion in the open space of the structure. Then, according to this model, the effective normalized diffusion coefficient is expresses as: ! �!"" = 1 − (1 − �) ! Eq.6 The generalized Scaling relation approach �!"" = �! Eq.7 was used for comparison. In a previous work of Fiedler et al.135 it was found that Eq.7 gives the best agreement for the scaffolds obtained with PU as template when � ≈ 2.

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List of Figures

Fig. 1.1 Schematic overview of the research approach of the present work: 3 preparation of BG scaffolds with improved mechanical properties and development of functional coatings with ordered mesoporous silica materials, i.e. MCM-41 and Co-MBG and SrCu-MBG, for the release of metallic ions and drugs once in contact with the biological fluids.

Fig. 2.1 Schematic illustration of bone formation/bone remodeling process 7 including osteoblast and osteoclast differentiation from mesenchymal and hematopoietic stem cell, respectively; matured osteoblast become osteocytes and are entrapped in the bone matrix.

Fig. 2.2 Bone mechanical homeostasis mediated by different loads. When 8 loads are in a physiological range (εmin<ε<εmax) deposition and resorption of bone matrix are in perfect balance. When ε<εmin the resorption of bone is preponderant, when ε<εmax the deposition is faster than resorption.

Fig. 2.3 Schematic diagram of key factors in the design of optimal scaffolds 9 for bone tissue engineering (figure modified from Gerhardt and Boccaccini6).

Fig. 2.4 a) powder particles compacted together, b) particles beginning to bond 12 together, c) fully sintered ceramic (figure modified from Narayan et al.17).

Fig. 2.5 Relation between glassy, solid and liquid state. Heating a crystal its 15 specific volume increases (line F-E) and in correspondence of the melting temperature (Tm) there is high increase of the specific volume (line E-B) because the solid-liquid transition; later the volume increase linearly (line B-A). When the liquid is cooled, if the cooling rate is slow, the curve will follow the same curve of heating (line A-B-E-F), if the cooling rate is fast the liquid passes through point B without abrupt change of specific volume and without crystallizing (line A-B- 27 C-D). Tg is the temperature of glass transition .

Fig. 2.6 Structure of the crystalline and amorphous silica. 15

Fig. 2.7 Structure of a glass SiO2-Na2O 16

Fig. 2.8 Overview of biological response to ionic dissolution products of 17 bioactive glasses (figure modified from Hoppe et al.26).

Fig. 2.9 Compositional diagram for bone bonding. Materials belonging to the 18 region A are materials able to bond with bone; region S is a region of class A bioactivity where bioactive glasses bond to both bone and soft tissue and are gene activation, as 45S5 Bioglass® (figure modified from Hench9).

Fig. 2.10 Schematic illustration of the surface stages (i-v) reactions on bioactive 20

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glasses, forming double SiO2- rich layer and Ca-P- rich layer (figure modified from Peitl et al.26). Fig. 2.11 Sequence of interfacial reactions involved in forming a bond between 21 bone and bioactive glass (figure modified from Peitl et al.26).

Fig. 2.12 Two different mechanisms proposed for the production of silica 29 ordered mesoporous materials. Observe that the main difference occurs when adding the silica precursor (TEOS). After the surfactant has added, the supramicellar phase is formed (route 1), or at the same time as the micellar mesophase in being formed (route 2). This figure has been adapted from Vallet-Regí et al.43.

Fig. 2.13 Schematic representation of the post-synthesis functionalization 30 method of ordered mesoporous materials. This figure has been adapted from Vallet-Regí et al.43.

Fig. 2.14 Schematic representation of the co-condensation functionalization 31 method of ordered mesoporous materials. This figure has been adapted from Vallet-Regí et al.43.

Fig. 2.15 Schematic representation of some of the most used and investigated 32 ordered mesoporous silica particles with their respective pore sizes.

Fig. 2.16 Schematic representation of synthesis of MBGs through the EISA 34 method (picture modified from Hum et al.113).

Fig. 4.1 Removal of the extra slurry from the templates by compressed air in 47 order to keep an open porosity of the resulting BG scaffolds.

Fig. 4.2 Scheme for the preparation of Bioglass® based scaffolds via replica 48 method and heat treatment programme designed for burning out templates (400°C – 1h) and sintering BG (1050°C – 1h).

Fig. 4.3 Digital camera images of SA (a1) and SL (a2) sponges and PU foam 52 (b) templates; inhalant surfaces of SA (c) and SL (d) sponges; exhalant SA (e) and SL (f) sponges; fibrous network of SA (g) and SL (h) sponges (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis60)

Fig. 4.4 Replicated scaffolds derived from Spongia Agaricina (BG-SA), 53 Spongia Lamella (BG-SL) and polyurethane packaging foam (BG- PU): light microscopy and scanning electron microscopy (SEM) images. (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”129)

Fig. 4.5 a)b) BG-PU scaffolds at different magnification, c) inhalant and d) 54 exhalant surface of BG-SA, e) and f) inhalant and exhalant surface of BG-SL. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis60)

Fig. 4.6 SEM micrographs of cross sections of (a) BG-PU and (e-f) BG-SA 55

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scaffolds to observe the strut microstructure of these BG scaffolds; schematic cross section of (b) BG-PU and (c) BG-SA struts.

Fig. 4.7 (a) SEM micrographs of the surface of the BG scaffolds after the 56 sintering process, which induces a partial crystallization of the starting BG. The different phases formed during the thermal treatment are: (b) crystalline phases, combeite and rhenanite (see XRD results below), (c-d) amorphous glassy particles fused together.

Fig. 4.8 EDX analysis of the BG scaffold surface in order to identify the 57 different phase compositions.

Fig. 4.9 2D section and 3D reconstruction of a),d) BG-PU, b),e) of BG-SL and 58 f,g BG-SA obtained by µ-CT analysis. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis60)

Fig. 4.10 XRD spectra of BG powder, BG-SA and BG-SL scaffolds showing 59 the crystalline phases formed after the thermal treatment. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis60)

Fig. 4.11 Stress-displacement curves of BG foams manufactured with different 60 sacrificial templates. (©2015 Boccardi, Philippart, Juhasz-Bortuzzo, Novajra, Vitale-Brovarone, Boccaccini. Published by Taylor & Francis60)

Fig. 4.12 (a) Weight and (b) porosity variation of BG-SA and BG-SL scaffolds 61 after different immersion time in SBF for up to 28 days. The standard deviation values of the data for BG-SL in (b) were too low to be visible. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”128)

Fig. 4.13 Compressive strength variation of BG-SA and BG-SL scaffolds after 62 immersion in SBF for up to 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale- Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”128)

Fig. 4.14 Stress-displacement curves of BG-SA scaffolds after immersion in 62 SBF for up to 28 days.

Fig. 4.15 Concentration of (a) Si, (b) Ca and (c) P ions released from BG 63 scaffolds during the first hours of immersion in SBF for up to 3 days; (d) pH variation in the first 48 h of immersion in SBF. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6),

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Fig. 4.16 SEM micrographs of BG-SA (a-f), BG-SL (g-l) and BG-PU (n-s) 64 scaffolds at different immersion times in SBF (for up to 28 days) showing evolution of HCA) deposit. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale- Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”128)

Fig. 4.17 SEM micrographs of cross sections of single struts of the BG-PU (a,c) 65 and BG-SA (b,d) scaffolds after immersion in SBF for 14 days showing the evolution of the HCA phase both on the external surface and inside the hollow structure of the struts of the BG scaffolds.

Fig. 4.18 FTIR analysis on BG-SA (a), BG-SL (b) and BG-PU (c) scaffolds 66 before and after immersion in SBF for different number of days and SEM micrograph (d) showing the needle like structure of HCA formed on the BG-SA surface after immersion in SBF for 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”128)

Fig. 4.19 µ-CT reconstructions of BG-SA, BG-SL and BG-PU scaffolds before 67 and after immersion in SBF for up to 28 days (a), pore size variation of BG-SA (b), and BG-SL (c) scaffolds after immersion in SBF for up to 28 days. (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone, Fey, Boccaccini (©2016 Biomedical Engineering Society) “With permission of Springer”128)

Fig. 4.20 Porosity variation in BG-SA (a), BG-SL (b) and BG-PU (c) scaffolds 68 in as-fabricated conditions and after the immersion in SBF for 28 days. (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”129)

Fig. 4.21 (a) Cell viability results from indirect cell culture test with different 69 dilutions of the extraction volumes, SEM micrographs of (b) BG-PU, (c) BG-SA and (d) BG-SL scaffolds after 14 days of direct cell culture study, (e) cell viability results from direct cell culture for up to 14 days (* statistically different). (Annals of Biomedical Engineering, Bioactivity and mechanical stability of 45S5 bioactive glass scaffolds based on natural marine sponges, 44(6), 2016, 1881, Boccardi, Philippart, Melli, Altomare, De Nardo, Novajra, Vitale-Brovarone,

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Fig. 4.22 Normalised oxygen diffusivity data plotted versus scaffold porosity 70 for BG-SA, BG-SL and BG-PU scaffolds. (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”129)

Fig. 4.23 Anisotropy of normalized scaffold bioactivity showing the porosity 72 variation inside the BG scaffolds along different directions. (J Mater Sci: Mater Med, Oxygen diffusion in marine-derived tissue engineering scaffolds, 26, 2015, 200, Boccardi, Belova, Murch, Boccaccini, Fiedler (©Springer Science+Business Media New York 2015) “With permission of Springer”129)

Fig. 4.24 Summary of 45S5 BG structural transformations identified by 74 Lefebvre et al.141

Fig. 4.25 Schematic representation of the two phenomena taking place after 75 immersion of BG scaffolds in SBF: the BG starts to dissolve on the external surface and in the inner core of the struts, the HCA starts to form as a homogenous layer on the external surface of the scaffolds and inside the BG struts.

Fig. 5.1 Flowchart of the different methods developed for the production of the 83 composite system MCM-41 coated BG scaffolds.

Fig. 5.2 HRTEM images of sample MCM-41_A (a,b) and sample MCM-41_B 87 (c,d) which were characterized by high ordered mesoporosity; sample MCM-41 (e,f) and sample MCM-41_D (g,h), which were characterized by a disordered porosity. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.3 High-resolution image of the ordered mesoporous structure of MCM- 88 41_B after analysis with FFT and inverse FFT (a); plot of the distance between the parallel pore channels obtained with ImageJ plug in Plot Profile along the yellow line (b). (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.4 SAXRD analysis of samples MCM-41_A (a), MCM-41_B (b), MCM- 89 41_C (c) and MCM-41_D (d). MCM-41_A particles were characterized by the three peaks, labelled 100, 110 and 200. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale- Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.5 Nitrogen adsorption desorption isotherm of MCM-41_B particles. 89 (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering

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and Biotechnology64) Fig. 5.6 HRTEM images of the porous structure of samples prepared with 90 different water/ethanol ratios: (a-b) sample MCM-41_A, no ethanol and 20 min of stirring following the standard synthesis procedure, (c- d) sample MCM-41_B 60% of ethanol and 20 min of stirring, (e-f) sample MCM-41_C, 90% of ethanol and 2 h of stirring. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale- Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.7 Ibuprofen released profile form MCM-41_A (a), MCM-41_B (b), 91 MCM-41_C (c) and MCM-41_D (d) particles. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.8 SEM micrographs of MCM-41 particles as synthetized (a), after the 92 immersion in SBF for 1 day (b) and 3 days (c); SEM micrographs of MCM-41_Ref particles as synthetized (d), after the immersion in SBF for 1 day (e) and 3 days (f).

Fig. 5.9 SEM-EDS analysis of MCM-41_B particles before (a) and after (b) 93 the immersion in SBF for 28 days; SEM-EDS analysis of MCM- 41_Ref particles before (c) and after (d) the immersion in SBF for 28 days.

Fig. 5.10 HRTEM images of the MCM-41_B particles as synthetized (a) and 94 after the immersion in SBF for 3 days (b), 14 days (c) and 28 days (d); HRTEM images of MCM-41_Ref as synthetized (e) and after the immersion in SBF for 3 days (f), 14 days (g) and 28 days (h).

Fig. 5.11 Plot profile of the distance between the pore channels characteristic of 95 the MCM-41_B and MCM-41_Ref particles before and after immersion in SBF for up to 28 days.

Fig. 5.12 FTIR spectra of MCM-41_B particles before and after the immersion 96 in SBF for up to 28 days. In the figure the peaks associated to the Si-O stretching mode are marked.

Fig. 5.13 pH variation of the SBF containing the silica particles and the control 97 for up to 28 days of test (a); ICP analysis of the Si ion release for up to 7 days (b) for MCM-41_Ref and MCM-41_B samples.

Fig. 5.14 SEM micrographs of MCM-41_B particles (a) and Ag_MCM-41_B 98 particles obtained with different concentrations of silver nitrate: 0.6 mg mL-1 (b), 3 mg mL-1 (c) and 6 mg mL-1 (d). Backscattering SEM micrographs of Ag_MCM-41_B particles obtained using the highest concentration of silver nitrate. FTIR spectra (f) of the non- functionalized and the Ag-functionalized MCM-41_B particles obtained with different silver nitrate concentrations.

Fig. 5.15 HRTEM images of MCM-41_B (a) and Ag_MCM-41_B (c) particles, 99 which were characterized by high ordered mesoporous structure. High-resolution images of the ordered mesoporous structures of

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MCM-41_B (b) and Ag_MCM-41_B (d) particles after analysis with FFT and inverse FFT. Plot of the distance between the pore channels obtained with ImageJ plug in plot of MCM-41_B (e) and Ag_MCM- 41_B (f) particles.

Fig. 5.16 HRTEM image of Ag-MCM-41_B particles (a); EDX analysis of a 100 single silver nanoparticle on the surface of a MCM-41_B particle (b); silver nanoparticles size distribution (c); XRD analysis of functionalized Ag_MCM-41_B and not functionalized MCM-41_B mesoporous silica nanoparticles.

Fig. 5.17 SEM images of Ag_MCM-41_B particles before and after immersion 101 in SBF solution for up to 7 days (a-d); EDX analysis of Ag-MCM- 41_B particles before and after 7 days of immersion in SBF.

Fig. 5.18 FTIR spectra of Ag_MCM-41_B particles before and after 7 days of 102 immersion in SBF.

Fig. 5.19 Concentration of (a) Si and (b) Ag ions released from Ag_MCM-41_B 103 particles for up to 7 days of immersion in SBF; (c) pH variation after the immersion of functionalized and non-functionalized MCM-41_B in SBF.

Fig. 5.20 Ibuprofen release profile from MCM-41_B and Ag_MCM-41_B 104 particles for up to 14 days.

Fig. 5.21 BG-PU scaffolds coated with MCM-41_A (a-b), MCM-41_B (c-d) 105 and MCM-41_D (e-f) particles by CCM procedure. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.22 HRTEM images of MCM_41_B particles coating the BG-PU 105 scaffolds at different magnifications in order to show that the ordered mesoporous structure of the material was not affected by the CCM method. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.23 SEM micrographs of BG-PU scaffolds uncoated (a) and coated with 107 MCM-41_B (b-d) after 1 week in SBF; and of MCM-41_B_BG-PU scaffolds, after immersion in TRIS buffered solution for 10 days (e-g). (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 5.24 SEM micrographs of CCM MCM-41_B_BG-SA external (a) and 108 internal (b) surfaces and of CCM MCM-41_B_BG-PU external (c) and internal (d) surfaces.

Fig. 5.25 SEM micrographs of BG-PU scaffolds coated with MCM-41_B using 109 the Post Coating Method A (PCM_A method, wet particles) using

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different concentration of wet particles in ethanol.

Fig. 5.26 SEM micrographs of BG-PU scaffolds coated with MCM-41_B 110 particles PCM_B method obtained using different concentrations for scaffolds impregnation.

Fig. 5.27 SEM micrographs of external (a) and internal (b) surface of BG-SA 110 scaffolds coated with MCM-41_B using PCM_B method.

Fig. 5.28 SEM micrographs of CCM MCM-41_B BG-SA (a) BG-PU (b) 112 scaffolds and PCM_B MCM-41_B BG-SA (c) and BG-PU (d) scaffolds after immersion in SBF for 1day.

Fig. 5.29 SEM micrographs of a cross section of a PCM_B MCM-41_B BG-SA 113 scaffold after the immersion in SBF for 3 days at different magnifications showing a) that the strut is seen to be covered in HCA and that b-c) the MCM-41_B particles have been incorporated in the developing HCA.

Fig. 5.30 FTIR spectra of CCM MCM-41_B BG-PU and BG-SA scaffolds and 114 PCM_B MCM-41_B BG-PU and BG-SA scaffolds before and after the immersion in SBF for up to 7days. The relevant peaks are discussed in the text.

Fig. 5.31 pH variation of after immersion in SBF for up to 7 days: CCM MCM- 115 41_B BG-SA, CCM MCM-41_B BG-PU, PCM_B MCM-41_B BG- SA, and PCM_B MCM-41_B BG-PU scaffolds.

Fig. 5.32 Dru129g-release profile from BG-PU scaffolds coated with ordered 115 mesopo130rous silica particles compared to non-coated scaffolds. (©2015, Boccardi, Philippart, Juhasz-Bortuzzo, Beltràn, Novajra, Vitale-Brovarone, Spiecker, Boccaccini. Frontiers in Bioengineering and Biotechnology64)

Fig. 6.1 Optical images of Co-MBG and SrCu-MBG before EISA (a,b), after 128 EISA (c,d) and after the thermal treatment at 700°C for 6 h (e,f).

Fig. 6.2 HRTEM images of SrCu-MBG (a) and Co-MBG (b) glasses; plot of 129 the distance between the parallel pore channels obtained with ImageJ plug in Plot Profile.

Fig. 6.3 XRD analysis of SrCu-MBGs and Co-MBGs after the thermal 130 treatment.

Fig. 6.4 SEM images of Co-MBG and SrCu-MBG before (a,b) and after the 130 immersion in SBF for 1 h (c,d).

Fig. 6.5 SEM micrographs of Co-MBG and SrCu-MBG powders immersed in 130 SBF for 1d (a,b) and 3d (c,d).

Fig. 6.6 ICP analysis of Si, P, Ca, Co, Cu and Sr ions release for up to 7 days 131 in SBF.

Fig. 6.7 Drug release profile of Co-MBG (a) and SrCu-MBG loaded for with 132

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drug using three different up-take times, i.e. 6, 12 and 24 h.

Fig. 6.8 SEM micrographs of BG pellets heat treated at 1050°C, similarly to 133 the sintering of the scaffold before (a) and after (c) the immersion in TRIS buffered solution for 14 days, BG pellets double thermal treated, first at 1050°C and after at 700°C before (b) and after (d) immersion in TRIS buffered solution; SEM-EDS analysis of the BG pellets after 14 days in SBF (e,f).

Fig. 6.9 Digital camera images of BG-PU scaffolds coated up to 6 times by dip 134 coating procedure with Co-MBG synthesis solution after the thermal treatment at 700°C for 6h.

Fig. 6.10 SEM micrographs of BG BG-PU scaffolds dip coated in sol-gel glass 134 synthesis solution increasing the number of coating cycles, from 2 up to 5.

Fig. 6.11 SEM micrographs of Co-MBG BG-PU dip coated 4 times with sol-gel 135 glass solution and immersed in SBF for 1 day (a) and 7 days (b).

Fig. 6.12 Weight variations of BG-SA scaffolds impregnated up to 5 times with 136 Co-MBG and SrCu-MBG synthesis solutions.

Fig. 6.13 SEM micrographs of BG-PU scaffolds impregnated 1, 2 and 3 times 136 with Co-MBG sol-gel glass synthesis solution.

Fig. 6.14 SEM micrographs of Co-MBG BG-SA (a), SrCu- MBG BG-SA (b), 137 Co-MBG BG-PU (c) and SrCu- MBG BG-PU (d) scaffolds after one impregnation with sol-gel glass synthesis solutions.

Fig. 6.15 SEM micrographs of BG-SA scaffolds impregnated with Co-MBG 138 sol-gel glass synthesis solution at different magnifications (a, b and d) compared to the non-coated BG-SA scaffold (c). SEM image of a cross section of a Co-MBG BG-SA scaffold strut (e).

Fig. 6.16 Co-MBG BG-SA and SrCu-MBG BG-SA scaffolds weight (a) and 139 porosity (b) variation after immersion in SBF for up to 28 days.

Fig. 6.17 Maximum compression strength variations of Co-MBG and SrCu- 140 MBG BG-SA scaffolds after immersion in SBF for up to 28 days. The SBF was refreshed every 2-3 days.

Fig. 6.18 ICP Concentration of (a) Si, (b) P, (c) Ca ions released from Co- 141 MBG BG-SA and SrCu-MBG BG-SA scaffolds up to 7 days in SBF, without SBF refresh every 2-3 days; pH variation in 7 days of immersion in SBF, without SBF refresh.

Fig. 6.19 SEM micrographs of the external surfaces of Co-MBG BG-SA (a), 142 SrCu-MBG BG-SA (b), Co-MBG BG-PU (c) and SrCu-MBG BG-PU (d) scaffolds after immersion in SBF for 1 day.

Fig. 6.20 SEM micrographs of the inner core of Co-MBG BG-SA (a), SrCu- 143 MBG BG-SA (b), Co-MBG BG-PU (c) and SrCu-MBG BG-PU (d) scaffolds after the immersion in SBF for 1 day.

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Fig. 6.21 SEM micrographs of the external surface of Co-MBG BG-SA 144 scaffolds before the immersion in SBF (a) and after the immersion for 3 days at different magnifications (b-dc) and high magnification images of the HCA layer formed (e).

Fig. 6.22 SEM micrographs of Co-MBG BG-SA (a), SrCu-MBG BG-SA, Co- 145 MBG BG-PU (c) and (b) and SrCu-MBG BG-PU (d) scaffolds after immersion in SBF for 7 days.

Fig. 6.23 SEM micrographs of inner strut cross sections of Co-MBG BG-SA 146 scaffolds (a-e) and SrCu-MBG BG-SA scaffolds (f-l) before and after immersion in SBF for 1, 7, 14 and 28 days.

Fig. 6.24 SEM micrographs of inner strut cross sections of Co-MBG BG-PU 147 scaffolds (a-d) and SrCu-MBG BG-PU scaffolds (e-h) after immersion in SBF for 1, 7, 14 and 28 days.

Fig. 6.25 FTIR spectra of (a) Co-MBG BGPU, (b) SrCu-MBG BG-PU, (c) Co- 148 MBG BG-SA and (d) SrCu-MBG BG-SA scaffolds before and after the immersion in SBF for up to 28 days. The relevant peaks are discussed in the text.

Fig. 6.26 Optical images of BG-SA scaffolds (a) non-coated (b) coated with Co- 149 MBG and (c) coated with SrCu-MBG synthesis solutions; BG-PU scaffolds (d) non-coated, (e) coated with Co-MBG and (f) coated with SrCu-MBG synthesis solutions.

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List of Tables

Table. 2.1 Composition of bone tissue (modified from Redaelle and 5 Montevecchi16)

Table 2.2 Design criteria for bone tissue engineering scaffolds6,19 table 9 modified from Chen et al.19)

Table 4.1 Summary of the natural marine sponge’s architecture properties, 53 focusing on pore dimension and structural thickness.

Table 4.2 Architectural properties of BG based scaffolds prepared with PU 58 and natural marine sponges as sacrificial templates.

Table 4.3 Diffusivities and structural parameters of tissue engineering 71 scaffolds (table modified from Boccardi et al.129)

Table 5.1 Composition of four different synthesis solutions used for the 81 preparation of MCM-41 particles. MCM-41_A particles were prepared following the work of Zeleňák et al. MCM-41_D particles were prepared following the work of Grün et al. MCM- 41_B and MCM-41_C particles were developed combining the two previous synthesis solutions (table modified from Boccardi et al.129)

Table 5.2 Summary of the maximum compressive strength of the DTT and 111 coated scaffolds.

Table 6.1 Maximum compressive strength values of BG-PU scaffolds coated 135 with Co-MBG increasing the dip coating cycles up to 6.

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