Bioactive Glass Combined with Natural Derived Proteins As Composite Materials for the Application in Bone Tissue Engineering

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Bioactive glass combined with natural derived proteins as composite materials for the application in bone tissue engineering

Kompositmaterialien bestehend aus bioaktivem Glas und natürlichen Proteinen für die Anwendung im Knochen Tissue Engineering

Der Technischen Fakultät

der Friedrich-Alexander-Universität Erlangen-Nürnberg

zur Erlangung des Doktorgrades Dr.-Ing.

vorgelegt

von

Jasmin Katharina Hum

aus Nürnberg

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

Tag der mündlichen Prüfung: 17. Februar 2016

Vorsitzender des Promotionsorgans: Prof. Dr. rer. nat. Peter Greil

Gutachter: Prof. Dr.-Ing. habil. Aldo R. Boccaccini Prof. Valeria Cannillo Prof. Dr. rer. nat. Armin Bolz

Danksagung

An dieser Stelle möchte ich mich ganz herzlich bei all denjenigen bedanken, die direkt und indirekt an der Entstehung dieser Arbeit mitgewirkt haben. In erster Linie gilt mein Dank Prof. Aldo R. Boccaccini, der mir nicht nur die Möglichkeit gegeben hat, den Lehrstuhl Biomaterialien von Anfang an mit etablieren zu können, sondern auch durch die Aufgabenstellung der Doktorarbeit meine Kreativität angeregt und mir viele neue wissenschaftliche Erfahrungen näher gebracht hat. Seine Denkanstöße und Vorschläge habe ich stets geschätzt. Vielen Dank auch an Prof. Valeria Cannillo, Prof. Armin Bolz, Prof. Ulrich Lohbauer und Prof. Sannakaisa Virtanen, die sich die Zeit genommen haben, die Arbeit letztendlich zu begutachten und zu bewerten.

Dank mehrerer Kooperationen u.a. mit dem Lehrstuhl Glas und Keramik und dem Lehrstuhl Korrosion und Oberflächentechnik hatte ich Zugriff zu unterschiedlichsten Messmethoden, die mir die Charakterisierung meiner Proben wesentlich erleichterten. Auch wenn die Interpretation des ein oder anderen Ergebnisses manchmal herausfordernd war, konnte für jedes Rätsel doch eine Erklärung gefunden werden. Danke daher an Judith Roether, Yaping Ding, Dirk Dippold, Helga Hildebrand, Sabine Brungs und Eva Springer, die mir bei der Durchführung der Messungen hilfreich zur Seite standen.

Ebenso möchte ich mich beim Team Henkestraße bedanken, vor allem Alina Grünewald und Rainer Detsch, mit denen ich hilfreiche wissenschaftliche Diskussionen führen konnte. Prof. Ben Fabry danke ich für die Bereitstellung seiner Labore.

Während der letzten Jahre haben viele Studenten auf der Suche nach Rat an meine Tür geklopft. Meistens konnte ich helfen, doch auch ich habe viel Unterstützung erhalten. Besonders möchte ich zwei meiner Studenten erwähnen. Shiva Naseri, immer hilfsbereit und wissbegierig, und Jonathan Potschka, der seine Zeit im Labor mit der Herstellung und Funktionalisierung von Scaffolds verbrachte. Danke für eure Hilfe.

Neben der wissenschaftlichen Laborarbeit ist allerdings ein Punkt, der nicht vergessen werden darf: das soziale Umfeld. Ich hatte das Glück mit wunderbaren Menschen sowohl Labor wie auch Büro teilen zu können. Danke der spanischen Welle mit all eurem Temperament. Anahí Philippart, Sandra Cabañas Polo, Valentina Miguez Pacheco, Luis Eduardo Cordero Arias und Micael Alonso Frank, mit euch konnte ich immer herzhaft lachen. Danke auch Elena Boccardi, Stefanie Spallek und Petra Rosner für die schönen Erinnerungen. Es war eine wunderbare Zeit. Auch den restlichen Kollegen am Institut danke ich für die zahlreichen Diskussionen, die so manchen Geistesblitz während meiner Dissertation brachten.

Ganz herzlich danke ich auch meinen lieben Zimmernachbarn Sigrid Seuß, Giulia Rella und Liliana Liverani für wunderbar warmes Arbeitsklima, eine angenehme Wohlfühlatmosphäre in unserem Büro und die daraus erwachsenen Freundschaften, die ich nicht mehr missen möchte. Heinz Mahler, Gerhard Frank und Bärbel Wust danke ich ganz herzlich für jegliche Unterstützung im technischen und administratorischen Bereich, die zu jeder Zeit hilfsbereit waren. Auch für sonstige Probleme hattet ihr stets ein offenes Ohr. Danke, dass ich immer willkommen war.

Zu guter Letzt geht ein großes Dankeschön an meine Freunde und Familie, vor allem Steffi und Hariet, die mir außerhalb des wissenschaftlichen Alltags immer eine große Stütze waren und die trotz vieler offener Fragen Interesse an meiner Arbeit gezeigt haben. Danke Christian Dolle für Ahorn oder Eiche, unsere Kaffeepausen am See und dafür, dass Du immer für mich da bist.

Während meiner Promotion war ich nicht nur Doktorandin, sondern auch Laborleitung, Sekretärin und noch viel mehr. Flexibilität war an der Tagesordnung und ich nehme aus dieser Zeit sehr viel Erfahrung mit. Es war ein langer Weg, aber es ist geschafft. Danke an alle!

Index

Abbreviations and symbols...... IV Abstract...... VII Zusammenfassung...... IX

1 Introduction...... 1

2 Fundamentals – Application of scaffolds in bone tissue engineering...... 4

2.1 State of the art...... 4 2.1.1 Bone tissue engineering...... 4 2.1.2 Scaffolds and their fabrication technologies...... 7

2.2 Materials selection...... 9 2.2.1 45S5 Bioglass® – a silicate-based bioactive glass...... 9 Production, structure and properties...... 10 45S5 Bioglass® as scaffold material...... 12 2.2.2 Collagen – a natural polymer...... 13 Structure of collagen...... 13 Collagen for tissue engineering...... 14 2.2.3 Zein – a protein derived from corn...... 15 Structure and morphology...... 15 Zein as biomaterial...... 16

2.3 Objective of research...... 17

3 Collagen – Composite scaffolds with bioactive glass and collagen...... 19

3.1 Reinforcement of porous collagen scaffolds with bioactive glass particles – Manufacturing techniques, results and discussion...... 19 3.1.1 Scaffold preparation...... 19 Porous collagen scaffolds without bioactive glass particles...... 19 Reinforced porous collagen scaffolds with bioactive glass particles...... 20 Crosslinking process...... 21 3.1.2 Morphological and microstructural characterization...... 25 Pure collagen...... 25 Collagen with bioactive glass reinforcement...... 27 3.1.3 Swelling properties and degradation behavior...... 29 Swelling properties...... 29 Degradation behavior...... 31 3.1.4 Evaluation of bioactivity...... 32

Index

Preparation of simulated body fluid...... 32 Bioactivity study...... 32

3.2 Collagen as coating material on bioactive glass-based scaffolds – Manufacturing techniques, results and discussion...... 37 3.2.1 Sample preparation...... 37 Scaffold and pellet fabrication...... 37 Surface functionalization, collagen coating and optimization of parameters...... 39 Crosslinking process...... 51 3.2.2 Swelling properties and release behavior...... 54 Swelling properties...... 54 Release behavior...... 55 3.2.3 Evaluation of bioactivity...... 59 3.2.4 Mechanical characterization...... 63

4 Zein – Composite scaffolds with bioactive glass and zein...... 66

4.1 Reinforcement of porous zein scaffolds with bioactive glass particles – Manufacturing techniques, results and discussion...... 66 4.1.1 Scaffold preparation...... 66 Porous zein scaffolds without bioactive glass particles...... 66 Reinforced porous zein scaffolds with bioactive glass particles...... 66 4.1.2 Morphological and microstructural characterization...... 67 Pure zein...... 67 Zein with bioactive glass reinforcement...... 69 4.1.3 Swelling properties and degradation behavior...... 71 Swelling properties...... 71 Degradation behavior...... 72 4.1.4 Evaluation of bioactivity...... 73 4.1.5 Mechanical characterization...... 76

4.2 Zein as coating material on bioactive glass-based scaffolds – Manufacturing techniques, results and discussion...... 77 4.2.1 Sample preparation...... 77 Scaffold fabrication...... 77 Zein coating...... 78 Crosslinking process...... 80 4.2.2 Swelling properties and degradation behavior...... 81 Swelling properties...... 82 Degradation behavior...... 82 4.2.3 Evaluation of bioactivity...... 83 4.2.4 Mechanical characterization...... 87

Index

5 In vitro studies and comparison of the investigated systems...... 89

5.1 In vitro studies – a static cultivation system...... 89 5.1.1 Sample preparation...... 89 5.1.2 Cell culture and seeding...... 90 5.1.3 Cell viability and relative proliferation...... 91 5.1.4 Cell morphology...... 94

5.2 In vitro studies – a dynamic cultivation system...... 96 5.2.1 Evaluation of a bioreactor...... 96 Assembly of the system...... 96 Selection of parameters...... 97 Cell cultivation...... 100 5.2.2 Cell viability and LDH activity...... 101 5.2.3 ALP activity...... 103 5.2.4 Cell morphology...... 104

5.3 Comparison of the investigated systems...... 106 5.3.1 Polymeric scaffolds with and without ceramic reinforcement...... 106 Pore size and porosity...... 106 Mechanical properties and biodegradability...... 107 Bioactive behavior...... 108 5.3.2 Biopolymer-coated bioactive glass-based scaffolds...... 109 Pore size and porosity...... 109 Mechanical properties...... 108 Bioactive behavior...... 110 Biocompatibility and osteoinduction...... 110

6 Conclusion and future work...... 112

Appendix I – Chemicals and materials...... 114 Appendix II – Characterization methods...... 117 Literature...... 121

III

Abbreviations and symbols

η Eta (Denotation for viscosity)  Theta (Denotation for contact angle) ρ Rho (Denotation for density)  Sigma (Denotation for strength)  Phi (Denotation for work function)

A Cross-section ALP Alkaline phosphatase APTS (3-Aminopropyl)triethoxysilane β-TCP Beta-tricalcium phosphate BG Bioactive glass BrdU Bromodeoxyuridine BTE Bone tissue engineering

CaCl2 Calcium chloride CaP Calcium phosphate CDU Collagen digestion unit Coll Collagen

CO2 Carbon dioxide cl Crosslinked d Diameter DE Germany DMEM Dulbecco’s Modified Eagle’s Medium DMF N,N-Dimethylformamide

EB Binding energy

Ekin Kinetic energy EDC N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide EDTA Ethylenediaminetetraacetic acid ELISA Enzyme-linked immunosorbent assay FBS Fetal bovine serum FTIR Fourier transform infrared spectroscopy GA Glutaraldehyde h Height HA Hydroxyapatite He Helium HCl Hydrochloric acid

H2O Water

Abbreviations and symbols

H2SO4 Sulfuric acid HEPES 4-(2-Hydroxyethyl)piperazine-1-ethanesulfonic acid KBr Potassium bromide KCl Potassium chloride

K2HPO4 Dipotassium hydrogen phosphate LDH Lactate dehydrogenase m Mass MES 4-Morpholineethanesulfonic acid

MgCl2 Magnesium chloride

N2 Nitrogen NaCl Sodium chloride NAD Nicotinamide adenine dinucleotide NADH Reduced form of nicotinamide adenine dinucleotide

NaHCO3 Sodium hydrogen carbonate

NaN3 Sodium azide NaOH Sodium hydroxide

Na2HPO4 Disodium hydrogen phosphate

Na2SO4 Sodium sulfate NBO Nonbridging oxygen NHS N-Hydroxysuccinimide P Porosity PenStrep Penicillin-streptomycin PBS Phosphate buffered saline PCL Polycaprolactone PCR Polymerase chain reaction PGA Polyglycolic acid PLA Polylactic acid PLGA Poly(L-lactic-co-glycolic acid) pNP p-Nitrophenol pNPP para-Nitrophenylphosphate ppi Pores per inch PU Polyurethane PVA Polyvinyl alcohol Re Reynolds number rpm Revolutions per minute RT Room temperature SBF Simulated body fluid SEM Scanning electron microscopy

Abbreviations and symbols

Tg Glass transition temperature TGA Thermogravimetric analysis THF Tetrahydrofuran TRIS Tris(hydroxymethyl)aminomethane TRIS-HCl Tris(hydroxymethyl)aminomethane hydrochloride uc Uncrosslinked UPW Ultrapure water UV-vis Ultraviolet–visible v Flow velocity V Volume VEGF Vascular endothelial growth factor XPS X-ray photoelectron spectroscopy

ZnCl2 Zinc chloride

Abstract

The present study has focused on the development and characterization of novel composite materials based on bioactive glass and natural derived proteins for the possible application in bone tissue engineering. Due to the progressive world population aging and the limited ability of the human body to restore lost tissue, there is an increasing demand for the regeneration of tissue damaged by disease or trauma. Especially, the replacement and repair of bone tissue represents a significant challenge. Nowadays, methods for the treatment of damaged bone tissue focus mainly on autologous or allogeneic bone grafts. Nevertheless, many drawbacks are associated with such implants like limited availability, risk of infections, incompatibility or ethical concerns. Therefore, the interdisciplinary field of bone tissue engineering (BTE) considers the development of so-called scaffolds which can provide an alternative for auto- and allografts. Some type of scaffolds for BTE comprises inorganic synthetic materials such as hydroxyapatite or bioactive glasses, e.g. the 45S5 Bioglass® composition. The application of 45S5 bioactive glass is attracting increasing attention as this material offers many advantages like bioactivity, osteoconductivity and angiogenic effects. However, the high brittleness of bioactive glass-based scaffolds limits their application area in bone tissue engineering, particularly in load-bearing parts. To overcome this problem, composite materials are developed to combine the key characteristics of different material classes. In this work, improvement of the mechanical properties of bioactive glass-based scaffolds (of the 45S5 composition) fabricated by the foam replica technique was attempted by the combination with biopolymers. In particular, scaffolds were coated with the natural polymers collagen and zein. For the coating procedure with collagen, scaffolds were surface- functionalized, which was a necessary step for the covalent bonding of collagen to the bioactive glass surface. For the zein coating, scaffolds were dip-coated in a solution with 8 wt.% of zein. The collagen coating as well as the zein coating were further stabilized by chemically induced crosslinks. The microstructure and the homogeneity of the applied coatings were investigated by scanning electron microscopy (SEM). For the evaluation of the bioactivity, uncoated and polymer-coated scaffolds were immersed in simulated body fluid (SBF) for up to 14 days. The formation of hydroxyapatite was confirmed by SEM observation and Fourier transform infrared spectroscopy (FTIR). Results showed no influence of the coating material on the bioactive behavior of the 45S5 bioactive glass-based scaffolds. The mechanical performance of the different scaffolds was investigated by compression strength tests. Compressive strength increased from 0.04 ± 0.02 MPa for uncoated scaffolds to 0.21 ± 0.03 MPa, and 0.18 ± 0.02 MPa for uncrosslinked and crosslinked collagen-coated samples, respectively. Zein-coating showed values of 0.21 ± 0.02 MPa and 0.19 ± 0.03 MPa for uncrosslinked and crosslinked samples, respectively. The enhancement of the mechanical performance of the different scaffolds was attributed to the polymeric infiltration of the microcracks present on the surface of the struts. The values of compression strength are still at the lower boundary for BTE applications but coated scaffolds were sufficiently robust to be handled safely which indicates that they can be used in clinical settings. The scaffolds exhibited the expected cell response when

VII

Abstract tested in vitro in contact with MG-63 cells. In addition, collagen- and zein-coated 45S5 bioactive glass-based scaffolds indicated an enhanced osteogenic differentiation behavior of ST-2 cells compared to uncoated samples. Altogether, the results of the present investigations showed that zein- and collagen-coated scaffolds have promising potential for the application in bone tissue engineering. Next to polymer-coated scaffolds, collagen and zein were also used for the fabrication of composites based on porous polymer matrices incorporating 45S5 bioactive glass particles as reinforcement. Primarily, this approach was applied to estimate the characteristics of collagen and zein as biomaterial. One the one hand, porous collagen sponges with and without bioactive glass incorporation were produced by lyophilization. In addition, salt leaching was employed to fabricate porous zein-based structures (with and without 45S5 bioactive glass). Even though collagen scaffolds exhibited high bioactivity and porosity of around 98 %, the mechanical properties were considered to be more relevant for soft tissue engineering applications. Zein scaffolds reinforced with bioactive glass exhibited compressive strength of 2.2 ± 0.9 MPa. In addition, the samples were bioactive with a porosity of around 84 %. The novel developed scaffolds provide a well- founded basis that warrants their further investigations in the field of bone tissue engineering. Composite scaffolds based on the combination of natural derived proteins and bioactive glass are therefore interesting candidates for BTE and the results of the present investigations should be expanded in future studies in relevant in vivo models.

VIII

Zusammenfassung

Die vorliegende Arbeit beschreibt und erläutert die Entwicklung und Charakterisierung neuartiger Kompositmaterialien auf der Grundlage von bioaktivem Glas und natürlichen Polymeren, die im Knochen Tissue Engineering Anwendung finden sollen. Die steigende Nachfrage nach Möglichkeiten, beschädigtes Körpergewebe zu ersetzen, ist nicht nur auf die fortschreitende Alterung der Bevölkerung zurückzuführen, sondern liegt vor allem an der eingeschränkten Fähigkeit des Körpers verlorenes Gewebe eigenständig zu regenerieren. Vor allem der Ersatz von verlorener Knochensubstanz stellt für Mediziner immer wieder eine große Herausforderung dar. Heutzutage hat sich der Einsatz von autogenen oder allogenen Knochenimplantaten als Standard etabliert. Allerdings birgt die Anwendung auch viele Nachteile. Insbesondere die eingeschränkte Verfügbarkeit, die Möglichkeit von Infektionen oder auch auftretende Unverträglichkeiten müssen hierbei genannt werden. Ebenso spielen ethische Bedenken eine Rolle. Um diese Problematik umgehen zu können, konzentriert sich Knochen Tissue Engineering auf die Entwicklung sogenannter Scaffolds, die als Alternative für autogene bzw. allogene Knochentransplantate dienen können. Manche Scaffolds bestehen aus anorganischen synthetischen Materialien, z.B. Hydroxylapatit oder auch Bioglass®, wobei die Anwendung von bioaktivem Glas mehr und mehr Aufmerksamkeit auf sich zieht, da die Vorteile (u.a. Bioaktivität, Osteokonduktivität oder auch die angiogene Wirkung) für sich selbst sprechen. Allerdings ist das Einsatzgebiet von bioaktivem Glas in Form von Scaffolds aufgrund der niedrigen mechanischen Festigkeiten begrenzt, insbesondere in lasttragenden Bereichen. Demzufolge beschäftigt sich die Forschung mit der Entwicklung von Kompositmaterialien, um Eigenschaften unterschiedlicher Werkstoffgruppen zu vereinen. Das Ziel dieser Arbeit war unter anderem die Verbesserung der mechanischen Eigenschaften von bioaktivem Glas (mit der Zusammensetzung 45S5) basierten Scaffolds (die mit Hilfe der Schaumreplikationsmethode hergestellt wurden), indem diese mit Kollagen oder Zein beschichtet wurden. Für die Kollagenbeschichtung wurde die Oberfläche der Proben funktionalisiert, um Kollagen durch Peptidbindungen an die Oberfläche zu koppeln. Die Zeinschicht wurde durch einfache Tauchbeschichtung in eine Lösung mit 8 Gew.% Zein aufgetragen. Sowohl Kollagen wie auch Zein wurden des Weiteren durch chemisch induzierte Quervernetzungen stabilisiert. Neben der Mikrostruktur der Scaffolds und der Homogenität der Beschichtung, die mit Hilfe von Rasterelektronenmikroskopie (REM) untersucht wurden, spielt auch die Bioaktivität eine große Rolle. Um diese beurteilen zu können, wurden Studien durchgeführt, bei denen beschichtete und unbeschichtete Scaffolds bis zu 14 Tagen in simulierter Körperflüssigkeit (SBF) ausgelagert wurden. Die Abscheidung von Hydroxylapatit wurde sowohl durch REM Aufnahmen als auch durch Fourier-Transformations-Infrarot-Spektroskopie (FTIR) bestätigt. Es konnte gezeigt werden, dass die Bioaktivität durch die Beschichtung mit Kollagen oder Zein nicht beeinträchtigt wurde. Die mechanischen Festigkeiten der unterschiedlichen Proben wurden mittels Druckversuchen bestimmt. Die Ergebnisse zeigten eine leichte Erhöhung der Druckfestigkeit von 0.04 ± 0.02 MPa für unbeschichtete Scaffolds auf 0.21 ± 0.03 MPa (unvernetzt) und 0.18 ± 0.02 MPa (vernetzt) für kollagen-

Zusammenfassung beschichtete Scaffolds. Zein-beschichtete Scaffolds zeigten Werte im Bereich von 0.21 ± 0.02 MPa (unvernetzt) und 0.19 ± 0.03 MPa (vernetzt). Die erhöhten Werte können hauptsächlich auf die Infiltration des jeweiligen Polymers in die Bruchstellen an der Oberfläche der einzelnen Scaffoldstege zurückgeführt werden. Die gemessenen Werte liegen an der unteren Grenze für die Anwendung im Knochen Tissue Engineering, zeigen aber trotzdem eine ausreichende Stabilität auf, um im klinischen Bereich handgehabt zu werden. Beschichtete Scaffolds zeigten in vitro nicht nur gute Ergebnisse bezüglich Zellvitalität (MG-63 Zellen) auf, sondern lassen auch eine verbesserte osteogene Differenzierung von ST-2 Zellen im Vergleich zu unbeschichteten Scaffolds vermuten. Die vorliegenden Ergebnisse von kollagen- und zeinbeschichteten Scaffolds liefern also vielversprechendes Potential für die Anwendung im Knochen Tissue Engineering. Neben Polymerbeschichtungen wurden Kollagen und Zein auch verwendet, um poröse Polymermatrizen herzustellen, die durch die Einlagerung von bioaktiven Glaspartikeln verstärkt werden sollten. In erster Linie diente dieser Ansatz der Charakterisierung von Kollagen und Zein als Biomaterial. Einerseits wurden poröse Kollagenschwämme durch das Verfahren der Gefriertrocknung hergestellt. Zusätzlich diente die Salzauslaugung dazu, poröse zeinbasierte Strukturen zu erschaffen (mit und ohne bioaktivem Glas). Obwohl Kollagenschwämmchen eine hohe Bioaktivität und Porosität um die 98 % aufwiesen, sind sie aufgrund ihrer geringen mechanischen Festigkeit wesentlich besser für Tissue Engineering von Weichgewebe geeignet. Hingegen zeigten Zeinscaffolds, die mit bioaktivem Glasteilchen verstärkt wurden, Druckfestigkeiten im Bereich von 2.2 ± 0.9 MPa auf, bei einer Porosität um die 84 %. Zusätzlich konnte auch bioaktives Verhalten nachgewiesen werden. Die Ergebnisse dieser neuartig entwickelten porösen Strukturen liefern daher eine fundierte Grundlage für weitere Forschungen auf diesem Gebiet. Kompositmaterialien basierend auf der Kombination von natürlichen Proteinen und bioaktivem Glas bieten interessante Möglichkeiten für zukünftige Anwendungen in der regenerativen Medizin. Zusätzlich rechtfertigen die hier gezeigten Ergebnisse weitere Studien in relevanten in vivo Modellen.

Introduction

1 Introduction

The demographic evolution shows a worldwide progressive population aging over the past centuries. This phenomenon can be assigned to the improved medical care and high standards of hygiene, achieved especially in the developed countries. However, this development not only influences the economic situation but also entails a high demand for the healthcare system, in particular the availability of “body parts” to replace damaged and non-functional tissue and organs. The replacement or repair of tissue damaged by disease or trauma is a growing challenge faced by thousands of surgeons every day. Transplants can be divided into four categories1: (I) autografts, (II) allografts, (III) isografts and (IV) xenografts. Autografts are cells, tissue or organs which are taken from the same individual. Examples are transplantations of hair roots or bone parts (usually iliac crest1), autologous blood donation, vein extractions, skin which is used to cover surficial defects among others. Allografts are transplants taken from genetically non-identical donors of the same species. Examples for allotransplantations are the transplantations of whole organs, e.g. heart, liver or kidney. But also stem cells or bone parts can be transferred. A special type of allotransplantation is the so-called composite tissue allotransplantation2,3 where whole body parts (like extremities) are replanted. To avoid a rejection reaction of the body because of the genetic differences, immunosuppressant agents are administered. An exception to this are isografts which are taken from the identical twin. Xenografts are transplanted from one species to another. A common example for this is the porcine heart valve transplant. The different types of transplants show both advantages and drawbacks. Even though autografts are currently the gold standard4–6, the availability is strongly limited. Allografts as well as xenografts carry the risk of infection, rejection and incompatibility. Next to this, also ethical concerns are a significant issue when applying xenografts. Therefore, a solution approach is based on the development and application of alloplastic grafts which consist of non-biological engineered materials like metals, ceramics or polymers. A main field for the application of alloplastic grafts is bone tissue engineering. In general, tissue engineering combines approaches from natural and engineering sciences to develop artificial and three-dimensional constructs which support or replace damaged tissues or organs1,7,8. Bone tissue engineering (BTE) focuses on the regeneration process of native bone tissue9–11.

Figure 1: Schematic diagram showing the different steps in scaffold-guided bone tissue engineering, adapted from Ref.12.

Introduction

Bone is a complex, vascularized tissue indicating that BTE is an interdisciplinary field based on the combination of three different components, as represented in Figure 1: cells, signaling molecules and a scaffold which acts as temporary matrix to support the attachment, growth and proliferation of cells. As allografts and xenografts show disadvantages like limited availability, infections or ethical concerns, BTE usually involves an alloplastic graft as scaffold13,14. The scaffold mimics the extracellular matrix and should offer suitable mechanical properties, adequate pore size and porosity to enable tissue ingrowth and vascularization15. Ideally, the applied material should possess not only osteoconductivity but also osteoinductivity and it should be biodegradable. A wide range of biomaterials is available for the application as bone scaffold material. As part of natural bone, hydroxyapatite is an appropriate material and is widely used in this field16–21. Also tricalcium phosphate, which is similar to hydroxyapatite in terms of chemical composition, is used very often22–25. In this context, bioactive glasses are increasingly attracting attention based on their high bioactivity, osteogenic potential, biodegradability and angiogenic effects12,26–28. However, bioactive glasses are naturally very brittle and thus scaffolds fabricated from bioactive glasses are usually not suitable for load-bearing applications29. To overcome this problem, composite materials are developed to combine different material characteristics30–32, in this context, bioactive glass-polymer combinations should exhibit improved mechanical properties. In the framework of this thesis, bioactive glass (with the 45S5

33 composition: 45 % SiO2, 24.5 % CaO, 24.5 % Na2O and 6 % P2O5 in wt.%) was combined with two natural polymers, namely collagen and zein, forming novel composite materials. Collagen was used as it is the most abundant protein inside the human body34. However, its extraction is not only time-consuming but also very expensive. Especially, the high costs are a big drawback for future clinical applications. Because of this, zein, a plant-derived protein from corn35, was used to investigate its capability to replace collagen. The potential of the new developed composite scaffolds was evaluated in terms of mechanical properties, bioactivity, biodegradability and cell biocompatibility for bone tissue engineering applications. The thesis is organized in the following manner. After a summary of fundamentals in chapter 2, describing the application of scaffolds in bone tissue engineering, manufacturing techniques, results and discussion of composite scaffolds with collagen and bioactive glass are presented in chapter 3. Chapter 3 is divided into two sections (3.1 and 3.2), which focuses on the reinforcement of porous collagen matrices with bioactive glass particles and collagen as coating material on bioactive glass-based scaffolds, respectively. Chapter 4 is organized likewise but highlights the application of zein. In vitro studies, which were carried out in static and dynamic cultivation systems, are presented in chapter 5. This chapter also contains the comparison of the investigated systems. The thesis is completed by the conclusion and discussion of future work in chapter 6. Figure 2 shows a graphical representation of the research carried out in this thesis, showing the materials investigated, composite systems developed and characterization conducted.

Introduction

Figure 2: Graphical representation of the research carried out in the framework of this doctoral thesis.

Fundamentals

2 Fundamentals Application of scaffolds in bone tissue engineering

2.1 State of the art

The history of tissue engineering can be traced way back to the antiquity36,37. Since then, the mankind imagines an improved way of life by the replacement or repair of damaged tissue. Fra Angelico, an Early Italian Renaissance painter, depicted the twin brothers Cosmas and Damian (died in the 3rd century) in a famous painting entitled “The healing of Justinian” showing the transplantation of an allograft limb. Until today, medicine has made great progress in realizing this imagination. The interdisciplinary field of tissue engineering “...follows the principles of cell transplantation, materials science and engineering towards the development of biological substitutes that can restore and maintain normal function”7. However, many drawbacks are left before the vision of tissue replacement can be realized entirely. Especially, bone tissue engineering is a growing challenge38. The present chapter not only explains the need for bone tissue engineering but also summarizes the current state of biomaterials for the application in the field of BTE38,39 and introduces different manufacturing techniques for biomedical scaffolds.

2.1.1 Bone tissue engineering

In general, bone plays an essential role and performs different tasks inside the body. First of all, it supports the whole human body and provides a lever system for muscle activity1. Furthermore, the skeleton protects internal organs. Next to the structural function, bone also acts as metabolic reservoir as it stores calcium, phosphate, magnesium and other trace elements1,40. Bone tissue consists of osteocytes embedded in the bone matrix which contains 30 wt.% of organic matter (mainly collagen I) and 70 wt.% of inorganic components (mainly bone apatite, a carbonate-rich form of hydroxyapatite)1,40. In addition, osteoblasts, the bone forming cells, produce organic bone matrix whereas osteoclasts resorb bone tissue by acidification. Bone tissue is highly vascularized which ensures the nutrient and oxygen supply of bone cells. In contrast to other types of tissue, bone exhibits the unique character of regeneration without scar tissue41,42. Bone healing is a very complex process and can be divided into direct and indirect fracture healing40 (Figure 3). It is influenced by many factors like type of fracture and bone, blood supply, age of patient, injury severity among others. The direct healing can be differentiated between contact and gap healing. If there is a close contact between the fracture parts (up to 200 µm), contact healing can take place where osteons bridge the fracture gap and enable a direct growth of osteoblasts, osteoclasts and vessels. If the gap is between 200 µm and 1 mm, woven bone is formed and gradually transformed into lamellar bone. The indirect fracture healing is implemented into several stages (Figure 3). Simplified, during stage 1 a fracture hematoma is developed in between the bone fragments. This is followed by an inflammation reaction (stage 2) where macrophages and granulocytes remove necrotic tissue. Fibroblasts start producing collagen (stage 3) and

Fundamentals granulation tissue is formed which harden in stage 4 (callus). Under the influence of bone morphogenetic proteins (BMPs), osteoprogenitor cells are differentiated into osteoblasts which induce the mineralization of the callus and the transformation to woven bone. During the remodeling in stage 5, woven bone is converted to lamellar bone and the original bone structure is restored40.

Bone fracture

Indirect healing Direct healing 1. Injury (fracture hematoma)  Contact healing (up to 200 µm) 2. Inflammation  Gap healing (200 µm – 1 mm)_ 3. Granulation 4. Callus formation and osteogenesis 5. Remodeling

Figure 3: Schematically overview of the bone healing process.

However, the ability of bone regeneration is limited. The non-regenerative threshold of bone is called “critical-sized defect”43 and is defined as “the smallest size intraosseous wound in a particular bone and species of animal that will not heal spontaneously during the lifetime of the animal”43,44. Therefore, the treatment of larger bone defects caused by trauma or disease requires the application of bone grafts which can be derived from different origins. Nowadays, the gold standard for bone reconstruction are autografts (taken from the same individual) which exhibit a high osteogenic potential and excellent biocompatibility. But limited availability, pain and donor site morbidity are only a few of the know side effects and significant drawbacks of autografts38,39. Alternatives like allografts (taken from another human donor) or xenografts (of animal origin) are associated with risks of infections, incompatibility, ethical concerns among others. Hence, great hopes have been set on the research and development of bone tissue engineering technologies which are based on the combination of a suitable scaffold structure (made of engineered biomaterials) with appropriate growth factors and cells. An ideal scaffold should provide a temporary matrix for the growth and proliferation of bone cells with osteogenic characteristics like osteoconductivity and osteoinductivity, sufficient mechanical strength, bioactivity, high porosity to support tissue ingrowth and vascularization, biocompatibility and biodegradability. As autografts, allograft and xenografts are associated with limitations and drawbacks (mentioned above), new approaches focus on the design and development of artificial three- dimensional scaffolds made of biomaterials. Biomaterials for bone tissue substitute must be neither toxic nor

Fundamentals carcinogenic, corrosion resistant, biodegradable (at least partially), sterilizable, bioactive, economical and storable1,10,38,39,45–48. Many materials are available but only some of them can be considered as biomaterial.

Table 1: Overview of popular biomaterials available for applications in bone tissue engineering.

Material type Material Reference Chen et al.29 Gerhardt et al.49 Bioactive glasses Jones et al.50 Ceramics and bioactive glasses Fu et al.51 Oliveira et al.21 HA Deville et al.19 Venkatesan et al.52 Venkatesan et al.53 Alginate Valente et al.54 Augst et al.55 Natural derived polymers Li et al.56 Chitosan Levengood et al.57 Venkatesan et al.58 Salgado et al.59 Starch Martins et al.60 Williams et al.61 PCL Wong et al.62 Synthetic polymers Salerno et al.63 PLA Pan et al.64 Witte et al.65 Magnesium Liu et al.66 Balla et al.67 Metals Tantalum Rupérez et al.68 Zou et al.69 Dabrowski et al.70 Titanium Wang et al.71 Venkatesan et al.52 Valente et al.54 Alginate composites Lin et al.72 Castilho et al.73 Venkatesan et al.58 Composites Chitosan composites Zhang et al.74 Chesnutt et al.75 Collagen + bioactive glasses Sarker et al.76 PCL + HA Morelli et al.77 PGA + β-TCP Cao et al.78

Calcium phosphates79,80, like hydroxyapatite (HA), biphasic calcium phosphate (BCP) or beta-tricalcium phosphate (β-TCP) are very interesting candidates for the application in bone tissue engineering, as they offer similar composition compared to the inorganic phase of bone. Also bioactive glasses of different compositions have great potential due to their highly bioactive behavior50. However, these materials are

Fundamentals naturally very brittle and not suitable for load-bearing applications whereas polymers are very flexible and provide chemical versatility, being, however, in general, non-bioactive. Polymers can be classified into synthetic and natural origin. Natural polymers like collagen, alginate, silk, starch or chitosan are commonly used as biomaterial. Examples for synthetic polymers, applied in the biomedical field, are polycaprolactone (PCL), polylactic acid (PLA) or polyglycolic acid (PGA). Also metals have been proposed as bone substitute materials but compared to ceramics or polymers the lack of degradability is a big drawback (except for recent developments based on Mg alloys65,81). To combine key characteristics of different material classes, composite materials have been proposed since the early day of bone tissue engineering31,74,82. In Table 1, a selection of materials is shown which includes currently available popular biomaterials for bone tissue engineering applications10,38,39. This variety of biomaterials implies also the availability of many different fabrication techniques. Most relevant fabrications methods for three-dimensional scaffolds are presented in the following section.

2.1.2 Scaffolds and their fabrication technologies

In the last years various techniques have been introduced to fabricate three-dimensional and porous scaffold structures for tissue engineering applications, e.g. additive manufacturing83–88 (previously named rapid prototyping), solvent casting and particulate leaching89–92, gas foaming93–95, electrospinning96–99, freeze drying100–104 or phase separation105–107.

The concept of additive manufacturing summarizes different fabrication techniques which combine computer-assisted design (CAD) with computer-assisted manufacturing (CAM). The scaffolds are produced layer by layer. Different methods are available like selective laser sintering (SLS), stereolithography or 3D printing. Williams et al.61 describe the fabrication process of porous PCL scaffolds by selective laser sintering which is a fast and cost effective method108. Therefore, PCL particles were successively sintered inside a powder bed by laser scanning where particles are fused but not decomposed. By this method very complex structures can be produced. However, the system has drawbacks like oxidation of the polymer or material shrinkage108. Stereolithography uses an ultraviolet laser which hardens a photosensitive polymer.

Vozzi et al.109 applied this technique to produce poly(DL-lactide-co-glycolide) scaffolds by a poly(dimethylsiloxane) mold. Nevertheless, only a few photosensitive polymers suitable as biomaterial are available which limits this technique for tissue engineering applications. In contrast, 3D printing is the most commonly used rapid prototyping technology for the fabrication of both, polymer and bioceramic scaffolds108. For example, Cox et al.20 used this method to produce porous hydroxyapatite (HA) structures by combining HA with a binder and by subsequent sintering. However, this method can produce simple porous structure without high complexity.

Fundamentals

Solvent casting combined with particulate leaching is a very common and simple procedure to prepare polymeric scaffolds or structures with desired porosity. A polymer is dissolved in an organic solvent and a porogen (e.g. sodium chloride particles) is added to the solution. By the size and ratio of the porogen, the porosity is defined. Afterwards, the solution is casted. The final geometry can be a thin membrane (cast on a petri dish) or a scaffold (cast into a three-dimensional mold). The solvent can be removed by freeze drying or evaporation. To receive a porous structure, the porogen is leached out. Thadavirul et al.90 as well as

Liao et al.91 used this technique to produce porous scaffolds of PCL and poly(L-lactic-co-glycolic acid) (PLGA), respectively. PCL was mixed with polyethylene glycol (PEG) and dissolved in chloroform. In addition, sodium chloride particles were added. After the evaporation of the solvent, the scaffolds were leached in deionized water and dried. The scaffolds exhibited a highly interconnected porous structure. For the fabrication of PLGA scaffolds, PLGA was mixed with sodium chloride and dissolved in different solvents (1,4-dioxane, acetone, chloroform, dichloromethane and tetrahydrofuran (THF)). After the solvent removal, sodium chloride was leached in distilled water. Samples were dried and revealed highly porous three-dimensional structures.

Polymeric porous structures can also be produced by gas foaming. Several gases like carbon dioxide (CO2),

110 111 nitrogen (N2) or helium (He) are available as porogen among others . Sheridan et al. describes the fabrication of porous poly(lactic-co-glycolic acid) scaffolds by exposure to high-pressure CO2. The solid polymer was compressed into disks and saturated with the gas. A gradually reduction of the gas pressure induced an expansion of the gas phase which resulted in a porous structure. However, this method only generates scaffolds with a closed porosity and low pore interconnectivity which limits their application in the area of tissue engineering110. To overcome this problem, gas foaming is often combined with particulate leaching which increases the degree of interconnected porosity111,112. As this technique does not require the use of solvents, bioactive molecules can be incorporated into the polymer matrix without altering their biological function.

Fibrous nanostructured scaffolds can be produced by electrospinning. This technique produces fibers which are assembled in a non-woven mesh96. During the process of electrospinning, a polymer solution is charged in a high electric field. Due to electrostatic repulsion, a fiber is formed which is collected on a fixed grounded collector (resulting in randomly orientated fibers) or a rotating grounded collector (resulting in aligned fibers)113. The diameter of the fibers is affected by many parameters like type of polymer, viscosity of the polymer solution, distance to collector or flow rate. Normally, fibers can be produced in the range of a few nanometers up to micrometers96. Li et al.114 fabricated nanofibrous scaffolds for tissue engineering applications with this technique. Poly(L-lactic acid) (PLLA) was dissolved in a mixture of chloroform and N,N-Dimethylformamide (DMF). PCL was dissolved in THF and DMF. The polymer solutions were electrospun at 16 and 12 kV, respectively, and collected on a sheet of alumina foil. Results revealed fibrous

Fundamentals matrices with high ratio of surface area to volume, mimicking the structure of the native extracellular matrix.

For the fabrication of porous structures by phase separation, polymers are dissolved in a solvent. The inducement of liquid-liquid or solid-liquid phase separation can be triggered by the influence of chemicals (nonsolvent induced phase separation, NIPS) or due to thermal changes (thermally induced phase separation, TIPS)108,110,115. During liquid-liquid phase separation the polymer solution is divided into a polymer-rich and a polymer-poor phase. Solid-liquid phase separation appears when the polymer crystallizes from the solution110. After the removal of the solvent, porous polymeric scaffold are formed. Liu et al.105 produced porous nanofibrous gelatin scaffolds by this method. Gelatin was dissolved in a mixture of ethanol and water. The phase separation was induced at -76 °C. By the combination with porogen leaching, highly porous structures with high interconnectivity were successfully produced.

Depending on the desired application, a suitable fabrication technique has to be chosen. During this study, porous structures were produced by freeze drying, salt leaching and by the foam replica method. Information about the applied methods in this project is provided in the corresponding section.

2.2 Materials selection

As discussed previously, a wide range of materials can be selected to produce scaffolds for the application in bone regenerative medicine. The selection of materials, which were utilized in the framework of this thesis, was made based on their properties and is explained on the following pages.

2.2.1 45S5 Bioglass® – a silicate-based bioactive glass

In general, a glass is defined as non-crystalline, brittle and inorganic material. Its state can be characterized as frozen, supercooled liquid with amorphous structure and isotropic behavior. Specific metallic and non- metallic oxides build up a three-dimensional structure. So-called network formers provide the skeletal

116 117 structure, e.g. SiO2, B2O3 or P2O5. Based on the theory of Zachariasen and Warren , the structural unit of silicate glass consists of SiO4-tetrahedra which exhibit a distorted structure due to irregular distances and bond angles. Compared to this, a quartz crystal builds a regular structure with short and long range order

(Figure 4). The network modifiers (e.g. CaO, K2O or Na2O), as the name already implies, modify the

28 network, change its structure and properties by inclusions into the structure of the SiO4-tetrahedra (Figure 4). Therefore, the –Si–O–Si– bridging oxygen bonds are disrupted and instead of binding to Si, O is forming ionic bonds to the present alkali cations (e.g. Ca2+, K+ or Na+).

Fundamentals

Figure 4: Two-dimensional structure model of (I) quartz crystal, (II) silicate glass and (III) alkali-silicate glass with disrupted network118 (reprinted with permission from Hanser - Copyright © 2008 Carl Hanser Verlag München).

Predominantly, the properties of a glass depend on the composition. Bioactive glasses in general show the ability for interfacial bonding between the material and tissue by a specific biological response which results in the formation of a biologically active hydroxyapatite layer on the surface of the implant119,120. Nowadays, there are several types of bioactive glasses like phosphate- and borate-based glasses or conventional silicates, such as the well-known 45S5 Bioglass® on which this section is focused on.

Production, structure and properties

The story of 45S5 Bioglass® started in the late 60s of the last century, more specifically in 1969, and was the result of investigations led by Larry L. Hench (at University of Florida, USA)33. The glass with the composition 45 % SiO2, 24.5 % CaO, 24.5 % Na2O and 6 % P2O5 (in wt.%) is used for clinical applications since 1985121 and is produced by a conventional melt-derived synthesis route. In this method, silica, calcium oxide, sodium oxide and phosphorus pentoxide are homogenously mixed and melted in a platinum crucible at around 1400 °C. To obtain a solid glass block, the glass melt is casted into thermostable molds made of graphite. For the production of Bioglass® powder, the liquid glass is quenched in water, resulting in so- called glass frit which can be milled down to a desired particle size using a planetary mill. Originally, the composition of Bioglass® was chosen as it is close to a ternary eutectic making it easy to melt33,122. Compared to the composition of conventional soda-lima glass, Bioglass® exhibits three main differences123:

(I) lower amount of SiO2 (< 60 mol.%), (II) high content of CaO and Na2O and (III) high ratio of calcium to phosphate. Figure 5 shows the ternary phase diagram of SiO2-CaO-Na2O on which the composition of

45S5 Bioglass® is based. Thereby, SiO2 acts as network former. The structure of SiO4-tetrahedra is disrupted by the presence of Ca and Na ions which act as network modifier resulting in non-bridging oxygen (NBO) inside the glass structure. The high ratio of calcium to phosphate promotes the high bioactivity. The structure of the glass can be assessed by the glass composition using the concept of network connectivity (NC)121 which describes the number of bridging oxygen per network forming element . The

Fundamentals network connectivity can be calculated in consideration of the molar ratio of each component. In case of

121 Bioglass®, Si is the network forming element which exhibits a network connectivity NCSi of 2.11 . The special feature of 45S5 Bioglass® is its ability to bond to soft and hard tissue which can be attributed to its composition (Figure 5). This phenomenon is called bioactivity which describes the ability of a material to interact with living tissue resulting in a strong implant-bone interface. Biomaterials can be divided into three categories based on their biological behavior, namely bioinert, bioactive and resorbable123,124. A material is defined as bioinert if it is not toxic and not interacting with the surrounding tissue but encapsulated by fibrous tissue. Examples for bioinert ceramics are Al2O3 or ZrO2. If a material is completely dissolved and replaced by new tissue, it is resorbable (e.g. β-TCP). As already mentioned, Bioglass® is classified as bioactive material. The level of bioactivity is related to the time taken in days (t0.5) to bond more than 50 % of the

124 interface to bone . The bioactivity index IB can be calculated as

IB = 100/t0.5 [Eq. I]

Materials, which exhibit IB values > 8, are class A bioactive materials (like Bioglass®) and can bind to soft

124 and hard tissue whereas class B materials show IB below 8 but > 0 and can only bind to hard tissue .

Figure 5: Compositional diagram for bone-bonding. Note regions A, B, C, D. Region S is a region of Class A bioactivity where bioactive glasses bond to both bone and soft tissues and are gene activating33 (reprinted with permission from Journal of Materials Science: Materials in Medicine - Copyright © 2006 Springer).

As reported in the early literature about Bioglass®, the bone-bonding mechanism is enabled by the formation of a hydroxyapatite layer on the material’s surface and is divided into eleven stages124. Stages I-V are chemical reactions which can also be mimicked in vitro by the immersion in simulated body fluid (SBF). Stages VI-XI require implantation into the human body as cells and biological entities of the extracellular matrix are involved in the processes124,125. In contact with aqueous solutions (like body fluids), the Bioglass® network starts to dissolve. Chemical and structural changes occur as function of time124. Due to non-

Fundamentals bridging oxygen, the silica network in Bioglass® is less crosslinked compared to conventional glass leading to a more open structure which facilitates the penetration of water molecules into the glass network121. Stages I and II of the bioactivity mechanism are diffusion controlled where Na+ and Ca2+ ions, which are part of

+ + the Bioglass® structure, are rapidly exchanged by H or H3O from the aqueous solution. The silica groups are hydrolyzed, forming silanol (Si-OH) groups on the glass surface. During stage III, polycondensation of the Si-OH groups takes place which results in the formation of a silica-rich layer (SiO2) on the surface. In

2+ 3- stage IV, Ca and PO4 migrate to the silica-rich layer, forming an amorphous calcium phosphate layer

2+ 3- - 2- which is crystallized during stage V due to the incorporation of soluble Ca , PO4 , OH and CO3 ions from the solution resulting in a carbonated hydroxyapatite layer (HCA), structurally similar to the mineral phase in bone. The biological process of bone-bonding starts in stage VI when biological entities are adsorbed. Macrophages are activated (stage VII) and stem cells attach to the bioactive surface (stage VIII). In stage IX, stem cells differentiate into osteoblasts which start to generate the organic bone matrix (stage X). During the final stage XI, the organic matrix is mineralized by the incorporation of calcium phosphate, forming a natural bone structure. Furthermore, dissolution products of 45S5 Bioglass® are reported to stimulate osteogenesis and angiogenesis by the enhanced expression of genes in cells leading to bone formation and vascularization126. The dissolution products are thus key factors influencing the differentiation and proliferation behavior of osteoblastic cells and enhancing the secretion of relevant growth factors, e.g. VEGF (Vascular endothelial growth factor) to promote angiogenesis126.

45S5 Bioglass® as scaffold material

For the first time in 2006, Chen et al.29 produced 45S5 Bioglass®-based scaffolds with a highly interconnected porous structure by the foam replication technique. As Bioglass® offers excellent bioactivity, high biocompatibility and controllable biodegradability, it is a promising scaffold material for tissue engineering applications. The foam replica technique requires a sacrificial template (e.g. polyurethane foam) which is coated by a Bioglass® slurry resulting in porous green bodies. By controlling the dimensions of the template structure, the final scaffold geometry can be influenced. During a subsequent sintering process (two stages, 400 °C and 1050 °C), the template material is burned out and the bioactive glass microstructure is densified. The thermal treatment of 45S5 Bioglass® powder entails a phase

127,128 transformation showing the nucleation and growth of a crystalline phase (Na2Ca2Si3O9) which not only influences the densification process but may also affect the bioactive behavior29,127. However, the crystallinity can be controlled by the sintering temperature and the heating rate. Furthermore, as 45S5 Bioglass®-based scaffolds offer a high surface area due to the high porosity, bioactive kinetics can differ to the one of a bulk

29 and dense material. Chen et al. described the transition of Na2Ca2Si3O9 to a bioactive and biodegradable amorphous phase after the immersion in SBF for 28 days. In addition, the mechanically strong crystalline phase offers temporary mechanical support. Therefore, 45S5 Bioglass®-based scaffolds have the advantage

Fundamentals of combining sufficient mechanical competence, bioactivity and tailorable biodegradability for further investigations. However, the mechanical properties are not suitable for load-bearing applications.

2.2.2 Collagen – a natural polymer

Collagen, a natural structural protein, is the predominant component in connective tissue and therefore the most abundant protein in mammals129. Collagen forms a structural network providing strength and stability in various tissues. Up to now, 28 different types of collagen have been identified130. This thesis is focused on the characterization and application of collagen type I (henceforth referred to as collagen), as it is the main organic component in the extracellular matrix of bone tissue.

Structure of collagen

Collagen fibers consist of collagen fibrils which are partially synthesized intracellular in the ribosomes of the endoplasmic reticulum (ER). Inside the cells procollagen is formed by three polypeptide chains (known as α chains) which exhibit a left-handed triple helix.

The primary structure is characterized by a repeating peptide sequence, namely Glycin-X-Y (where X and Y represent any amino acid but frequently proline and hydroxyproline)131. Two identical α(1) chains and one

Fundamentals

α(2) chain are assembled intracellular by hydrogen bonds forming a typical right-handed superhelix (procollagen). The procollagen is released into the extracellular matrix and is around 300 nm long and 1.5 nm wide. By the enzymatic cleavage of the C- and N-terminal propetide, the molecule is converted into collagen. Collagen fibrils are formed by the alignment of collagen molecules with a gap of 40 nm and a displacement of 67 nm34. A multitude of collagen fibrils form collagen fibers, stabilized by inter- and intramolecular crosslinks. The collagen structure is schematically represented in Figure 6. For the production of collagen solutions, collagen can be extracted from rat tail tendon, bovine Achilles tendons or bovine skin by acids (hydrochloric acid or acetic acid)132. Collagen fibrils are partially solubilized resulting in a solution with collagen triple helices. In addition, terminal propeptides can be removed by the use of pepsin which reduces the immunogenicity of collagen132. So-called atelocollagen is less viscous and has a better yield of extractions133. However, due to the missing propeptides, the forming of collagen fibrils is very difficult as such biomolecules play a crucial role in the process of fibrillogenesis132. On the other hand, collagen fibrils received by acidic treatment are able to self-assemble depending on the pH and temperature but lack mechanical strength due to the lack of crosslinks132.

Collagen for tissue engineering

As already mentioned, collagen is the most abundant protein in the extracellular matrix and is therefore a very popular biomaterial. Due to its natural origin, collagen offers many binding sites for cellular attachment. Cells can bind through fibronectin or directly attach to the RGB sequence in the collagen network via integrin receptors134. Lee et al.135 provide the following comparison of advantages and disadvantages of collagen as biomaterial (Table 2).

Table 2: Advantages and disadvantages of collagen applied as biomaterial135.

Advantages Disadvantages

 High availability of bovine skin or tendons  High cost of pure collagen from which collagen can be easily extracted  Variable properties of extracted collagen  Non-antigenic  Hydrophilicity which leads to swelling  Biocompatible and non-toxic  Complex handling properties  Biodegradable and resorbable  Biodegradability can be controlled by the degree of crosslinking  Hemostatic  Collagen fibers offers high tensile strength  Compatible with additives

Fundamentals

 Chemical versatility  Many binding sites for cell attachment

Collagen can be processed into different shapes like sheets, sponges or injectable scaffolds among others136,137 which have different applications in the biomedical field. The main application of collagen films or sheets is located in the field of ophthalmology. They are used as graft material for corneal replacement or as treatment for infected tissue when loaded with antibiotics137,138. Furthermore, collagen films can be combined with bone morphogenetic proteins (BMPs) to achieve direct osteoinduction139. Three- dimensional collagen sponges are often applied as wound dressing as they can absorb exudates and allow a smooth adherence offering a barrier for bacterial infections135. Coatings of collagen sponges with growth factors showed improved healing of dermal and epidermal wounds135. In addition, collagen has been also introduced as injectable scaffold material. Collagen has been combined with polyhydroxyethyl methacrylate and the received hydrogel was loaded with anti-cancer drugs and implanted subcutaneously into rats. Results revealed a stable hydrogel structure without calcification after 6 months of implantation140. Despite attractive biological properties, collagen offers a high affinity to water and therefore very low mechanical properties, which limits its application in the field of bone tissue engineering. To overcome this drawback, collagen can be combined with different additives141–143, including inorganic phases. Sarker et al.76 have reviewed the combination of collagen with bioactive glasses. By the reinforcement with bioactive glass particles, the compressive strength and stiffness of collagen-based materials can be enhanced. Pek et al.144 introduced a nanocomposite material based on collagen and apatite which exhibited compressive stiffness of 37.3 ± 2.2 MPa. In consideration of results reported in the literature so far, collagen offers an extensively interdisciplinary field for further investigations as bone substitute material and it was considered in the present study for developing bioactive glass-collagen composites.

2.2.3 Zein – a protein derived from corn

Zein is a natural protein derived from corn. As it is extracted from renewable resources, it is very interesting to investigate its potential for biomedical applications145. Next to this, zein has shown to be biodegradable and biocompatible145,146. Zein is the major storage protein in the seeds of maize and comprises between 40 and 50 % of the total endosperm protein147. In this section, fundamentals about structure and morphology are explained. Furthermore, the use of zein as biomaterial is described.

Structure and morphology

Based on their solubility, corn proteins can be divided into four different categories: (I) albumins (water- soluble), (II) globulins (aqueous saline-soluble), (III) prolamins (water-insoluble) and (IV) glutelins (water- and alcohol-insoluble)145. The solubility of zein depends on the residues of the amino acid sequences. As zein 15

Fundamentals contains a high amount of leucine, proline and alanine, which are hydrophobic, zein is only soluble in alcohol-based solutions or organic solvents. Therefore, zein belongs to the class of prolamins145,147. Furthermore, zein is a heterogeneous mixture of different peptides (α-, β- and γ-zein) which differ in molecular size, solubility and charge. The fraction of α-zein is supposed to be around 80 % of the total prolamin amount and exhibits molecular weights of 22 and 24 kD35,145. β-zein represents between 10 and 15 % of the zein structure and is suggested to be a high molecular protein (14, 22 and 24 kD) from α-zein linked by disulfide bonds35. γ-zein comprises between 5 and 10 % depending on the genotype of the corn145. For α-zein different structural models have been proposed145,148,149.

Structural model proposed by Matsushima et al.150: Structural model proposed by Argos et al.151:

Glutamine-rich bridges

Side view Top view

Matsushima et al.150 investigated the structure of α-zein in 70 vol.% of ethanol by small-angle X-ray scattering (SAXS). They proposed the structural model represented in Figure 7. Successive helical segments are arranged in an anti-parallel manner which are connected by hydrophilic glutamine-rich bridges and stabilized by hydrogen bonds whereas the surface of the helices is supposed to be hydrophobic. Argos et al.151 performed measurements of the structure in methanol and proposed a cylindrical arrangement of the helices, as illustrated in Figure 7. Due to the helical content, zein is expected to exhibit a globular structure in non-aqueous solutions35. The molecular structure of zein is completely amorphous. The glass transition

145 temperature (Tg) is reported to be around 165 °C but can be reduced by the addition of a plasticizer. Up to 280 °C the structure of zein is thermally stable resulting in a single step degradation beyond this value.

Zein as biomaterial

As zein can form glossy, hydrophobic grease-proof coatings, which resist microbial attacks, it is widely used as coating material in the pharmaceutical and food industry35,147. Furthermore, the application of zein-based materials in the biomedical field is increasingly attracting attention due to its biodegradability and biocompatibility. Mainly, the fabrication of zein-based films is reported in the literature. Dong et al.146

Fundamentals describe the characterization of zein films which were prepared for culturing cells. Results revealed the successful attachment of human liver cells indicating the biocompatibility of zein films. Similar results were reported by Sun et al.152. Zein fibers can be fabricated by electrospinning (nanofibers) or melt-spinning (conventional fibers)153. However, morphological instability of the fibers in aqueous media combined with poor mechanical strength limits their application field. In contrast, the use of zein as controlled drug delivery system is quite promising. Drugs are entrapped inside the zein structure by stable protein-drug complexes based on the high amount of nonpolar amino acids153. Matsuda et al.154 describe the use of zein-based microspheres as vehicle for chemoimmunotherapy. Also the delivery of antibiotics has been reported elsewhere155. The protein derived from corn is not only applied for the fabrication of films, fibers or microspheres but can also be used for the development of three-dimensional scaffolds which are interesting for applications as bone substitute material. By the salt leaching technique, porous zein-based scaffolds can be produced, as reported by Gong et al.156. Compressive strengths in the range of 2.5 ± 1.2 MPa up to 11.8 ± 1.7 MPa were achieved. When seeded with cells, zein exhibited good biocompatibility. Moreover, in the presence of dexamethasone zein-based scaffolds are characterized as osteoconductive material. Compared to collagen, zein offers a higher availability and reduced cost, as it is derived from corn. Next to this, no immunogenic interactions have been noted in contact with zein-based materials. Zein is biocompatible and degradable and therefore it has great potential for the application in bone tissue engineering.

2.3 Objective of research

As there is a high demand of artificial bone substitute materials, this thesis focuses on the development of composite materials based on the combination of natural derived proteins and bioactive glass for the application in bone tissue engineering. Bioactive glass with the 45S5 composition was chosen as it offers high degree of bioactivity combined with biodegradability and biocompatibility. Next to this, the dissolution products of 45S5 bioactive glass are known to promote cell differentiation and angiogenesis. To enable vascularization and tissue ingrowth, scaffolds require a high porosity. 45S5 bioactive glass-based scaffolds were fabricated by the foam replica technique revealing a highly interconnected porous structure. However, these kinds of scaffolds are known to be very brittle which limits their application in load bearing systems. It was hypothesized that by the combination with natural derived proteins, mechanically improved scaffolds can be obtained with enhanced biological integrity. Therefore, natural derived proteins of different origin (animal and plant) were chosen on purpose to recognize and compare their potential for further investigations in the field of bone tissue engineering. Composites of bioactive glass-based scaffolds with collagen or zein are rarely reported in the literature. The development of novel composite materials is thus expected to expand the current knowledge in the field of biomaterials. Furthermore, porous matrices of collagen and zein were fabricated and reinforced with 45S5 bioactive glass particles. By the characterization

Fundamentals of the polymeric scaffolds a more fundamental knowledge of the interactions between the polymer and the inorganic phase was anticipated. As a result, the potential of natural derived protein matrices as bone substitute material can be characterized and estimated.

Collagen – As scaffold material

3 Collagen Composite scaffolds with bioactive glass and collagen

3.1 Reinforcement of porous collagen scaffolds with bioactive glass particles – Manufacturing techniques, results and discussion

As mentioned above, composite materials can be produced by many different methods. The present chapter reports on the manufacturing and characterization of composite scaffolds consisting of bioactive glass and collagen. The first section (3.1) presents the fabrication of porous matrices with different concentrations of collagen which were produced via lyophilization. To enhance the bioactive properties of these collagen sponges, they were reinforced with bioactive glass particles. The received collagenous and porous structures (with and without bioactive glass) were investigated in terms of microstructure, degradation behavior and change in bioactivity to characterize collagen as biomaterial. Results of this part of the study (3.1) were then transferred to the second part (presented in section 3.2) which describes and discusses the fabrication process of bioactive glass-based scaffold reinforced by collagen coating.

3.1.1 Scaffold preparation

For the preparation of scaffolds and performance of experiments, only reagents of analytical grade were used. A detailed list of all chemicals and associated suppliers can be found in appendix I.

Porous collagen scaffolds without bioactive glass particles

Porous collagen scaffolds can be produced by a variety of techniques. In this connection, lyophilization is often described in the literature due to its simplicity157–169. This process offers many advantages. Next to low costs, many parameters can be controlled like pore structure and pore size170. In addition, collagen can be easily combined with different additives, e.g. polymers160,163,166 or ceramics158,161,162,164,167,168 by suspension. The technique is based on the preparation of a collagen solution with desired component composition. Afterwards, the solution is frozen while ice crystals are formed. These crystals are sublimated and removed during the process of freeze drying and pores are formed. Other techniques for preparation of porous collagen scaffolds are e.g. electrospinning171–173, fiber extrusion174 or plastic compression (pc)175,176. In this thesis, lyophilization was the method of choice to produce porous collagen scaffolds. Therefore, a mixture of collagen, 1 M NaOH (sodium hydroxide) solution, 1 M HEPES buffer (4-(2-Hydroxyethyl)piperazine-1-ethanesulfonic acid dissolved in ultrapure water (UPW)), 10x concentrated Dulbecco’s Modified Eagle’s Medium (10x DMEM) and 0.01 M hydrochloric acid (HCl) was prepared. Collagen from calf was received as 0.5 % solution (5 mg/ml) dissolved in 0.01 M HCl. To receive scaffolds with different collagen concentrations, solutions with varying collagen

Collagen – As scaffold material

amounts were prepared. In detail, 1 M NaOH, 1 M HEPES buffer and 10x DMEM were mixed at a ratio of 1:1:2 (solution A). Depending on the final concentration of collagen (2, 3 and 4 mg/ml) the amounts of 0.5 % collagen solution and 0.01 M HCl were adjusted. The collagen mixture was then added slowly to solution A and gently stirred to avoid air bubbles. The final solution showed a pH between 7.5 and 8. For the preparation, all chemicals were refrigerated to a temperature of 4 °C to 8 °C. In addition, the whole process was carried out in an ice bath to avoid an early gelation of collagen. For further processing, the neutralized collagen solution was transferred to a 48-well plate and incubated at 37 °C for 6 h to initiate and complete the process of fibrillogenesis. Afterwards, the formed collagen gels were frozen. To investigate the influence of freezing temperature on the pore size, distribution and structure, two different freezing temperatures were chosen. Depending on the freezing conditions, ice crystals of different sizes are formed which is directly related to the size and distribution of pores in the final scaffold170. Therefore, collagen gels were frozen in a usual freezer at -20 °C and in a bath of liquid nitrogen at -196 °C before they were moved to the freeze dryer (Alpha 2-4 LSC plus, Christ, DE). After the collagen scaffolds were freeze-dried, they had to be crosslinked to enable a better handling (details given below, section Crosslinking process).

Reinforced porous collagen scaffolds with bioactive glass particles

As described by Cunniffe et al.162, there are two different methods to develop porous collagen scaffolds with ceramic reinforcement. The immersion method describes a process for indirect loading of ceramic particles into a collagenous structure. Porous collagen scaffolds, which were produced prior by lyophilization (as described above), are soaked in a ceramic suspension for a certain time period to incorporate ceramic particles and afterwards freeze-dried again. An important drawback using this technique is an uncontrollable amount of absorbed ceramic particles. Therefore, priority is given to the second method which enables a direct loading of ceramic particles into porous collagenous scaffolds being thus possible to control the relation between organic and inorganic phase. During the so-called suspension technique, ceramic particles are directly incorporated during the preparation of the collagen solution. For the reinforcement with bioactive glass particles, a concentration of 40 % collagen – 60 % bioactive glass (in wt.%) was chosen considering also results in the literature175. According to the fabrication process described above, collagen scaffolds with bioactive glass incorporation were produced as follows: solution A was prepared by mixing 1 M NaOH, 1 M HEPES buffer and 10x DMEM at a ratio of 1:1:2. Bioactive glass with the 45S5 composition (45S5 BG) with particle size of 2 µm was suspended in solution A by means of an ultrasound finger (Branson Digital Sonifier® 250, Emerson, DE) to receive a homogenous distribution of 45S5 BG particles. Depending on the final concentration of collagen, the amounts of 0.5 % collagen solution and 0.01 M HCl were adjusted and slowly added to solution A. To avoid early gelation of the collagen mixture, the process was carried out in an ice bath. In addition, all used chemicals were refrigerated to a temperature of 4 °C to 8 °C. The final mixture was transferred to a 48-well plate and incubated at 37 °C for 6 h to initiate and complete the process of

Collagen – As scaffold material fibrillogenesis. Afterwards, the formed collagen gels were frozen, lyophilized and crosslinked.

Crosslinking process

Collagen (as part of the native tissue) provides high strength and good mechanical properties due to inter- and intramolecular crosslinks within the triple helices and constituent collagen molecules (section 2.2.2 Structure of collagen). In addition, embedded cells in the collagen matrix support the mechanical stability. This advantage is missing when using collagen as purified biomaterial. During collagen processing, collagen fibrils can be extracted but intermolecular crosslinks are lost and so processed collagen exhibits a higher degradation rate and lower mechanical support. This weakness can be overcome by introducing additional bonds. Different methods have been introduced for crosslinking collagen. By ultraviolet irradiation (UV)177–182 and also by dehydrothermal treatment (DHT)177,182,183 physically induced crosslinks can be achieved. For chemically induced crosslinks different chemicals are available like carboxylic acids (e.g. acyl azide184–187), aglycones (e.g. genipin188–191), aldehydes (e.g. glutaraldehyde182,192–200), polyepoxy compounds201–204, carbodiimides (e.g. 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide182,204–209), isocyanates (e.g. hexamethylene diisocyanate185,199,200) or antioxidants (nordihydroguaiaretic acid (NDGA)210–213). Because of its high efficiency and low cost, glutaraldehyde (GA) is most extensively chosen as crosslinking agent. As described by Damink et al.192, glutaraldehyde readily enables crosslinking reactions. In case of collagen, it is generally assumed that aldehyde groups of GA covalently bind to -amine groups of lysine or hydroxylysine residues in collagen molecules (Figure 8). Newly formed intermediate molecules are included in the linkage which can entail serious consequences like toxicity and incompatibility with cells. Additionally, unreacted GA can be released causing inflammation and cell death which has been demonstrated in different studies198,199,214.

Lysin

A) e

-amine groups B) Glutaraldehyde Linkage

Aldehyde group

-amine groups

Figure 8: (A) Lysine as part of the triple helix in collagen. Illustration of reactive -amine groups. (B) Aldehyde groups of glutaraldehyde covalently bind to the amino groups of lysine or hydroxylysine residues in collagen molecules. By intermediate molecules collagen molecules can be crosslinked215 (created by ChemBioDraw Ultra 13.0).

Collagen – As scaffold material

To overcome problems associated with the application of GA, zero-length crosslinking chemicals are used. This means, the crosslinking reaction is facilitated but the agent does not remain in the chemical bond. Such reagents are carbodiimides. Carbodiimides like N-(3-Dimethylaminopropyl)-N’-ethylcarbodiimide (EDC) activate carboxyl groups of glutamic or aspartic residues to react with amino groups of other chains and amide-type crosslinks without any residual reactive groups are formed215,216(Figure 9). EDC is often used in combination with N-Hydroxysuccinimide (NHS) which increases the reaction efficacy and degree of crosslinking by suppression of side reactions205. The schematic mechanism of the crosslinking processes is shown in Figure 9. Compared to glutaraldehyde, EDC-crosslinked collagen showed superior bio- and cytocompatibility, slow degradation rate and reduced calcification182,204,215,217–219, therefore it was also chosen for this study. Thus, crosslinked (cl) collagen scaffolds were produced using the protocol below.

EDC activated carboxyl group EDC

Carboxyl and amino groups

NHS Amide-type crosslink

NHS activated carboxyl group

Figure 9: Schematically illustration of crosslinking collagen with carbodiimide205,215 (in this case 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimides (EDC) - created by ChemBioDraw Ultra 13.0).

According to Powell et al.220, a 50 mM MES (4-Morpholineethanesulfonic) acid solution in 40 vol.% ethanol was prepared using MES hydrate which resulted in a pH of 5.5. This pH showed to be the optimal environment for the activation of carboxyl groups with EDC. EDC and NHS were used at a ratio of 1:1 (60 mM)205. Therefore, N-(3-Dimethylaminopropyl)-N’-ethylcarbodiimide hydrochloride and NHS were dissolved in MES buffer solution. Due to the limited stability of carbodiimides in aqueous solutions, the crosslinking solution was always prepared freshly221. Lyophilized scaffolds were immersed for 4 h at room temperature. After crosslinking, scaffolds were removed and washed in 0.1 M Na2HPO4 (disodium hydrogen phosphate) for 2 h to hydrolyze remaining O-acylisourea of the carbodiimide205,216. Subsequently, scaffolds

Collagen – As scaffold material were washed in UPW, frozen and lyophilized. The success of crosslinking was investigated by Fourier transform infrared spectroscopy (FTIR). This effective technique enables identification of the chemical composition by exposure to infrared radiation. Therefore, KBr (potassium bromide) was mixed with 1 wt.% of the sample to be analyzed, pestled and pressed into pellets using an electrohydraulic pressing device (8104 N, Mauthe Maschinenbau, DE). The spectra were recorded in absorbance mode and collected between 1800 cm-1 and 600 cm-1. Pure KBr was used to correct the background noise. Figure 10 shows the FTIR spectra of pure

collagen before and after the crosslinking process. The characteristics peaks of interest are labeled.

-1

1632 cm 1632 I Amide

-1 Pure collagen - crosslinked

Pure collagen - uncrosslinked

1546 cm 1546 II Amide

-1

1240 cm 1240 III Amide

1082 cm 1082 stretching C-O-C

-1

1033 cm 1033 stretching C-O Absorbance [a.u.] Absorbance

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 10: FTIR spectra of pure collagen before and after the crosslinking process showing the formation of an additional peak at 1546 cm-1 (amide II) after crosslinking. The identified peaks are discussed in the text.

The amide I band, which is typical for the triple helical structure in non-denatured collagen175, shows characteristic frequencies in the region at 1632 cm-1 and can be associated with the C=O stretching vibrations175,222–226. Amide III can be identified at 1240 cm-1 and belongs to N-H deformation175,227. Crosslinked samples exhibit an additional peak at 1546 cm-1 (related to the N-H deformation and C-N stretching224,225,227) which indicates an extra amide-type bond (amide II) in the developed crosslinked network. This confirms the success of the crosslinking process. As described before227, the intensity of the amide II increases with the degree of crosslinking. In addition, peaks, which are visible at 1082 cm-1 and 1033 cm-1, can be associated with the C-O-C and C-O stretching of the carbonate present in the side chain of collagen228. For the determination of the degree of crosslinking of porous collagen-based scaffolds, the ninhydrin assay was performed229,230. Therefore, a ninhydrin solution was prepared by dissolving 2 wt.% of ninhydrin powder in ethanol (99.8 vol.%). Collagen scaffolds (50 ± 0.6 mg) with and without crosslinks were immersed in 5 ml of prepared ninhydrin solution using closable containers and heated in a water bath at 100 °C for 20 minutes.

Collagen – As scaffold material

During a complex reaction231, ninhydrin reacts with free amino groups and the so-called Ruhemann's purple is formed. Simplified, after decarboxylation of the amino acids, the amino group is transferred to ninhydrin. This complex reacts with a second ninhydrin molecule and forms the Ruhemann’s purple (Figure 11).

Ninhydrin α-amino acid

Ruhemann’s purple

Figure 11: Simplified mechanism of ninhydrin reaction with amino acids (created by ChemBioDraw Ultra 13.0).

The intensity of the violet color of Ruhemann’s purple is directly proportional to the concentration of the free amino groups in the collagenous scaffolds. For analysis, optical absorbance was measured at 570 nm via UV-Vis spectroscopy (Specord 40, Analytik Jena). The concentration of free amino groups can be determined from a standard calibration curve (glycine concentration vs. absorbance). Due to amide-type crosslinks, the concentration of free amino groups decreases with increasing degree of crosslinking. The degree of crosslinking (n=4) was calculated according to

(N (N Degree of crosslin ing [ ] ( uc cl)  [Eq. 2] (N uc

where (NH2)uc and (NH2)cl are the determined concentrations of free amino groups in uncrosslinked and crosslinked collagen scaffolds, respectively. After the crosslinking process with NHS and EDC, porous collagen-based scaffolds exhibit 50 ± 1 % of crosslinks.

To summarize the fabrication of porous collagen-based scaffolds with and without bioactive glass particles, the following flow chart (Figure 12) illustrates the different steps:

Collagen – As scaffold material

Collagen mixture: Porous collagen 1 M NaOH Collagen mixture transferred scaffolds without 1 M HEPES to 48-well plate, 45S5 BG particles 10x DMEM gelation at 37 °C for 6 h 0.5 % Collagen solution + 0.01 M HCl

Collagen gel frozen at -196 °C or Lyophilization Crosslinking -20 °C

Reinforced porous Collagen mixture: collagen scaffolds 1 M NaOH Collagen mixture transferred 1 M HEPES }+ 45S5 BG particles to 48-well plate, with 10x DMEM gelation at 37 °C for 6 h 45S5 BG particles 0.5 % Collagen solution + 0.01 M HCl

Collagen gel frozen at -196 °C or Lyophilization Crosslinking -20 °C

50 mM MES Washing in 0.1 M

in 40 vol.% ethanol Immersion for 4 h at Na2HPO4 for 2 h, Crosslinking + 60 mM EDC room temperature rinsing in UPW and + 60 mM NHS drying by lyophilization

Figure 12: Flow chart of the manufacturing processes for porous collagen scaffolds with and without bioactive glass reinforcement.

3.1.2 Morphological and microstructural characterization

Porous collagenous scaffolds with and without bioactive glass reinforcement, which were produced previously by the process described above, were characterized in terms of morphology and microstructure. The results are presented next.

Pure collagen

As described before, collagen scaffolds were produced with varying amounts of collagen (2, 3 and 4 mg/ml). In addition, two different freezing temperatures were chosen prior to lyophilization. After the crosslinking process, the obtained scaffolds were investigated to analyze their morphology and microstructure. Scaffolds with 2 mg/ml collagen were directly excluded for further experiments due to their extremely high rate of shrinking during the process of freeze drying. Scaffolds with 3 and 4 mg/ml were investigated via scanning electron microscopy (SEM). Figure 13 shows SEM images of cross-sections of collagen scaffolds with varying concentrations of collagen. The influence of different freezing temperatures on the pore structure and distribution can be observed. Comparing the different freezing conditions, freezing at -20 °C creates uniform and cylindrical formed scaffolds with a homogenous distribution of pores in the range of 100 to 500 µm. In contrast, scaffolds formed by liquid nitrogen appear very inhomogeneous. Even though liquid nitrogen is expected to form small ice crystals and thus small pores, large pores up to 2.5 mm are seen to have formed. Because of technical limitations, it was

Collagen – As scaffold material

not possible to keep the temperature of -196 °C during the whole process of freeze drying. It is assumed, that small ice crystals assemble during the increase in temperature to reduce their high surface energy. As a result, bigger ice crystals are formed which are removed during the freeze drying process and large pores are left. Thus, for further investigations, -20 °C was the chosen temperature for freezing scaffolds before they were lyophilized.

A - 3 mg/ml A - 4 mg/ml

B - 3 mg/ml B - 4 mg/ml

Figure 13: SEM images of porous collagen scaffolds with varying collagen concentrations after lyophilization. A) Collagen scaffolds frozen at -20 °C, B) Collagen scaffolds frozen at -196 °C.

The porosity P (n=4) of the collagenous sponges (prepared at -20 °C) was determined according to

ρ P ( - sample ) [Eq. 3] ρmaterial

where ρSample is the density of the collagenous scaffold and ρmaterial is the density of pure collagen (in this case

3 232 1.34 g/cm ) . For the determination of ρsample, the volume of each scaffold was calculated by equation [4]

d ( ) h [Eq. 4]

where d is the diameter and h the height of each scaffold. Table 3 shows the different values of the measured samples. Based on the volume V and the mass m (Table 3), the density ρsample was calculated as follows:

Collagen – As scaffold material

m ρ [Eq. 5] sample

Calculations showed a porosity of around 98 % (Table 3) for both concentrations (3 and 4 mg/ml collagen).

Table 3: Calculated values of density ρsample and porosity P of pure collagen scaffolds with different concentrations.

3 No. Mass [g] Diameter [cm] Height [cm] ρsample [g/cm ] Porosity [%] 3 mg/ml collagen: 1 0.0052 0.63 0.73 0.023 98.29 2 0.0064 0.61 0.72 0.030 97.73 3 0.0059 0.59 0.75 0.029 97.85 4 0.0055 0.61 0.73 0.026 98.08 4 mg/ml collagen: 5 0.0084 0.65 0.95 0.027 98.01 6 0.0076 0.63 0.93 0.026 98.04 7 0.0088 0.68 0.93 0.026 98.06 8 0.0079 0.67 0.95 0.024 98.24

Collagen with bioactive glass reinforcement

Figure 14 shows collagen scaffolds with 3 and 4 mg/ml collagen after the reinforcement with 60 wt.% of 45S5 BG particles. After the incorporation of bioactive glass into the collagen solution, an ultrasonic finger was used. This resulted in a homogenous distribution of the ceramic particles which can be observed in the SEM images (Figure 14). Single particles can be seen without agglomerations and also homogeneously distributed pores were formed between 50 and 150 µm (for both concentrations).

A - 3 mg/ml + 60 wt.% of 45S5 BG B - 4 mg/ml + 60 wt.% of 45S5 BG

Bioactive glass particles Bioactive glass particles

Figure 14: SEM images of collagen-BG scaffolds, frozen at -20 °C A) 3 mg/ml Collagen + 60 wt.% of 45S5 bioactive glass particles and B) 4 mg/ml Collagen + 60 wt.% of 45S5 bioactive glass particles.

Collagen – As scaffold material

Because no major difference could be observed between 3 and 4 mg/ml, a selection was made choosing only scaffolds with 4 mg/ml collagen (with and without 45S5 BG particles incorporation) for further experiments. For chemical analysis of the composite scaffolds, 1 wt.% of each sample was used and mixed with KBr, as described before. The KBr pellets of collagen sponges with and without bioactive glass were analyzed by FTIR in absorbance mode. The spectra were collected between 1800 and 600 cm-1. The most characteristic peaks

are labeled.

-1

1632 cm 1632 I Amide

-1 Pure collagen

Collagen + bioactive glass particles

1546 cm 1546 II Amide

Bioactive glass: 1100-1000 cm-1

Si-O-Si stretching, P-O stretching -1

Collagen:

-1 Amide III Amide 1240 cm 1240 1082 and 1033 cm

C-O-C and C-O stretching Absorbance [a.u.] Absorbance

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 15: FTIR spectra of pure collagen (4 mg/ml) and collagen (4 mg/ml) + 60 wt.% of 45S5 bioactive glass particles showing a growing peak (between 1100 and 1000 cm-1) after the addition of bioactive glass. The identified peaks are discussed in the text.

The FTIR spectrum of collagen with 45S5 BG particles slightly differs from that of pure collagen (Figure 15). Collagen shows typical peaks at 1632 cm-1 and 1240 cm-1 which can be attributed to the C=O stretching vibrations of the amide I175,222–226 and to the N-H deformation of amide III175,227, respectively. Amide II224,225,227 can be identified at 1546 cm-1 and belongs to the N-H deformation and C-N stretching of the amide-type bond in the crosslinked collagen network. Peaks, which are visible at 1082 cm-1 and 1033 cm-1, belong to the C-O-C and C-O stretching of the carbonate (carboxylic acids) present in the side chains of collagen228. However, typical peaks for 45S5 bioactive glass can be found in the same region between 1100 cm-1 and 1000 cm-1 (Si-O-Si stretching and P-O stretching175,233). The addition of bioactive glass to the collagen matrices is confirmed by the growing peaks in the region between 1100 cm-1 and 1000 cm-1 but due to the overlap it is hard to distinguish the different peaks and assign them explicitly to 45S5 bioactive glass. For the determination of the porosity P of the composite scaffolds (collagen sponges with 45S5 BG particles) equations [3], [4] and [5] were used. To ascertain the corrected density of the composite material ρComposite used in equation [3], the different ratios of collagen to bioactive glass must be considered. Before crosslinking,

Collagen – As scaffold material each scaffold contained 40 wt.% of collagen and 60 wt.% of bioactive glass. Because it is assumed that some material is lost during the crosslinking process, thermogravimetric analysis (TGA, STA 449 F3 Jupiter, Netzsch, DE) was carried out to quantify the amount of collagen and bioactive glass after the crosslinking process. Heating was performed in ambient atmosphere and with a heating rate of 5 K/min. Results revealed a ratio of 48:52 (in wt.%) for collagen and bioactive glass after crosslinking, respectively. Taking equation [5] into account,

ρComposite was determined as

mcomposite mcoll m ρcomposite [Eq. 6] composite coll

3 232 where mcoll and Vcoll represent the mass and volume of collagen (ρCollagen = 1.34 g/cm ) and mBG and VBG

3 29 represent the mass and volume of 45S5 bioactive glass (ρBG = 2.7 g/cm ) , respectively.

Based on the results of the TGA measurement, the density ρComposite of freeze-dried collagen-BG scaffolds after crosslinking was calculated as 1.82 g/cm3. Therefore, collagen-BG composites still exhibit a porosity of  95 % which means the overall porosity was not significantly influenced compared to pure collagen scaffolds.

3.1.3 Swelling properties and degradation behavior

The degradation behavior of hydrogels is affected by many parameters of the polymer itself like degree of crosslinking, molecular weight and also environmental factors (e.g. medium, pH, temperature) play a big role234,235. In addition, the water binding ability is an important feature to estimate the capability for tissue engineering applications236,237. The swelling kinetics describes the transportation of solvents into the polymeric network structure which directly influences the degradation behavior. The knowledge about water absorption is needed to understand and explain the release and degradation of collagen. Although collagen is insoluble in water, solvent molecules can occupy the inter- and intrafibrillar spaces which leads to swelling of the fibril structure.

Swelling properties

To evaluate the water absorption of a material, common and simple methods are sorption experiments, as described in the literature161,236–238. In this study, both collagen scaffolds, with and without bioactive glass reinforcement (n=5), were chosen and compared. The swelling study was performed as scaffolds were immersed in 10 mM phosphate buffered saline (PBS) at room temperature up to 24 h. After different time intervals (15 min, 30 min, 60 min, 24 h), samples were removed and excess solution was gently dried with a filter paper (qualitative filter paper 410). Water uptake was calculated as

Collagen – As scaffold material

m md ater upta e [ ] ( )  [Eq. 7] md

where mw and md are the mass of the wet and dry samples, respectively. Results revealed that equilibrium was reached after 1 h of immersion. Figure 16 shows the relative water uptake of collagenous porous scaffolds with and without 45S5 BG particles after 1 h in PBS.

1600 60 min 30 min 1400 15 min

1200

1000

800

600

400 Relative[%] water uptake

200

0 Collagen Collagen + BG

Figure 16: Results of the swelling study on collagen-based scaffolds: relative water uptake in PBS for collagen and collagen-BG scaffolds up to 1 h.

In general, swelling describes the binding of water molecules via hydrogen bonds which increases the distance between the single collagen fibers. The diagram (Figure 16) clearly demonstrates high water uptake ability up to 1600 % for pure collagen matrices due to a high diffusion rate which enables a good transportation of nutrients and oxygen during in vitro cell studies. However, during the swelling study a high increase of pore size and porosity entailed a total loss of shape observed in case of pure collagen scaffolds. Because collagen is insoluble in water, it transforms into a hydrogel. This means, high water content also leads to a decrease of mechanical properties which will be discussed later (3.2.4 Mechanical characterization). Collagen-BG composites proved to be more consistent in their cylindrical dimensions. Firstly, the decrease in swelling rate of collagen-BG composites can be attributed to the lower collagen amount compared to pure collagen scaffolds. In addition, it is assumed that bioactive glass particles and collagen fibers interact which leads to a reduction of hydrophilic groups and hence to a reduced swelling behavior and a more stable shape of the sample. Therefore, the swelling behavior of collagen can be positively controlled by addition of bioactive glass.

Collagen – As scaffold material

Degradation behavior

Degradability plays an important role in tissue engineering applications. The degradation behavior gives information about the stability of the scaffolds which is important for the formation of new tissue. The in vitro degradation of collagen and collagen-BG composite scaffolds was investigated by bacterial collagenase from Clostridium histolyticum with a collagenase activity of 125 CDU/mg (one collagen digestion unit (CDU) releases peptides from collagen equivalent in ninhydrin color to 1 µm of leucine in 5 h at pH 7.4 at 37 °C in the presence of calcium ions239). A buffer solution (pH = 7.4) was prepared containing 0.1 M Tris-HCl (Tris(hydroxymethyl)aminomethane hydrochloride), 0.005 M CaCl2 (calcium chloride) and 0.05 mg/ml NaN3 (sodium azide). Samples (n=4) of porous collagen sponges with and without bioactive glass reinforcement were immersed in 1 ml of Tris-HCl buffer solution at 37 °C. After one hour, 1 ml of Tris-HCl buffer with 200 CDU/ml was added to get the desired concentration of 100 CDU/ml. After different time points (1 h and 24 h), the enzymatic degradation was stopped by 0.25 M EDTA buffer (ethylenedinitrilotetraacetic acid) which binds Ca+. Scaffolds were washed three times in UPW, frozen and afterwards lyophilized. Scaffolds, which were immersed in pure Tris-HCl buffer solution without collagenase, were used as reference. The degradation rate was calculated as

mi ma Degradation [ ] ( )  [Eq. 8] mi

where mi is the initial weight of each scaffold and ma is the dry weight after degradation, respectively.

Table 4: Degradation rates of pure collagen and collagen-BG composites after the immersion in collagenase solution (100 CDU/ml).

1 h 24 h Reference after 24 h Pure collagen 33 ± 6 % 100 % 0.3 ± 0.2 % Collagen + 45S5 BG 21 ± 3 % 75 ± 5 % 5 ± 1 %

Table 4 shows the different degradation rates of pure collagen compared to collagen containing 45S5 bioactive glass particles after 1 h and 24 h of immersion in 0.1 M Tris-HCl buffer with collagenase (100 CDU/ml). Results revealed a higher stability of collagenous scaffolds after the reinforcement with bioactive glass. After 24 h of immersion, still 25 % of the composite scaffold material remains whereas pure collagen is completely digested. These results can be attributed to the reduced swelling behavior of reinforced collagen scaffolds, as already discussed. Reference samples of collagen with 45S5 BG showed a weight loss of 5 ± 1 % after 24 h in enzymatic-free buffer solution which can be explained with the slow release of bioactive glass particles over time. Similar results, which describe the reduced degradation behavior of polymeric matrices after ceramic reinforcement, are also reported in the literature240–243.

Collagen – As scaffold material

3.1.4 Evaluation of bioactivity

The bioactive behavior is an important aspect that anticipates the bone-bonding ability of a biomaterial. For the assessment of bioactivity in vitro, the standard protocol introduced by Kokubo et al.244 was applied although this method has recently attracted some criticism245,246. However, this well-known technique was applied to compare samples with the results received so far in the literature29,247–249. To study the acellular mineralization potential in vitro, samples (n=3) of pure collagen and collagen with bioactive glass reinforcement were put in closeable containers of polypropylene, immersed in 50 ml of simulated body fluid (SBF) and placed in an orbital shaker (90 rpm) at 37 °C. After different time points (3 d, 7 d and 14 d), samples were removed, gently washed in UPW and analyzed in terms of hydroxyapatite (HA) formation on the surface. For optical characterization of the biomineralization process, SEM observations were used. The chemical composition was confirmed by FTIR measurements.

Preparation of simulated body fluid

For the preparation of SBF, different chemicals were required which are listed in Table 5. For the preparation of 2 liters SBF, a volumetric flask was filled with 1.5 liters of UPW. The reagents were added one by one in the right order, given by Table 5, until every reagent was completely dissolved. The solution was continuously stirred and checked in terms of transparency, colorlessness and precipitation. After dissolving the ingredients, a pH of 7.4 at 37 °C was adjusted by a few drops of concentrated HCl. Afterwards, the bottle was filled up to 2 liters with UPW and stored in the fridge.

Table 5: List of chemicals used to prepare 1 liter of SBF.

Order Chemical Amount [g] 1 Sodium chloride (NaCl) 7.996

2 Sodium hydrogen carbonate (NaHCO3) 0.350 3 Potassium chloride (KCl) 0.224

4 Dipotassium hydrogen phosphate trihydrate (K2HPO4  3 H2O) 0.228

5 Magnesium chloride hexahydrate (MgCl2  6 H2O) 0.305 6 1 M hydrochloric acid (HCl) 35 ml

7 Calcium chloride dehydrate (CaCl2  2 H2O) 0.368

8 Sodium sulfate (Na2SO4) 0.071 9 Tris(hydroxymethyl)aminomethane (TRIS) 6.057*

Bioactivity study

Figure 17 and 18 show SEM images of pure collagen scaffolds before and after the immersion in SBF. After 3 days of immersion, first precipitates of calcium phosphate (CaP) can be recognized due to their typical

32 * TRIS was added particularly slowly to avoid precipitation.

Collagen – As scaffold material morphology. Small round-shaped crystals around 650 nm are seen to be randomly distributed on the surface of the collagen scaffold indicating its bioactive behavior (Figure 17).

Pure Collagen - 0 d Pure Collagen - 3 d Pure Collagen - 3 d

Figure 17: SEM images of pure collagen scaffolds (4 mg/ml) before and after 3 days of immersion in SBF at different magnifications showing the formation of round-shaped hydroxyapatite crystals already after 3 days.

Pure Collagen - 7 d Pure Collagen - 14 d Pure Collagen - 14 d

Figure 18: SEM images of pure collagen scaffolds (4 mg/ml) after 7 and 14 days of immersion in SBF at different magnifications showing the highly bioactive behavior of pure collagen. After 14 days of immersion the surface is homogenously covered with HA crystals.

After 7 days, the surface is homogeneously covered by a dense layer of HA crystals which can be clearly

Collagen – As scaffold material

identified at higher magnifications (Figure 18). These results indicate that the mineralization process of collagen is constantly progressing with longer immersion time. It is also observed that HA aggregates and form the typical cauliflower structure21,250 (particle size around 2.5 µm), visible after 14 days. The bioactive behavior of collagen has been also described before in the literature. Rhee et al.251 as well as Marelli et al.252 described the formation of hydroxyapatite crystals in collagen matrices after the immersion in SBF. According to expectations, collagen scaffolds with added 45S5 bioactive glass particles showed enhanced bioactivity. Compared to pure collagen, which exhibits small precipitates (650 nm) of HA after 3 days in SBF (Figure 17), collagen with 45S5 BG reinforcement showed HA nodules with diameter  1.5 µm during the same period of time (Figure 19). The amount of HA deposition accelerated with conditioning time in SBF. At day 14, the scaffold was seen to be completely covered by HA crystals with a mean size of 4 µm while the collagen structure was not altered.

Collagen + 45S5 BG - 0 d Collagen + 45S5 BG - 3 d Collagen + 45S5 BG - 14 d

Collagen + 45S5 BG - 0 d Collagen + 45S5 BG - 7 d Collagen + 45S5 BG - 14 d

Figure 19: SEM images of collagen composite scaffolds (4 mg/ml) with 60 wt.% of 45S5 bioactive glass particles before and after immersion in SBF for 3, 7 and 14 days (at different magnifications) showing the accelerated formation of hydroxyapatite crystals after the addition of 45S5.

Apatite formation was confirmed by FTIR. Samples were prepared as KBr pellets and analyzed in absorbance mode at IR wavenumbers between 1800 cm-1 and 600 cm-1. Figure 20 shows the FTIR spectra of pure collagen sponges immersed in SBF for different time periods. Amide peaks, visible at 1632 cm-1, 1546 cm-1 and 1240 cm-1, belong to the collagenous network and were already explained above (section 3.1.2). The apatite formation, which was visible in the SEM images, can now be proved by the growing peak in the wavenumber region between 1080 cm-1 and 1030 cm-1. With increasing

Collagen – As scaffold material immersion time in SBF the broad peak is seen to grow and this effect can be attributed to P-O stretching in the

hydroxyapatite layer253.

-1

1632 cm 1632 I Amide

1546 cm 1546 II Amide

-1

-1 1080 - 1030 cm

P-O stretching

1240 cm 1240 III Amide

14 d

7 d

3 d

Absorbance [a.u.] Absorbance

1 d

0 d

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 20: FTIR spectra of pure collagen scaffolds (4 mg/ml) after 0, 1, 3, 7 and 14 days of immersion in SBF showing a growing peak in the area between 1080 and 1030 cm-1 indicating the formation of HA. The identified peaks are discussed in the text.

-1

1020 cm 1020 stretching P-O

-1

960 cm 960 stretching P-O

-1

875 cm 875 bending C-O

1632 cm 1632 I Amide

-1

1240 cm 1240 III Amide

1546 cm 1546 II Amide

14 d

7 d Absorbance [a.u.] Absorbance

3 d

1 d

0 d

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 21: FTIR spectra of collagen composite scaffolds (4 mg/ml) with 60 wt.% of 45S5 bioactive glass after 0, 1, 3, 7 and 14 days of immersion in SBF showing growing peaks at 1020, 960 and 875 cm-1 indicating the formation of HA. The identified peaks are discussed in the text.

Collagen – As scaffold material

The FTIR spectra of collagenous composite scaffolds containing 45S5 BG reinforcement can be seen in Figure 21. These spectra show additionally peaks at 1020 cm-1, 960 cm-1 and 875 cm-1 after the immersion in SBF.

-1 -1 3- The growing peak at 1020 cm and its small shoulder at 960 cm are related to the P-O stretching of PO4 as part of HA formation on the surface253. The small peak developed at 875 cm-1 can be attributed to the C-O bending and brings out the carbonate present in the hydroxyapatite253,254.

Collagen – As coating material

3.2 Collagen as coating material on bioactive glass-based scaffolds - Manufacturing techniques, results and discussion

This chapter deals with the second concept proposed in this thesis to produce composite scaffolds using bioactive glass and collagen. Shortly summarized, highly interconnected porous bioactive 45S5 glass-based scaffolds were prepared by the foam replica technique29. To immobilize proteins on these scaffolds, the surface was functionalized and coated with collagen, followed by a crosslinking step. Since some characterization methods required a flat surface, dense 45S5 BG pellets were used in this case. Because scaffolds and pellets were fabricated under same conditions and processing parameters, the results were considered to be directly transferred to scaffolds although some differences in the results from pellets and scaffolds are discussed when relevant.

3.2.1 Sample preparation

For the composites and experiments all applied reagents were of analytical grade and purchased from different suppliers. A detailed list containing all chemicals can be found in appendix I.

Scaffold and pellet fabrication

A melt-derived bioactive glass powder (of the 45S5 composition, 45S5 BG) with a particle size of  2 µm was used to fabricate either porous scaffolds or dense pellets. Additionally, a binder (polyvinyl alcohol, PVA) and disperser (KV 9062) were used for the preparation of the slurry. Regarding the foam replica process originally described by Chen et al.29, the technique was applied in this investigation to produce three-dimensional bioactive glass-based scaffolds with a highly interconnected porous structure similar to cancellous bone. Commercially available polyurethane (PU) foam with 45 ppi (pores per inch) served as sacrificial template for the replication method. The inner structure of the scaffold complied with the selected ppi, so the component geometry can only be influenced by the external dimensions. Using a round cutting device, cylindrical specimens were punched out of the PU foam with dimensions of 7x12 mm (height x diameter). To provide a better adhesion of the bioactive glass particles to the surface of the PU foam, specimens were pretreated in an ultrasonic bath using 70 vol.% 2-propanol to remove residual contaminations left from the manufacturing process (e.g. oil or dust). The polymeric foam was then coated by immersion into the prepared 45S5 BG slurry. Subsequently, the slurry infiltrated the structure of the foam and 45S5 BG particles adhered on the surface of the template29. The slurry contained 50 % bioactive glass powder, 47.7 % ultrapure water, 2 % disperser and 0.3 % binder (in wt.%). For the slurry preparation, PVA was dissolved in water at 80 °C. After cooling down to room temperature (RT), KV 9062 and bioactive glass powder were added. To receive a homogenous slurry, the ingredients were added slowly step by step under vigorous stirring using a magnetic

Collagen – As coating material

stirrer. The cleaned PU foams were then immersed into the slurry, retrieved after 5 minutes and squeezed to remove excess slurry. After drying at 60 °C for minimum 2 h, a double coating was applied by repetition of the coating step to increase the thickness of the 45S5 BG layer in order to reduce the shrinkage rate and to receive enhanced mechanical stability which should result in a better handling of the scaffolds. Before sintering, the samples were dried again at 60 °C for at least 24 h. Finally, the received green bodies were heat treated to burn out the polymeric foam as well as organic components of the slurry and to sinter the bioactive glass structure. For the sintering, a progressive sintering profile was chosen, as represented in Figure 22. In a first step, a heating rate of 2 °C/min up to 400 °C was applied to burn out the template structure. After 1 h holding time, this was connected to a second step, heating up to 1050 °C (2 °C/min) and holding 2 h to densify the structure and complete the sintering process. After sintering, the scaffolds were left to cool down naturally in the furnace.

1200

1050 °C 1000 2 h

800

2 °C/min 600

400 °C 400

1 h Temperature [°C] Temperature

2 °C/min 200 Heating phase Sintering Natural cooling

2 4 6 8 10 12 14 16 Time [h]

Figure 22: Heating profile used to fabricate bioactive glass-based scaffolds.

The sintering process yielded porous bioactive glass-based foams with an average size of 4.2 ± 0.2 mm in height and 7.5 ± 0.3 mm in diameter which implies shrinkage of around 40 % during sintering.

Table 6: Calculated values of density ρsample and porosity P of bioactive glass-based scaffolds after sintering.

3 No. Mass [g] Diameter [cm] Height [cm] ρsample [g/cm ] Porosity [%] 1 0.0221 0.74 0.43 0.120 95.57 2 0.0264 0.79 0.42 0.128 95.25 3 0.0289 0.76 0.45 0.142 94.76 4 0.0374 0.73 0.42 0.213 92.12 5 0.0351 0.71 0.42 0.211 92.18 6 0.0251 0.77 0.38 0.142 94.75

The porosity P of the sintered scaffolds was determined considering equations [3], [4] and [5] where ρsample is

Collagen – As coating material

the calculated density of the sintered scaffold and ρmaterial is the true density of amorphous 45S5 bioactive glass (in this case 2.7 g/cm3)29. Table 6 shows the different values of the measured samples. The calculations indicated an average porosity of  94 % which is in accordance with the literature29. For investigation of the macro- and microstructure of the obtained scaffolds scanning electron microscopy (SEM) observations were used.

Hollow strut

Figure 23: SEM images of as-fabricated bioactive glass-based scaffolds after sintering at different magnifications showing the highly porous structure of the scaffold with interconnected porosity and hollow struts.

Figure 23 illustrates the typical macro- and microstructure of as-fabricated bioactive glass-based scaffolds obtained by the polymer replica technique at different magnifications. The structure is porous and highly interconnected and shows a macrostructure similar to cancellous bone whereas the sintered microstructure exhibits the presence of the typical hole in the center of the foam struts. The SEM images show a pore size of sintered samples in the range of 250 – 500 µm. The struts diameter is in the range of 50 to 100 µm. The hollow nature of the struts is typical for scaffolds produced by this method and can be associated with the burning out of the sacrificial PU foam. For the fabrication of bioactive glass pellets, 0.3 g of 45S5 BG powder was filled into a pressing tool ( = 10 mm), which was lubricated with glycerin, and compressed by applying 104 N in an electrohydraulic pressing device (Mauthe Maschinenbau, DE). The pellets did not require a drying step and could be sintered right after the fabrication. The sintering profile was similar to the one used for scaffolds (Figure 22) without holding time at 400 °C.

Surface functionalization, collagen coating and optimization of parameters

Relevance of functionalization

Interactions between surfaces of biomaterials or medical implants and cells play an essential role in determining the success of application of the biomaterial. Body responses and inflammatory processes are mediated by protein adsorption due to cellular interactions between adhesive molecules and cells255–259. If the protein is recognized as part of the natural tissue, cells can attach, proliferate and differentiate. Otherwise, inflammation

Collagen – As coating material

can cause cell death and, as a consequence, the failure of devices259–262. Due to this sensitive reaction of cells, a common strategy in tissue engineering is the modification of surfaces with different proteins. In 2003, Rosengren et al.263 described selective adsorption of proteins from human plasma onto two bioactive glass- ceramics by immersion. However, due to electrostatic and hydrophobic interactions, this method may bear a challenge. Proteins reveal chemical complexity and a fragile nature and as a result of this, the native conformation of the used protein can change considerably247,264–267. Therefore, to overcome this problem and to prevent the alteration of the biological functionality, proteins can be grafted to the carrier surface by chemical bonding which has been applied in different studies268–270. Additionally, the protein release kinetics can be controlled by surface functionalization (see section 3.2.2 Release behavior). One typical approach to modify the surface of bioactive ceramics is silanization271–274 which describes the chemical bonding of organo-functional silanes onto surfaces. Silanes are a class of molecules bearing at least one silicon (Si) molecule in oxidation state +IV as central heteroatom while organosilanes contain at least one silicon-carbon (Si-C) bond275 (Figure 24).

Figure 24: Example of the chemical structure of a silane (left) and an organosilane (right - created by ChemBioDraw Ultra 13.0).

Above-mentioned organo-functional silanes have two different reactive moieties bond to their Si atom and can couple organic molecules to inorganic surfaces via covalent bonds275,276. In the 1940s, during the development of fiberglass-reinforced composites, the value of these silanes as coupling agent was discovered275,277. Silane coupling agents can be described by the simplified molecular structure X-R-Si(OR)3. The X group represents one of the reactive groups, the organo-functional group, and is selected to be reactive with the compound of interest. The R group is the chemical linkage between the organo-functional group and Si. OR are hydrolysable ester bonds such as methoxy- (-OMe), ethoxy- (-OEt) or acetoxygroups (-OAc) and provide the second reactive group275,276.

Ethoxy group: -OEt Amino group: -NH2 Propyl group: Linkage

Figure 25: Chemical structure of APTS, the protein coupling agent used in this study (created by ChemBioDraw Ultra 13.0).

Collagen – As coating material

One of the most studied and applied coupling agent is (3-aminopropyl)triethoxysilane (APTS, a sol-gel precursor), which has been used in several studies before247,259,271,278–281, and it was also chosen for this study. APTS consists of three ethoxy groups and one propyl group as linkage between Si and the amino group (Figure 25). The surface functionalization via silanization is a very complex process and is influenced by many different parameters. Reaction time, temperature, solvent as well as silane concentration and nature influence can affect the silanization procedure282,283. The following simplified scheme (Figure 26) describes the surface functionalization of 45S5 bioactive glass using APTS in four steps which can be divided into hydrolysis (I), condensation reaction (II), hydrogen bonding (III) and bond formation (IV). For the first reaction, adsorbed moisture on the surface is sufficient to start the process of hydrolysis (I). Water molecules attack the ester bonds in the ethoxy groups (Si-OEt). They are readily hydrolyzed and replaced by hydroxyl groups (-OH) to form reactive silanol groups (Si-OH). Every hydrolyzed ethoxy group produces one ethanol molecule275,282,284. Hydrolysis (I) is followed by a slower condensation reaction (II). Through a self-condensation process, silanol groups can react with each other which results in a polymeric structure with very stable siloxane bonds (Si-O- Si) and water as byproduct. Additionally, silanol groups can react with remaining ethoxy groups and form the

275,282,284 same polymeric structure by elimination of ethanol (C2H5OH) . The final condensed product is now able to react with -OH groups located on the surface of inorganic surfaces (e.g. 45S5 bioactive glass). By hydrogen bonding (III), a tight polymeric network of siloxanes assembles in a uniform layer on the surface of 45S5 BG stabilized by coordinative van-der-Waals hydrogen bonds.

(I) Hydrolysis2–4 APTS Hydrolyzed APTS

(II) Condensation reaction2–4 Final condensed product

R: -OH -OEt

Collagen – As coating material

(III) Hydrogen bonding2–5

Bioactive glass surface Hydrogen bond Bioactive glass surface

(IV) Bond formation2–5

Reactive amino groups



Covalent bond

Bioactive glass surface Functionalized bioactive glass surface

Figure 26: Schematic representation of the surface functionalization of 45S5 bioactive glass with APTS divided into (I) hydrolysis, (II) condensation reaction, (III) hydrogen bonding and (IV) bond formation (created by ChemBioDraw Ultra 13.0).

However, these bonds are not very stable. By heat treatment, water is removed and the hydrogen bonds are replaced by more stable covalent bonds (IV). After this, the surface is functionalized with reactive amino groups and compounds of interest can be directly coupled (e.g. proteins like collagen).

Collagen coating

The process of collagen coating on bioactive glass surfaces by silanization was firstly described by Chen et al.247, so preliminary work was based on this protocol. Therefore, an aqueous solution of 0.45 M APTS was prepared. The pH value was adjusted to 8 using 1 M hydrochloric acid (HCl). The produced bioactive glass- based scaffolds were immersed into the solution under stirring conditions and heated up to 80 °C. After an

Collagen – As coating material immersion time of 4 hours, the samples were removed, cleaned in UPW and dried at ambient atmosphere. For the coating, collagen was received as 0.5 % solution (5 mg/ml) dissolved in 0.01 M HCl. For the protein loading, a 0.01 M HCl solution was prepared with 0.350 mg/ml collagen. Afterwards, the functionalized samples were immersed into this solution and incubated at room temperature. After 24 hours of immersion, they were removed, rinsed with UPW and dried at ambient atmosphere. Due to the acidic solution, collagen is present as pro-collagen with single collagen molecules and reactive carboxyl groups on the side chains of the triple helical structure (for details see section 2.2.2 Structure of collagen). These carboxyl groups react with the amino groups present on the functionalized bioactive glass surface through peptide bonds and collagen molecules can attach. SEM images (Figure 27) show the successful loading of bioactive glass-based scaffolds with a collagen monolayer (labeled as dip-coated).

Figure 27: SEM images of collagen-coated bioactive glass-based scaffolds obtained by dip coating at different magnifications showing the inhomogeneous distribution of collagen inside the scaffold structure.

One noticeable feature is the inhomogeneous distribution of collagen inside the coated 45S5 BG-based samples. A thin collagen film with a few nanometers thickness partially covers the surface of the scaffold and blocks the pores though the overall connectivity is not significantly hampered. However, the aim of better handling of the brittle scaffolds could not be achieved by using this coating procedure. In addition, a uniform coating is desired. To optimize this process, another coating method was performed by using the following protocol. Solution A was prepared by mixing 1 M NaOH, 1 M HEPES buffer, 10x concentrated Dulbecco’s Modified Eagle’s Medium (10x DMEM) at a ratio of 1:1:2. Solution A was then mixed with collagen (four parts of 0.5 % collagen solution) to receive solution B. The 0.5 % collagen solution was carefully added to solution A and gently stirred to avoid air bubbles. The pH was monitored with pH-indicator strips. Ideally, the solution showed a pH value between 7.5 and 8. In case of slight differences, the pH was adjusted by adding a few drops of 0.1 M HCl or 0.1 M NaOH. All chemicals were refrigerated (4 – 8 °C) to avoid an early gelation of the neutralized collagen solution B. In addition, the whole process of preparation was carried out in an ice bath. The functionalized 45S5 BG-based scaffolds were placed in a 48-well plate and the neutralized collagen solution B was pipetted on each sample. The samples were incubated over night at 37 °C to initiate and complete the collagen fibrillogenesis. After gelation, the well plate was left uncovered until the samples were

Collagen – As coating material

completely dry. Due to the high shrinkage rate of the collagen gel during the drying process, the collagen fibers contracted and were orientated along the structure of the 45S5 BG-based scaffold (labeled as air-dried). This resulted in a dense fibrous collagen layer wrapped around the struts. SEM images (Figure 28) show a well distributed coating all over the scaffold. Cross-sections exhibit a layer thickness of a few micrometers (compared to a thin monolayer of collagen obtained by dip coating). Altogether, meshed collagen sheets cover the surface of the bioactive glass-based scaffold homogeneously while the scaffold’s macroporosity is not affected. At the interface, the difference between the smooth bioactive glass surface and the fibrous collagen layer can be clearly seen. The handling of the scaffolds was positively improved by the presence of this type of coating. Therefore, the air-dried collagen coating procedure was applied for all further experiments.

Collagen sheet

Fibrous collagen layer

BG surface

Figure 28: SEM images of air-dried collagen-coated bioactive glass-based scaffolds at different magnifications showing a well distributed collagen coating inside the scaffold structure while the macroporosity is not affected.

According to section 3.1.1 (Reinforced porous collagen scaffolds with bioactive glass particles), the effect of incorporated bioactive glass particles in the collagen coating was investigated. Based on the literature175, a concentration ratio of 40:60 (in wt.%) was chosen for 45S5 bioactive glass particles (particle size of  2 µm) and collagen, respectively. To prepare the solution for the collagen coating with incorporated 45S5 BG particles, the protocol for air-dried collagen-coated bioactive glass-based scaffolds was used (as described above). The 45S5 BG particles were dispersed in solution A (1 M NaOH, 1 M HEPES buffer and 10x DMEM at a ratio of 1:1:2) by using an ultrasonic finger. After the particles were well dispersed, the collagen was added (four parts of 0.5 % collagen solution) and the resulting solution was gently mixed. The steps of the coating procedure referred to the one described previously for the air-dried collagen-coated scaffolds.

Collagen – As coating material

Bioactive glass particles

Figure 29: SEM images of collagen-coated bioactive glass-based scaffolds with incorporated 45S5 BG particles at different magnifications showing blogged pores after the coating procedure (left) although the bioactive glass particles are well distributed inside the collagen (right).

SEM images (Figure 29) show the structure of the scaffolds after the collagen coating with incorporated 45S5 BG particles. As it can be seen in Figure 29 (left), the collagen coating with incorporated 45S5 BG particles blocks the pores of the porous scaffold structure completely. Due to the ion release of bioactive glass particles in the collagen solution, the pH changed over time during the coating procedure and, as a consequence of this, also the fibrillogenesis of collagen is affected. Under normal circumstances (pH 7.4, 37 °C) the collagen gelation takes around 30 minutes. It is assumed that the uncontrollable change of pH accelerated the gelling behavior of collagen which led to the clogging of the pores. However, 45S5 BG particles show a fairly homogenous distribution inside the collagen matrix (Figure 29, right). For further experiments, this coating procedure was excluded.

Optimization of parameters – Initial cleaning step

After improvement of the coating procedure, the next step was the modification of the functionalization process. To achieve a uniform deposition of reactive amino groups on the bioactive glass surface, small amounts of water and free hydroxyl groups are required to start the process of functionalization (details explained above). However, due to fast condensation of silanol groups to silica gel on the bioactive glass surface, the presence of free -OH groups is hindered271. Uniformity and reproducibility can be achieved by performing an initial cleaning step to remove organic compounds and other contaminations, thus exposing and activating reactive hydroxyl groups on the surface285. A multitude of cleaning methods are described in the literature using various mixtures of acids, bases and organic solvents271,280,285,286. Based on the work of Verne et al.271, who specifically introduced different methods for cleaning bioactive glasses, the following three methods were chosen and evaluated.

Collagen – As coating material

Method 1 (acetone): Samples were gently swiveled in acetone for 5 minutes followed by rinsing in UPW. This process was repeated three times. Afterwards the samples were dried at 60 °C.

Method 2 (acetone + acid): Samples were gently swiveled in acetone for 5 minutes followed by rinsing in UPW. This process was repeated three times. Further, samples were soaked in 0.1 M sulfuric acid (H2SO4) for 1 minute followed by rinsing in UPW. This process was also repeated three times. Afterwards the samples were dried at 60 °C.

Method 3 (acetone + base): Samples were gently swiveled in acetone for 5 minutes followed by rinsing in UPW. This process was repeated three times. Further, samples were soaked in 0.1 M NaOH solution for 3 minutes followed by rinsing in UPW. This process was also repeated three times. Afterwards the samples were dried at 60 °C.

The effect of the applied cleaning methods can be described as follows. Cleaning the samples in acetone (method 1) may activate hydroxyl groups on the BG surface according to the following reactions271:

(I)

(II)

By reaction with water (I), the ester bonds in the silica network can be cleaved and form two terminals with OH-functionalities. Additionally, Na+ ions, which are capable of being leached out of the 45S5 BG, support this reaction. The pH of the solution shows a low increase and the ion exchange between Na+ and H+ facilitates the formation of -OH on the surface (II). The ion exchange can be stimulated by treatment in acidic solution (III). According to the following reaction271, cleaning method 2 was applied which included soaking in H2SO4.

(III)

Method 3 was carried out in a basic environment. Due to the catalytic effect of the hydroxide ions, a hydroxylation of the silica network may occur, as depicted in the following reactions271 (IV + V):

(IV)

Collagen – As coating material

(V)

Considering the literature271, the success of the applied cleaning methods was evaluated by measuring the contact angle. In this case, a flat surface was required and therefore bioactive glass pellets were used. The results of the contact angle measurement are represented in Figure 30. Pure 45S5 bioactive glass samples without any treatment showed a contact angle of 15 ± 2 °. After cleaning in acetone, the angle decreased to 7 ± 2 °. Further treatment in NaOH did not show any improvement. Greatest success could be achieved using cleaning method 3 which included further soaking in H2SO4. Previously, it was suggested that the concentration of -OH groups on materials surfaces is closely related to the hydrophilicity of the material287. Because of the electronegativity of the oxygen atom in -OH groups, the functional group becomes polar. The polar H2O molecule is attracted by this polarity and a hydrophilic effect can be observed. Taking all results into account, a suitable way of cleaning could be the procedure in acidic solution.

20 Untreated Acetone

Acetone + H2SO4 Acetone + NaOH

15 

10 Contactangle 5

Figure 30: Contact angle  of 45S5 bioactive glass surfaces before and after cleaning (using the different methods described in the text).

However, the surface of the samples was additionally investigated by SEM observation after the treatments which show the negative effect of the applied acid treatment on the 45S5 BG surfaces (Figure 31).

SEM images of the bioactive glass samples cleaned in 0.1 M H2SO4 revealed a porous and etched surface. Soaking the samples in acidic solution attacked the surface to such an extent that this cleaning method was excluded. Accordingly, based on all results of this part of the study, all samples were cleaned only in acetone before functionalization.

Collagen – As coating material

Untreated Acetone Acetone + NaOH Acetone + H2SO4

Figure 31: SEM images of sintered bioactive glass surfaces (pellets) before and after applying different cleaning methods.

Optimization of parameters – Solvent and immersion time

As already mentioned, the silanization process can be affected by many parameters including applied solvent, time and APTS concentration amongst others. Preliminary, 45S5 BG samples were functionalized by adapting the protocol of Chen et al.247 who showed a successful silanization of bioactive glass surfaces using deionized water as solvent. However, studying the literature, a higher potential for protein loading by the use of organic solvents during the process of functionalization has been reported269. Based on those previous studies, the way of functionalization for 45S5 BG was modified by changing solvent and immersion time. To analyze the impact of the different parameters, X-ray photoelectron spectroscopy (XPS) was chosen for further characterization. All experiments were carried out with 45S5 bioactive glass pellets which were cleaned in acetone before functionalization (based on the results of the initial cleaning process).

Solvent Time Treatment  UPW  5 min  Acetone + 2 vol.%  1 h Rinse in 2 h at  Ethanol (99.8 vol.%) APTS  4 h UPW 100 °C  Toluene  6 h

Figure 32: Optimization of applied solvents. 45S5 BG pellets were immersed in different solvents and removed after varying time points. The efficiency of the functionalization process was assessed by XPS.

The influence of different solvents and immersion times on the efficiency of the functionalization process was checked first. Therefore, the process of functionalization was performed in UPW281, acetone288, ethanol271 and toluene285. Solutions were prepared with a concentration of 2 vol.% APTS. The samples were immersed for 5 minutes271, 1 hour285, 4 hours281 and 6 hours271, respectively. After each time point, samples were removed, rinsed in UPW two times and immediately heat treated at 100 °C for 2 h to consolidate the bonding between the silane layer and the surface (section 3.2.1 Relevance of functionalization). An overview of the applied methods is given in Figure 32. After heat treatment, the pellets were examined via XPS to observe the chemical modification during the different treatments. By determination of the atomic concentration of the elements on the surface, the

Collagen – As coating material functionalization process was estimated. To confirm the presence of APTS molecules on the bioactive glass surface, different markers can be applied. Nitrogen (N) and carbon (C) were considered to be adequate candidates because this elements are not contained in the 45S5 BG substrate281. However, carbon can also represent a source of contamination, so the nitrogen concentration was chosen as respective marker. The results of the XPS measurements are represented in Table 7. To compare the results, 45S5 bioactive glass pellets without APTS treatment were used as reference.

Table 7: XPS results showing the different atomic concentration of elements present on the surface after the functionalization process. As reference, 45S5 BG pellets cleaned in acetone but without APTS treatment were used.

Atomic concentration Sample Time N1s Si2p Ca2p Na1s P2p Reference 0 min 0.01 6.58 5.83 24.68 4.25 5 min 0.85 7.13 2.58 23.17 3.61 1 h 1.85 7.78 1.87 5.69 1.66 UPW 4 h 1.81 8.27 0.43 2.11 0.99 6 h 1.95 8.83 0.92 0.98 0.19 5 min 1.25 7.37 3.43 11.04 4.00 1 h 3.29 11.36 3.07 9.42 2.54 Acetone 4 h 3.40 11.57 2.51 8.26 1.99 6 h 3.47 11.65 1.79 7.73 1.48 5 min 1.15 8.41 3.56 15.12 3.00 1 h 1.63 9.17 3.43 12.65 2.77 Ethanol 4 h 1.81 9.29 3.39 10.53 2.59 6 h 1.83 9.23 3.30 9.84 2.31 5 min 1.17 11.05 3.83 9.69 2.73 1 h 3.59 11.43 2.57 9.51 2.02 Toluene 4 h 3.69 11.51 1.76 9.10 1.81 6h 3.81 11.59 0.98 9.03 1.19

After functionalization, all samples showed an increasing amount of nitrogen which was already stable after 1 hour and only changed slightly with longer immersion times. The increasing nitrogen concentration indicates the successful attachment of APTS molecules on the surface. This thesis could be supported by the increasing amount of silica. Concentrations of calcium (Ca), sodium (Na) and phosphorus (P), which are part of the 45S5 BG substrate, were continuously depleted which indicates the dissolution of the 45S5 bioactive glass. It is interesting to confirm the silanization process was successful on every surface independently of the applied solvent. Comparing the values of nitrogen and silica, samples treated in water showed a poor functionalization of the sample surface. In addition, the dissolution rate is seen to be relatively high. Treatment in ethanol slowed down the dissolution of the 45S5 BG but the attachment of APTS molecules was not enhanced. In addition, high amounts of nitrogen could be detected using acetone or toluene as solvent during the functionalization process. Although toluene reveals better results than acetone, it was excluded due to its higher hazardousness. Thus

Collagen – As coating material

acetone, which shows similar results, was the solvent of choice for further experiments. Concerning the progressive dissolution of the bioactive glass substrate, the immersion time was chosen to be 1 h. For clarification, Figure 33 shows the XPS spectra of 45S5 BG samples before and after the functionalization process in acetone. A remarkable feature is the reduced intensity of the peaks after the silanization which indicates the fast dissolution process on the surface of the 45S5 BG pellets during the immersion process. Moreover, because of the reaction of free -OH groups on the surface of bioactive glass with APTS molecules, the peak at 533 eV is seen to decrease after the functionalization process. In return, a peak at around 400 eV appears confirming the presence of N-containing groups (amino groups of APTS molecules) on the surface.

400000 O 1s O Na 1s Na before functionalization after functionalization (1 h in acetone)

350000 KLL O Na KLL Na

300000

250000 Si2p

200000

N 1s N c/s

150000 2p P

P 2s P

Ca 2p3 - Ca 2p1 - Ca 2p Ca 2p1 2p3Ca Ca - - C 1s C

100000 2s Ca Si2s 50000

1200 1000 800 600 400 200 Binding energy [eV]

Figure 33: XPS spectra of 45S5 bioactive glass surfaces before and after the silanization in acetone (+ 2 vol.% APTS) for 1 h showing a reduced intensity of peaks after the functionalization process indicating a progressing dissolution of the 45S5 BG substrate. In return, a peak at 400 eV appears confirming the presence of amino groups on the surface. Relevant peaks are marked and further discussed in the text.

In consideration of all results so far, the manufacturing process of collagen-coated bioactive glass-based scaffolds, which were used for further investigations, can be summarized as follows and is illustrated in a flow chart (Figure 34). 45S5 bioactive glass-based scaffolds were produced by the foam replica method29. Cylindrical shaped PU samples (Ø = 12 mm, h = 7 mm) were initially cleaned in 70 vol.% of 2-propanol and immersed into a slurry of 45S5 BG particles, UPW, disperser and binder in order to obtain green bodies. Two layers of BG coating were applied and subsequently sintered at 1050 °C for 2 h. Bioactive glass pellets were prepared by compaction of 45S5 BG powder and sintered under the same conditions of the scaffolds. Before the samples were coated with collagen, an initial cleaning step was introduced which was carried out in acetone and UPW

Collagen – As coating material

(3 times, 5 minutes). After drying at 60 °C, the scaffolds were soaked in a solution of acetone with 2 vol.% of APTS for surface functionalization. After 1 hour of immersion, samples were removed, rinsed two times in fresh UPW and heat treated at 100 °C for 2 h to consolidate the bondings. The functionalized 45S5 BG samples were then placed in a 48-well plate. By mixing 1 M NaOH, 1 M HEPES, 10x DMEM and 0.5 % solution of collagen, a neutralized collagen solution was prepared and added to each sample. After incubation overnight, which initiated and completed the gelation process, the well plate was left uncovered until the samples were completely dry.

Cylindrical specimens of Manufacturing of Slurry: Sintering: polyurethane foam 45S5 BG, UPW, 2 h at 1050 °C 45S5 BG-based (cleaning in 70 vol.% 2- disperser and binder and natural cooling scaffolds propanol )

Functionalization of Initial cleaning in Functionalization: Heat treatment: acetone and UPW, Acetone + 2 vol.% 2 h at 100 °C 45S5 BG surfaces drying at 60 °C APTS for 1 hour and natural cooling

Collagen mixture: Functionalized scaffold Collagen coating on Incubation overnight at 1 M NaOH placed in 48-well plate 1 M HEPES 37 °C, afterwards 45S5 BG-based + 500 µl neutralized 10x DMEM drying uncovered at RT scaffolds 0.5 % Collagen solution collagen solution

Manufacturing of 45S5 BG powder Sintering: compressed 2 h at 1050 °C 45S5 BG pellets with 104 N and natural cooling

Figure 34: Flow chart explaining the manufacturing process for uncoated and collagen-coated 45S5 bioactive glass-based scaffolds.

Crosslinking process

As already described (section 3.1.1 Crosslinking process), additional crosslinks can be introduced by a chemical reaction. According to the protocol described there, crosslinked samples (cl) were prepared by treatment with EDC and NHS. Therefore, 60 mM of EDC and 60 mM of NHS were dissolved in 50 mM MES acid solution. Collagen-coated bioactive glass-based scaffolds were immersed in the solution for 4 h at room temperature.

Afterwards, they were removed and washed in 0.1 M Na2HPO4 for 2 h to hydrolyze the remaining O- acylisourea of the carbodiimide. Subsequently, scaffolds were washed in UPW and left for drying at 37 °C. Uncrosslinked samples are labeled as uc. Figure 35 shows the FTIR spectrum of 45S5 BG-based scaffolds without any coating. For comparison, the spectra of collagen coatings on the surface of these scaffolds are illustrated before and after the crosslinking process. The most relevant peaks are labeled.

Collagen – As coating material

45S5 BG-based scaffold + Collagen (cl) 45S5 BG-based scaffold + Collagen (uc)

45S5 BG-based scaffolds

-1

532 cm 532 Si-O-Sibending

621 cm 621 Si-O-Sibending

-1

926 cm 926 Si-O-2NBOstretching

455 cm 455 Si-O-Sibending

1658 - 1647 cm 1647 - 1658 I Amide

-1

1100 - 1000 cm 1000 - 1100 stretching Si-O-SiP-O stretching,

-1

1221 cm 1221 III Amide

1568 cm 1568 II Amide

Transmission[a.u.]

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 35: FTIR spectra of 45S5 BG-based scaffolds, collagen-coated 45S5 BG-based scaffolds before and after crosslinking showing the formation of additional peaks (amide I, II and III) after the collagen coating confirming the presence of the protein. The identified peaks are discussed in the text.

Pure 45S5 bioactive glass shows typical peaks at 455 cm-1 and in the region between 1100 cm-1 and 1000 cm-1. These peaks can be attributed to the Si-O-Si vibrational modes in the glass network233,289: Si-O-Si bending vibration at 455 cm-1 and Si-O-Si stretching vibration at around 1088 cm-1. The region between 1100 cm-1 and 1000 cm-1 is difficult to distinguish due to the overlap of Si-O-Si stretching and P-O stretching175. According to Filho et al.290, the spectral area between 450 cm-1 and 625 cm-1 is of special interest because the sharp peaks at 455 cm-1, 532 cm-1 and 621 cm-1 indicate the high crystallinity of the 45S5 bioactive glass-based scaffolds attributed to the Si-O-Si vibrational modes of the developed crystal phase (Si-O bending). The peak, visible at 926 cm-1, can be assigned to the vibrational mode of the Si-O-2NBO (non-bridging oxygen) bond175,233. The non-bridging oxygen results from the formation of SiO- groups in the glass network233,289,291 due to the cations of the network modifiers (Ca2+ and Na+, section 2.2.1 45S5 Bioglass® - a silicate-based bioactive glass). The spectra of the collagen-coated 45S5 BG-based scaffolds exhibit the characteristic peaks related to the bioactive glass phase. In addition, the presence of the collagen coating can be confirmed by the formation of additional peaks in the area between 1700 cm-1 and 1200 cm-1. These peaks are namely amide I (C=O stretching vibrations175,222–226) in the region between 1647 cm-1 and 1658 cm-1, which are typical for the triple helical structure in non-denatured collagen175, and amide III (1221 cm-1) which belongs to N-H deformation175,227. After the crosslinking process, an additional peak at 1568 cm-1 appears (related to the N-H deformation and C-N stretching224,225,227) which confirms the extra amide-type bond (amide II). For further characterization of uncrosslinked and crosslinked collagen-coated samples, thermogravimetric analysis (TGA) was carried out in order to quantify the amount of collagen before and after the crosslinking

Collagen – As coating material process. Heating was performed in ambient atmosphere with a heating rate of 5 K/min. Temperature and weight loss were continuously recorded and revealed the following curve progressions (Figure 36). Collagen shows a continuous mass loss which can be divided into three different stages292. Stage I, which is in the range between room temperature and 200 °C, can be attributed to the evaporation of water molecules. The mass loss during the second stage (200 - 450 °C) can be associated with the decomposition of collagen molecules. Stage III (450 - 600 °C) exhibits the mass loss during the pyrolysis of residual organic components. Considering the mass loss of collagen between 200 and 450 °C, the collagen amount of uncrosslinked samples can be determined as  6.5 wt.% whereas only  2 wt.% of collagen is left after the crosslinking process. During the crosslinking process, uncrosslinked collagen is released and only collagen fibrils, which are fixed in the crosslinked network, remain. This means, as a corollary, a crosslinking degree of  30 %.

-2 I.

-4 II. -6

-8 III.

Mass loss [%] loss Mass -10

-12

-14 45S5 BG-based scaffolds coated with collagen (crosslinked) 45S5 BG-based scaffolds coated with collagen (uncrosslinked)

0 100 200 300 400 500 600 Temperature [°C]

Figure 36: Mass loss of uncrosslinked and crosslinked collagen on 45S5 BG-based scaffolds during TGA in ambient atmosphere (heating rate 5 K/min).

If the degree of crosslinking is compared with results achieved in section 3.1.1 Crosslinking process, the reduction from  50 % for pure collagen compared to  30 % for collagen coating is obvious. Because the carboxyl groups, which are part of the crosslinked network, are already involved in the functionalization process, the degree of crosslinking is reduced.

Collagen – As coating material

3.2.2 Swelling properties and release behavior

One important aspect regarding coatings in general is their stability against release and degradation. Coatings are applied to fulfill several requirements like enhanced mechanical properties, drug delivery capability, stimulation of cell attachment or improved bioactivity. To ensure the effectiveness of collagen (section 2.2.2 Collagen for tissue engineering) as coating, the importance is given to analyze its stability on 45S5 BG-based scaffolds. The release behavior of collagen can occur by two mechanisms. The physical mechanism of release is ruled by dissolution and/or diffusion. The chemical mechanism is based on degradation which can be chemical or enzymatic.

Swelling properties

To determine the difference in water binding ability of collagen-coated 45S5 BG-based scaffolds with and without crosslinks, samples were soaked in PBS for up to 24 h at room temperature, as already described before (section 3.1.3 Swelling properties and degradation behavior). Samples were removed and excess water was dried with a filter paper (qualitative filter paper 410). Water uptake was calculated according to equation [7].

60 min 30 min 600 15 min

500

400

300

200 Relative[%] water uptake 100

0 Collagen uncrosslinked Collagen crosslinked

Figure 37: Results of the swelling study on collagen-coated samples: relative water uptake of collagen-coated BG-based scaffolds with and without crosslinking in PBS after 1 h.

Results revealed that equilibrium was reached after 1 h of immersion in PBS. Therefore, Figure 37 shows the relative water uptake of 45SS bioactive glass-based scaffolds coated with collagen after 1 h in PBS. The diagram shows the limited swelling behavior of collagen before the crosslinking process (310 ± 70 %) compared to

Collagen – As coating material crosslinked samples (520 ± 100 %). However, the literature describes a higher potential of water absorption for uncrosslinked collagen219,293. Normally, smaller distances between collagen molecules and reduced/occupied binding sites for water by introduced crosslinks explain the limited swelling behavior after the crosslinking process. This fact can be visualized by SEM images of the collagen-coated samples before and after crosslinking (Figure 38). The opposite behavior, which was observed during this study, can be explained by considering that even if the swelling behavior of uncrosslinked collagen is higher during the immersion in PBS, collagen is released very fast due to missing crosslinks and as a result a high amount of collagen is lost over time.

Figure 38: SEM images of collagen coating on 45S5 bioactive glass-based scaffolds before (A) and after (B) crosslinking showing the different network structures of collagen before and after crosslinking. .

Release behavior

The release behavior of the collagen-coated bioactive glass-based scaffolds was evaluated in different media to investigate the release kinetics independent from the formation of hydroxyapatite (immersion in PBS) and as a function of the HA formation (immersion in SBF). For this study, a selection of samples (Table 8) was made.

Table 8: Selected scaffolds. Collagen release was determined depending on the cleaning and functionalization process.

Sample Cleaning Functionalization Collagen Crosslinking  (A) - - + -  (B) + - + -  (C) - + + -  (D) + + + -

Collagen – As coating material

Chosen samples were immersed in 5 ml of phosphate buffered saline or simulated body fluid (SBF, see section 3.1.4 for details of preparation), respectively, and incubated in an orbital shaker247 at 37 °C. After varying time points, the solution was completely removed and replaced by fresh one. The released collagen amount in the removed solution was determined colorimetrically from a standard calibration curve (collagen concentration vs. absorbance), based on a modification of Lo ry’s method294,295. The Lowry method consists of two chemical reactions. First, the biuret reaction296 takes place where complex compounds between the protein and an alkaline copper tartrate are formed. In a second step, Folin & Ciocalteu’s Phenol reagent is added and reduced hich yields a blue color. ased on the manufacture’s instructions297 (L3540, Sigma-Aldrich), the following protocol was applied. Lowry reagent, which was received as powder, was mixed with 40 ml of UPW to get the Lo ry reagent solution. The Folin & Ciocalteu’s Phenol reagent working solution was prepared by mixing 18 ml of 2 N Folin & Ciocalteu’s Phenol reagent with 90 ml of UPW. The solution required storage at room temperature in an amber bottle due to light sensitivity. For analysis, 250 µl of protein solution (PBS or SBF) was transferred to a 1 ml cuvette, respectively. 250 µl of Lowry reagent solution was added and incubated at room temperature. After 20 minutes, 125 µl of Folin & Ciocalteu’s Phenol reagent working solution was added and incubated for additionally 30 minutes to allow the color to develop. According to the application, the solution was irradiated by visible light with wavelength of 595 nm using a single beam spectrometer (Specord 40, Analytik Jena, DE) and absorbance values were recorded. For the determination of the collagen concentration, a calibration curve was applied (known collagen concentration vs. absorbance). First, collagen release was determined as a function of the cleaning and functionalization process. The time period ranged from 1 h to 28 days. Pure 45S5 BG-based scaffolds without further treatment or collagen coating were used as blank sample. For statistical evaluation of the release rate, samples were tested in triplicate. The results of the cumulative collagen release in different media are illustrated and described in the following figures.

(A) uncleaned 45S5 BG-based scaffold, no functionalization, uc (A) uncleaned 45S5 BG-based scaffold, no functionalization, uc (B) Acetone cleaned 45S5 BG-based scaffold, no functionalization, uc (B) Acetone cleaned 45S5 BG-based scaffold, no functionalization, uc (C) uncleaned 45S5 BG-based scaffold, APTS, uc 11.0 (C) uncleaned 45S5 BG-based scaffold, APTS, uc (D) Acetone cleaned 45S5 BG-based scaffold, APTS, uc (D) Acetone cleaned 45S5 BG-based scaffold, APTS, uc Pure PBS 100 10.5

10.0

80 9.5

9.0

60 pH 8.5

8.0

40 7.5 Cumulative[%] release collagen 7.0 in PBS in PBS 20 0 100 200 300 400 500 600 700 0 100 200 300 400 500 600 700 Time [h] Time [h]

Figure 39: Cumulative collagen release from different 45S5 BG-based scaffolds in PBS (left) and pH values during release (right).

Collagen – As coating material

Figure 39 shows the different curves of the cumulative collagen release in PBS for up to 28 days. In parallel, the pH values during the release were measured. All samples exhibit a similar shape of the curves which show an initial burst release followed by a plateau. The curves differ only in the amount of released collagen. This phenomenon can be explained as follow: during the coating procedure, a collagen mesh consisting of single collagen fibers is wrapped around the struts of the 45S5 BG-based scaffold (Figure 28). In case of unfunctionalized samples (A+B in Figure 39), the collagen layer is only physically adsorbed on the surface due to missing binding sites. Additionally, the collagen fibers do not have any crosslinks. It is supposed that water penetrates the collagen mesh very fast which leads to swelling of the fibrous structure (section 3.2.2 Swelling properties). Macromolecules are released over time and can diffuse. This causes the fast collagen release from collagen-coated samples without any surface treatment. Around 90 % of the collagen is already released in the first 24 h. An initial cleaning step is seen to have slowed down the release in the beginning but could not prevent the high release rate during the short period of time. After 28 days of immersion, only  10 % of collagen is left on the surface of the scaffolds which requires enzymatic degradation for complete release. By introducing a silanized surface, collagen can bind covalently to the bioactive glass surface (see section 3.2.1 Surface functionalization) and the release rate is seen to have slowed down (C+D Figure 39). Similar results were described by Chen et al247. Despite the functionalization, the release profile shows an initial burst release up to 50 % because the functionalization process only affects collagen close to the surface. Collagen molecules, which are only adsorbed and not bonded to the BG surface, still diffuse and are released very fast. After a fast release of adsorbed collagen in the beginning, a plateau is reached at 60 % for samples without an initial cleaning step and 54 % is reached for initially cleaned samples. Conversely, results showed that by introducing a cleaning step, the amount of covalently bonded collagen could be increased from 40 % to 46 % due to more reactive -OH groups on the surface.

(A) uncleaned 45S5 BG-based scaffold, no functionalization, uc (B) Acetone cleaned 45S5 BG-based scaffold, no functionalization, uc (A) uncleaned 45S5 BG-based scaffold, no functionalization, uc (C) uncleaned 45S5 BG-based scaffold, APTS, uc (B) Acetone cleaned 45S5 BG-based scaffold, no functionalization, uc (D) Acetone cleaned 45S5 BG-based scaffold, APTS, uc (C) uncleaned 45S5 BG-based scaffold, APTS, uc (D) Acetone cleaned 45S5 BG-based scaffold, APTS, uc 60 8.4 Pure SBF in 10mM PBS

8.2

8.0

7.8 pH

7.6

7.4 Cumulative[%] release collagen 20 in SBF 7.2 in SBF 0 100 200 300 400 500 600 700 0 100 200 300 400 500 600 700 Time [h] Time [h]

Figure 40: Cumulative collagen release from different 45S5 BG-based scaffolds in SBF (left) and pH values during release (right).

Collagen – As coating material

Figure 40 shows the cumulative collagen release in SBF over time and the corresponding pH values. The trend of collagen release kinetics is similar to the one in PBS (Figure 39), although the overall amount is remarkable lower. Even samples, which did not receive any surface treatment, released less than 60 % of the initial collagen amount. Because of the immersion in SBF, nodules of hydroxyapatite are expected to grow on the surface becoming embedded in the collagen matrix. As a result, collagen starts to mineralize and the release rate is reduced (details about HA formation are presented in section 3.2.3 Evaluation of bioactivity). In addition, pH values show an increment from 7.0 up to 10.0 in PBS and from 7.4 up to 8.4 in SBF, respectively, which can be attributed to the fast ion release of the 45S5 bioactive glass surface, as described also by Hoppe et al.126. Conditions for optimal collagen attachment are around pH  7.4, as discussed previously. The basic environment influences the stability of collagen235,247,298,299 and represents a second reason for the fast initial collagen release. In a second step, collagen release was evaluated depending on the degree of crosslinking. Selected samples are presented in Table 9.

Table 9: Selected scaffolds. Collagen release was determined depending on the degree of crosslinking. 672 336 168 Sample72 Cleaning Functionalization Collagen Crosslinking 24  (D) + + + -

6  (E) + + + +

To ascertain the release behavior, collagen-coated 45S5 bioactive glass-based scaffolds were immersed in PBS and SBF before and after the crosslinking process, respectively. The release kinetics is illustrated in Figure 41.

60 40

55 35

50 30 45 Acetone cleaned 45S5 BG-based scaffold, APTS, uc Acetone cleaned 45S5 BG-based scaffold, APTS, cl 25 40

20 35 Acetone cleaned 45S5 BG-based scaffold, APTS, uc Acetone cleaned 45S5 BG-based scaffold, APTS, cl 15

Cumulative[%] release collagen 30 Cumulative[%] release collagen

in PBS in SBF 25 10 0 100 200 300 400 500 600 700 0 100 200 300 400 500 600 700

Time [h] Time [h]

Figure 41: Collagen release kinetics of collagen-coated 45S5 BG-based scaffolds in PBS (left) and SBF (right) before and after crosslinking.

As expected, the cumulative collagen release of around 54 % in PBS and 32 % in SBF, respectively, could be further decreased to 36 % and 26 % (Figure 41), respectively, by a crosslinking process. In addition to covalently bonded collagen on the 45S5 bioactive glass surface, the collagen network is crosslinked which slows

Collagen – As coating material down the release rate. In conclusion, the release rate of collagen can be easily controlled by introducing a silanized surface in combination with a crosslinked collagen coating, depending on the application.

3.2.3 Evaluation of bioactivity

As already described above, the bioactive behavior is an important aspect describing the bone-bonding ability of biomaterials and can also explain the release behavior of collagen. According to section 3.1.4 Evaluation of bioactivity, samples were conditioned in SBF for up to 10 days. After different time points (1, 3, 7 and 10 days), samples were removed and analyzed in terms of hydroxyapatite formation on their surfaces. For optical characterization of the biomineralization process, SEM was used and the chemical composition was confirmed by FTIR measurements (summarized in Figure 42-47). Figure 42 reveals the typical surface morphology of 45S5 BG-based scaffolds after 1, 3 and 7 days of immersion in SBF. After 1 day, calcium phosphate precipitates cover the surface homogeneously which is qualitative assessed by the typical morphology (cauliflower-like21,250). The characteristic shape of HA can be recognized after 3 days in SBF. At higher magnification, hydroxyapatite can be clearly identified by its cauliflower structure with a particle size of  2 µm after 3 days and  3.5 µm after 7 days. With longer immersion time in SBF, particles grew and the entire surface is covered by a dense apatite layer. The FTIR spectra, which can be seen in Figure 43, confirm the formation of hydroxyapatite on the scaffold surface during immersion in SBF.

45S5 BG - 1 d 45S5 BG - 3 d 45S5 BG - 7 d

45S5 BG - 1 d 45S5 BG - 3 d 45S5 BG - 7 d Typical cauliflower structure

Typical cauliflower structure

Figure 42: SEM images of as-fabricated 45S5 BG-based samples after the immersion in SBF for 1, 3 and 7 days (at different magnifications) showing the formation of hydroxyapatite on the surface (qualitative assessed by the typical cauliflower-like structure).

The FTIR spectra of as-fabricated 45S5 bioactive glass-based scaffolds (Figure 43), recorded in transmission

Collagen – As coating material

mode, display the changes occurred during immersion in SBF. With increasing immersion time, conspicuous peaks at 532 cm-1, 621 cm-1 and 926 cm-1 are reduced in their intensity and disappear after 10 days which implies changes in the composition. The peak at 455 cm-1 is slightly shifted due to the presence of cations291. In addition, significant new peaks appear. A growing double peak is detected at 570 cm-1 and 600 cm-1 which can

3- 253,254,300 be attributed to bending vibrations of P-O related to the PO4 group in crystalline HA . The growing broad peak between 1100 cm-1 and 1000 cm-1 and the small shoulder at 960 cm-1 can be related to the P-O stretching in the phosphate group253,301,302. The peak at 800 cm-1 is related to the formation of a silica rich layer291. Peaks, which are developed at 875 cm-1 and in the region 1500 - 1400 cm-1, can be attributed to

2- 253,254 C-O bending and stretching vibrations of CO3 groups in the carbonated HA layer , respectively.

-1 -1

570 / 600 cm-1

-1 P-O stretching P-O

960 cm 960 P-O bending

-1 -1

875 cm 875 bending C-O

-1

Si-O-Si

800 cm 800 

10 d

1400-1500 cm 1400-1500 cm 1500-1400 bending P-O stretching C-O

1100-1000 cm 1100-1000 stretching P-O

7 d

3 d

1 d

0 d

Transmission[a.u.]

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 43: FTIR spectra of as-fabricated 45S5 BG-based samples after 0, 1, 3, 7 and 10 days of immersion in SBF showing additional peaks at 960 cm-1, 875 cm-1 and 800 cm-1, indicating the formation of HA. The identified peaks are discussed in the text.

Figure 44 shows SEM images demonstrating the bioactive behavior of collagen-coated 45S5 BG-based scaffolds without crosslinking (uc). The images show the growing of hydroxyapatite nodules in form of microcrystals along the collagen fibers after 1 day of immersion in SBF. With increasing immersion time, the formation of hydroxyapatite becomes clearer and the density of precipitates increases. Large HA particles are directly deposited on the fibrous structure and are well embedded inside the collagen matrix. After 10 days, the collagen layer is seen to be mostly mineralized. In addition, HA crystals form thin sheets and cover the surface of the scaffold. These findings are in accordance with the literature248,303 which describe the existence of nucleation sites in collagen for the formation of hydroxyapatite. Such nucleation sites are e.g. –COOH groups which interact with Ca2+ ions304,305. This interaction is also visible in the FTIR spectra (Figure 46 and 47) due to

Collagen – As coating material the shifted peak of amide I and the disappearance of amide II and amide III peaks after the immersion in SBF. This result explains the excellent bioactive behavior of the collagen coating. Same results were obtained in case of crosslinked (cl) collagen-coated samples (Figure 45).

45S5 BG + Coll (uc) - 1 d 45S5 BG + Coll (uc) - 7 d 45S5 BG + Coll (uc) - 10 d

HA sheets

45S5 BG + Coll (uc) - 3 d 45S5 BG + Coll (uc) - 7 d 45S5 BG + Coll (uc) - 10 d

HA sheets

Figure 44: SEM images of collagen-coated 45S5 BG-based scaffolds (uc) after the immersion in SBF for 1, 3, 7 and 10 days (at different magnifications) show the growing of hydroxyapatite nodules in form of microcrystals along the collagen fibers after 1 day of immersion. After 10 days, the collagen layer is mostly mineralized.

45S5 BG + Coll (cl) - 1 d 45S5 BG + Coll (cl) - 7 d 45S5 BG + Coll (cl) - 10 d

HA sheets

45S5 BG + Coll (cl) - 3 d 45S5 BG + Coll (cl) - 7 d 45S5 BG + Coll (cl) - 10 d

Figure 45: SEM images of collagen-coated 45S5 BG-based scaffolds (cl) after the immersion in SBF for 1, 3, 7 and 10 days (at different magnifications) show the growing of hydroxyapatite nodules in form of microcrystals along the collagen fibers after 1 day of immersion. After 10 days, the collagen layer is mostly mineralized. Crosslinking does not affect the bioactive behavior of collagen.

Collagen – As coating material

-1 -1

10 d

-1

Si-O-Si

-1

875 cm 875 bending C-O

-1

800 cm 800 

-1 570 / 600 cm-1

1643 cm 1643 I Amide P-O stretching P-O

960 cm 960 P-O bending Amide III Amide

7 d cm 1221

1500-1400 cm 1500-1400 stretching C-O

1110-1020 cm 1110-1020 stretching P-O

3 d

1 d

Transmission[a.u.]

0 d

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 46: FTIR spectra of collagen-coated 45S5 BG-based scaffolds (uncrosslinked) after 0, 1, 3, 7 and 10 days of immersion in SBF showing additional peaks at 960 cm-1, 875 cm-1 and 800 cm-1, indicating the formation of HA. The identified peaks are discussed in the text.

-1

-1 -1

10 d -1 C-O bending C-O

875 cm 875 -1

-1 Si-O-Si

-1 570 / 600 cm

800 cm 800 

1662 cm 1662 I Amide

1226 cm 1226 III Amide Amide II Amide

1568 cm 1568 P-O bending

960 cm 960 stretching P-O

1500-1400 cm 1500-1400 stretching C-O P-O stretching P-O

7 d cm 1110-1020

3 d

1 d

Transmission[a.u.]

0 d

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 47: FTIR spectra of collagen-coated 45S5 BG-based scaffolds (crosslinked) after 0, 1, 3, 7 and 10 days of immersion in SBF showing additional peaks at 960 cm-1, 875 cm-1 and 800 cm-1, indicating the formation of HA. The identified peaks are discussed in the text.

Collagen – As coating material

FTIR spectra, which were obtained from the collagen-coated 45 BG-based samples (with and without crosslinking) after immersion in SBF, evidenced the formation of hydroxyapatite. Figure 46 and 47 revealed “smoothed” spectra after immersion in S F suggesting the progressive change in composition. Peaks at 528 cm-1, 621 cm-1 and 932 cm-1 vanish after day 1 indicating the faster formation of hydroxyapatite compared to uncoated samples. Similar to the FTIR spectra of as-fabricated 45S5 BG-based samples (Figure 43), characteristic peaks appear: the double peak at 570 cm-1 and 600 cm-1, peaks at 800 cm-1, 875 cm-1and in the region 1500 -1400 cm-1 which have been explained above. Furthermore, a broad peak located between 1110 cm-1 and 1020 cm-1 and a small shoulder at around 960 cm-1 can be recognized which

3- 253,301,302 are attributed to the stretching mode of the phosphate phase (PO4 ) in the HA layer . In summary, collagen coating was shown to enhance the bioactive behavior due to the nucleation sites in collagen.

3.2.4 Mechanical characterization

In order to investigate the potential of collagen-coated 45S5 BG-based scaffolds for the application in bone tissue engineering, the compressive strength σ of different scaffolds was determined.

45S5 BG-based scaffolds, collagen-coated (uc) 45S5 BG-based scaffolds, collagen-coated (cl) 45S5 BG-based scaffolds, collagen-coated (cl - dry) 45S5 BG-based scaffolds, collagen-coated (cl - wet) II. 45S5 BG-based scaffolds, uncoated 0.2 0.2 I. III.

0.1 0.1

Compressive stress [MPa] Compressive Compressive stress [MPa] Compressive

1 2 1 2 Displacement [mm] Displacement [mm]

Figure 48: Stress-displacement curves in comparison for 45S5 BG-based scaffolds with and without collagen coating. Left: curves of uncoated and collagen-coated BG-based scaffolds (uc and cl); right: curves of collagen-coated BG-based scaffolds (cl) in dry and wet state.

Therefore, uniaxial compression tests were performed with a crosshead velocity of 1 mm/min, a preload of 0.1 N and maximum applied force of 50 N (Zwick Z050, DE). Five specimens of each series were uniaxial compressed until densification of the porous structure occurred. Figure 48 shows typical stress-displacement curves of uncoated and collagen-coated 45S5 BG-based scaffolds. In order to simulate the environment in vitro, crosslinked collagen-coated BG-based scaffolds were additionally immersed in PBS before measurement. After 1 h of rehydration, stress-displacement curves were recorded and compared with the ones of dry samples (Figure 48, right). The typical stress-displacement curves, which were recorded during the measurements,

Collagen – As coating material

show three different regions29 and can be explained as follows. In region I, the compressive stress constantly increases until the scaffold struts are fractured and maximum stress is reached. As a result, the stress- displacement curve shows a negative slope in region II. Due to progressive compression, the broken structure is densified and the compressive stress increases again in region III. This behavior is typical for compressed foams and has been described in the literature before29. Due to the hollo nature of the struts and the presence of microcrac s on the struts’ surfaces, which are a result of the crystallization process during sintering29, the curve of uncoated bioactive glass-based scaffolds exhibits a jagged behavior (Figure 48) and low values (0.04 ± 0.02 MPa) were reached for the compressive strength σ. In the literature, values around 0.3 – 0.4 MPa are reported for scaffolds with porosity ~ 90 %. Due to collagen coating, the mechanical strength could be enhanced by a factor of 5. Compressive strength values of 0.21 ± 0.03 MPa and 0.18 ± 0.03 MPa were achieved for uncrosslinked collagen- and for crosslinked collagen-coated scaffolds, respectively. It is supposed that by further optimization of the 45S5 BG-based scaffolds (e.g. multiple layers of 45S5 BG during the fabrication process), higher values can be reached. The literature reports compressive strength of cancellous bone in the range of 2 – 12 MPa51,306. However, also lower values are reported (0.2 – 4 MPa) when the relative density of bone tissue is ~ 0.129, thus coated scaffolds are still interesting for application in bone tissue engineering. Ingrowth of tissue in vivo will provide additional mechanical support which was reported in case of porous hydroxyapatite scaffolds307 (compressive strength increased from 10 to 30 MPa). ased on the fact that literature reported Young’s moduli for collagen in the range 0.2 – 21.5 GPa308, the present results are below expectations. Nevertheless, collagen coating enhanced the mechanical stability of scaffolds for further acellular and cellular evaluations in vitro which is suggested to be the result of collagen filling cracks on the surface of the struts and bridging gaps. However, to reach high values of mechanical properties as reported by Wenger et al.308, collagen fibers require a parallel orientation which is not the case in the present applied collagen coating (Figure 38). It was also observed that rehydration transforms collagen into a hydrogel which results in an additional drop of the compressive strength (0.069 ± 0.004 MPa). Based on the natural character of collagen it is difficult to reach higher values of compression strength as it is the case when synthetic polymers are used as coating material of similar scaffolds

(e.g. poly(D, L-lactide) acid or poly(3-hydroxybutyrate-co-3-hydroxyhexanoate)309 or polycaprolactone232). Nevertheless, due to the homogeneous collagen distribution, the scaffolds retained their cylindrical shape after compression (Figure 49). Non-coated 45S5 BG-based scaffolds, which are fundamentally very brittle, collapsed instantly and were completely destroyed after compression, as shown in Figure 49.

Collagen – As coating material

45S5 BG-based scaffolds 45S5 BG-based scaffolds without coating with collagen coating (dry state)

Figure 49: Digital camera image of 45S5 bioactive glass-based scaffolds with and without collagen coating after compression.

Compressive strength σ of scaffolds was measured also after the immersion in SBF for 7 and 14 days. In case of collagen-coated 45S5 BG-based scaffolds, the embedding of HA into the collagen, as illustrated in Figure 44 and 45, is expected to enhance the mechanical properties. The proof for this hypothesis is provided in Figure 50 which shows the calculated mean values of compressive strength.

0.25 0.30

0.25

0.20 [MPa]

[MPa] 0.20  0.15 

0.15 45S5 BG-based scaffolds, uncoated - after 7 d in SBF 45S5 BG-based scaffolds, collagen-coated (uc) - after 7 d in SBF 45S5 BG-based scaffolds, collagen-coated (cl) - after 7 d in SBF 0.10 0.10 45S5 BG-based scaffolds, uncoated - after 14 d in SBF 45S5 BG-based scaffolds, collagen-coated (uc) - after 14 d in SBF 45S5 BG-based scaffolds, uncoated 45S5 BG-based scaffolds, collagen-coated (cl) - after 14 d in SBF 0.05 45S5 BG-based scaffolds, collagen-coated (uc)

0.05 Compressive strength Compressive 45S5 BG-based scaffolds, collagen-coated (cl) strength Compressive

Figure 50: Mean values of compressive strength of 45S5 BG-based scaffolds with and without collagen coating after immersion in SBF for 7 and 14 days (dry state).

The compressive strength increased to 0.27 ± 0.03 MPa and 0.26 ± 0.03 MPa for uncrosslinked and crosslinked collagen, respectively, after 14 days of immersion in SBF. In contrast to this, the compressive strength of as-fabricated 45S5 bioactive glass-based scaffolds deteriorated rapidly with immersion time in SBF which is a result of the progressive dissolution of the 45S5 BG126,310,311. The dissolution of 45S5 BG also takes place in collagen-coated scaffolds but it is decelerated because of the presence of the coating and compensated due to the formation of HA crystals in the collagen matrix.

Zein – As scaffold material

4 Zein Composite scaffolds with bioactive glass and zein

4.1 Reinforcement of porous zein scaffolds with bioactive glass particles – Manufacturing techniques, results and discussion

Approach A highlighted the manufacturing process of composite scaffolds based on collagen and bioactive glass. Approach B focuses on zein, a rarely used but promising natural protein for the application in bone tissue engineering (2.2.3 Zein – a protein derived from corn). The first part of this chapter describes the creation of porous zein scaffolds (section 4.1) by the so-called salt leaching technique. Pure zein scaffolds were additionally reinforced with bioactive glass particles to enhance their bioactive behavior. Received scaffolds were characterized in terms of microstructure, degradation behavior and bioactivity to assess zein as biomaterial. In the second part (section 4.2), zein was applied as coating material on bioactive glass-based scaffolds (produced by the foam replica method) to enhance their mechanical stability.

4.1.1 Scaffold preparation

Experiments were performed under equal conditions to keep the possibility of reproducibility. Only chemicals of analytical grade were used which are listed detailed in appendix I.

Porous zein scaffolds without bioactive glass particles

For the fabrication of porous polymeric scaffolds many different techniques are available and described in the literature, e.g. scaffold preparation by induced phase separation of polymer solutions106,107,312,313, electrospinning96,314–316 or rapid prototyping88,110,317. In case of zein, particle leaching156,318–321 is the most frequently applied technique to produce porous scaffolds. Within the framework of this thesis, porous zein scaffolds were also produced by the leaching technique whereas salt was used as porogen. Therefore, zein was homogeneously mixed with sodium chloride particles (< 125 µm) at a ratio of 1:2 (zein:NaCl). The mixture was then compressed applying a load of 104 N in an electrical hydraulic pressing device, resulting in cylindrical specimens with dimensions of Ø = 10 mm and h = 6.5 mm. To receive a porous structure, the samples were leached at 80 °C in a water bath (UPW) for 2 hours, washed in ultrapure water and lyophilized for further use.

Reinforced porous zein scaffolds with bioactive glass particles

For the reinforcement of porous zein scaffolds with bioactive glass particles, a concentration of 70 % zein - 30 % bioactive glass (in wt.%) was chosen based on similar studies reported in the literature322–324. According

Zein – As scaffold material to the fabrication process described above, a dry and homogeneous mixture of zein and 45S5 BG was prepared in the correct ratio by using a mortar. The zein-BG powder was then mixed with sodium chloride particles at a ratio of 1:2 (zein-BG:NaCl) and pressed into cylindrical shape (Ø = 10 mm, h = 6.5 mm). For leaching, samples were immersed in a water bath (UPW) at 80 °C for 2 hours. The received scaffolds were gently rinsed in ultrapure water and lyophilized for further use.

4.1.2 Morphological and microstructural characterization

Porous zein-based scaffolds with and without bioactive glass reinforcement, which were produced previously by particulate leaching, were characterized in terms of morphology and microstructure. The results are presented next.

Pure zein

For the characterization of the microstructure, the obtained scaffolds were analyzed by scanning electron microscopy (SEM). In Figure 51, two different views are presented.

Top Top Top

Cross-section Cross-section Cross-section

Figure 51: SEM images of porous zein-based scaffolds (top and cross-section) at different magnifications showing the porous structure after the leaching process.

The top of the scaffolds shows many concavities indicating a highly porous structure with macropores between 100 and 400 µm. Struts diameter is in the range of 75 to 400 µm. At higher magnification, also micropores < 1 µm on the surface are visible. The cross-section of the samples shows a porous structure with many

Zein – As scaffold material

macro- and micropores which form a cave-like structure with interconnected porosity. The porous structure, which was already observed on the top of the scaffold, continues in the inside of the scaffold and confirms the successful leaching of the sodium chloride particles not only at the outside but also inside of the zein scaffold. The dimensions of the obtained zein scaffolds (n=4) were determined (Table 10) to evaluate their porosity. Based on equations [3], [4] and [5], a porosity of ~ 75 % was calculated for pure zein scaffolds produced by

3 321 the salt leaching technique (ρmaterial of zein = 1.22 g/cm ) .

Table 10: Calculated values of density ρsample and porosity P of zein-based scaffolds after the leaching process.

3 No. Mass [g] Diameter [cm] Height [cm] ρsample [g/cm ] Porosity [%] 1 0.197 1.12 0.77 0.26 78.71 2 0.199 1.05 0.81 0.28 76.74 3 0.2 1.08 0.66 0.33 72.89 4 0.196 1.07 0.67 0.33 73.33

After the salt leaching process, a structural change of zein was observed which was documented and investigated by FTIR spectroscopy. Pellets were prepared by mixing 1 wt.% of each sample with KBr, as already described before. The spectra were recorded in absorbance mode at IR wavenumbers between 4000 cm-1 and 400 cm-1. Most characteristic peaks are labeled in Figure 52.

-1

Pure zein (as-received) Amide I Amide

1662 cm 1662 Zein after leaching

-1

1535 cm 1535 II Amide

-1

1240 cm 1240 III Amide

3500 - 2800 cm-1

Amide A Absorbance [a.u] Absorbance

4000 3500 3000 2500 2000 1500 1000 500 Wavenumber [cm-1]

Figure 52: FTIR spectra of pure zein before and after the leaching process showing the reduced intensity of the peaks after leaching indicating a structural change of the zein. The identified peaks are discussed in the text.

Zein – As scaffold material

Pure zein (as-received), which was available as dry powder, shows a typical amide A band in the region between 3500 cm-1 and 2800 cm-1. Amide A belongs to the stretching of N-H and O-H bonds of the amino acids145,325,326. Peaks representing amide I, II and III are visible at 1662 cm-1, 1535 cm-1 and 1240 cm-1, respectively. Amide I can be attributed to the C=O stretching145,326,327 of the peptide whereas amide II is associated to the C-N stretching and N-H deformation145,326,327. Amide III is related to N-H and C-N deformation145,326,327. During the leaching of sodium chloride particles, water is able to penetrate the scaffolds’ structure. As described

328 by Gillgren et al. , the glass transition temperature (Tg) of zein is reduced with increasing water content. Zein,

35 which normally shows a Tg at 165 °C in dry state , exhibits a Tg below 60 °C with 10 wt.% of water. Amide- amide hydrogen bonds in the structure of zein (for details see section 2.2.3 Zein – a protein derived from corn) are very important for the secondary and tertiary structure329. It is assumed that the penetrating water interacts with the hydrogen bonds in the amide-amide linkage330,331 which can explain the reduced intensity of the peaks in the FTIR spectrum after leaching (Figure 52). Due to the interaction of water with the hydrogen bonds

(Figure 53), the zein structure is weakened and consequently Tg is reduced. During the leaching process, which takes place at 80 °C, zein becomes viscous and is plasticized332. The effect can also be observed in the SEM images (Figure 51). Scaffolds show a smooth and continuous surface caused by the plasticization. However, a full scientific explanation of the process during leaching needs more experimental data (formation of primary, secondary and tertiary structure) and also more information about the structure of zein as the mechanism of self-assembly is not fully understood yet.

A) B) Hydrogen bond

Figure 53: Schematic illustration of possible (A) amide-amide linkage and (B) amide-water linkage329,330 (created by ChemBioDraw Ultra 13.0).

Zein with bioactive glass reinforcement

Figure 54 shows SEM images of zein scaffolds reinforced with 30 wt.% of 45S5 bioactive glass particles after the leaching process. Small agglomerations (< 10 µm) of bioactive glass particles are homogeneously distributed inside the scaffold. The scaffold’s structure itself exhibits many macro- and micropores with interconnected porosity, as already described above. In addition, in can be observed that there is no visible connection

Zein – As scaffold material

between the organic and the inorganic phase. The BG particles are not embedded in the zein, as it is the case with 45S5 BG and collagen (section 3.1.2 Collagen with bioactive glass particles, Figure 14). It is assumed that this behavior is associated with the hydrophilicity of the BG and the hydrophobicity of the zein matrix.

Figure 54: SEM images of porous zein scaffolds + 30 wt.% of 45S5 BG particles after the leaching process at different magnifications showing small agglomerations (< 10 µm) of bioactive glass particles are homogeneously distributed inside the scaffold. However, due to the hydrophilicity of the BG and the hydrophobicity of the zein matrix, the BG particles are not embedded inside the zein.

In the FTIR spectra (Figure 55), collected between 1800 cm-1 and 650 cm-1, the peaks corresponding to 45S5 bioactive glass can be clearly identified.

-1 Zein + 45S5 BG after leaching

Zein after leaching

1040 cm 1040 Si-O-Sistretching

-1

940 cm 940 Si-O-2NBOstretching

-1

1662 cm 1662 I Amide

-1

1535 cm 1535 II Amide

1240 cm 1240 III Amide Absorbance [a.u] Absorbance

1800 1600 1400 1200 1000 800 Wavenumber [cm-1]

Figure 55: FTIR spectra of zein after leaching with and without BG reinforcement (30 wt.%) showing additional peaks (1040 cm-1 and 940 cm-1) after the addition of bioactive glass. The identified peaks are discussed in the text.

The great peak at 1040 cm-1 can be attributed to the Si-O-Si stretching vibrations233,289 in the glass network whereas the smaller peak at 940 cm-1 is assigned to the vibrational mode of the Si-O-2NBO bond175,233. Amide I, II and III can be seen at 1662 cm-1, 1535 cm-1 and 1240 cm-1, respectively, and belong to the zein

Zein – As scaffold material structure, as already explained. However, it was found that the intensity of the peaks, which was reduced after leaching due to the interaction of water with amide-amide linkages, (explanation in section above) is increased again after the addition of BG. It is assumed that the interaction of water with amide-amide linkages is weakened due to the interaction of water with BG. As a consequence, also the plasticization process is affected. This hypothesis can be supported by observations during handling of the different samples. Pure zein scaffolds showed a rubber-like structure after the leaching process whereas zein-BG composites exhibited a more brittle and fragile behavior (see section 4.1.5 Mechanical characterization). It is expected that during the leaching process not only sodium chloride but also bioactive glass particles are dissolved. Therefore, for the determination of the porosity of the zein-BG scaffolds, the correct ratio of zein:BG has to be ascertained. For this determination, TGA was carried out in ambient atmosphere and with a heating rate of 5 K/min. Results showed a ratio of 74:26 (in wt.%) for zein and bioactive glass after leaching, respectively. Referring to equation [6], the density ρcomposite of the zein-BG composite was calculated as

mcomposite mzein m ρcomposite [Eq. 9] composite zein

3 321 where mzein and Vzein represent the mass and volume of zein (ρzein = 1.22 g/cm ) and mBG and VBG represent

3 29 the mass and volume of 45S5 bioactive glass (ρBG = 2.7 g/cm ) , respectively.

3 Based on the results of the TGA measurement, ρcomposite was calculated as 1.42 g/cm which indicates that zein-BG composite scaffolds exhibit a porosity of ~ 84 %. For comparison, pure zein scaffold showed a porosity of ~ 75 %. The difference is likely caused by the additional leaching of bioactive glass particles.

4.1.3 Swelling properties and degradation behavior

Zein exhibits an amphiphilic character and consists of many hydrophilic-hydrophobic groups in the zein chains145 (section 2.2.3). Leucine, proline and alanine are hydrophobic amino acid residues and are the main reason for the poor solubility of zein in water, next to the presence of hydrocarbon groups in the side chains35,145. However, the swelling properties in aqueous solutions are important in order to explain and assess the degradation behavior of zein in a biological environment.

Swelling properties

For the evaluation of the swelling properties, simple sorption experiments were carried out. Therefore, samples (n=5) of pure zein scaffolds and zein composites with bioactive glass particles were immersed in PBS at room temperature up to 24 h. After different time points, the scaffolds were removed and excess solution was dried with a filter paper (qualitative filter paper 410). The water uptake was determined considering

Zein – As scaffold material

equation [7] and the results are presented in Figure 56. Results show that the equilibrium for pure zein was reached after 24 hours. Due to the hydrophobic character and the plasticization of zein, which occurs during leaching, the relative water uptake is limited to 400 %. In case of the bioactive glass reinforcement, equilibrium was already reached after 3 h of immersion. It is obvious that during the initial state the swelling is faster for zein-BG composites. This behavior can be attributed to the bioactive glass particles which additionally absorb water in the polymer matrix. In addition, the composite scaffolds exhibit a higher porosity. These results suggest that zein-BG samples show a higher susceptibility to degradation compared to pure zein which is discussed next.

24 h 3 h 1 h 500

400

300

200

Relative[%] water uptake 100

0 Zein Zein + BG

Figure 56: Results of the swelling study on zein-based scaffolds: relative water uptake in PBS up to 24 h for zein and zein-BG scaffolds.

Degradation behavior

To retrieve information about the structural stability of zein and zein-BG composite scaffolds in vivo, samples (n=3) were investigated in vitro in the presence of enzymes. The literature describes enzymatic degradation studies using pepsin or collagenase156. Pepsin, a digestive enzyme, is often used in contact with zein due to the application of zein in food industry. However, the present zein scaffolds are designed for a possible application in bone tissue engineering. Collagenase, an extracellular peptidase, is part of the tissue remodeling process in the whole body and therefore more relevant in the context of this study compared to pepsin. According to that, the degradation behavior of zein and zein-BG composites was investigated by bacterial collagenase from Clostridium histolyticum. Referring to section 3.1.3 Degradation behavior, a buffer solution of 0.1 M Tris-HCl,

0.005 M CaCl2 and 0.05 mg/ml NaN3 was prepared. Samples were soaked in 1 ml of the Tris-HCl buffer solution at 37 °C. After 24 h of immersion, 1 ml of Tris-HCl solution with 200 CDU/ml of collagenase was

Zein – As scaffold material added to reach the desired concentration of 100 CDU/ml of collagenase. The enzymatic degradation was stopped after different time points (1 h, 24 h, 7 d and 14 d) by addition of 0.25 M EDTA buffer. Scaffolds were washed in UPW, frozen and lyophilized. The degradation rates were determined according to equation [8] and are presented in Table 11. Reference samples were prepared without collagenase. After 14 days of incubation with collagenase, pure zein revealed a high stability with a degradation rate of 35.2 ± 0.8 % which has been also reported in the literature before156. In contrast, zein-BG composites were completely destroyed after 14 days in contact with enzymes. However, it was observed that zein-BG composites were not completely digested but the three-dimensional structure of the scaffolds was damaged. The higher degradation rate of zein-BG composites can be explained by the higher swelling behavior and the lower degree of plasticization of composite scaffolds compared to pure zein.

Table 11: Degradation rates of pure zein and zein-BG composites after the immersion in collagenase solution (100 CDU/ml).

1 h 24 h 7 d 14 d Reference (14 d) Pure zein 6 ± 1 % 7 ± 1 % 31 ± 1 % 35 ± 1 % 0 % Zein + 45S5 BG 10 ± 1 % 41 ± 2 % 100 100 13 ± 1 %

On the other hand, results revealed a high degradation resistance of zein in absence of enzymes. Pure zein did not show any weight loss after 14 days of immersion in PBS which suggests a high stability of zein in water-based solutions due to its insolubility. Zein-BG composites exhibit a degradation of 13 ± 1 % after 14 days of immersion without collagenase. It is assumed that the weight loss is not connected to a degradation of the zein matrix but to the dissolution of the bioactive glass particles.

4.1.4 Evaluation of bioactivity

For the application of zein or zein-BG scaffolds in bone tissue engineering, the evaluation of the bioactive behavior is essential. The bioactivity can be assessed in vitro, as already described in section 3.1.4 Evaluation of bioactivity. Simulated body fluid was prepared following the protocol described there. Samples (n=3) of pure zein and zein-BG composites were immersed in 50 ml of SBF and incubated at 37 °C in an orbital shaker. After different time points, samples were removed, washed in UPW and dried in a drying chamber at 60 °C. The formation of hydroxyapatite was investigated by SEM observations (Figure 57 and 59) and analyzed by FTIR (Figure 58 and 60).

Zein – As scaffold material

Pure zein - 7 d Pure zein - 14 d Pure zein - 14 d

Figure 57: SEM images of pure zein after 7 and 14 days of immersion in SBF at different magnifications showing no formation of hydroxyapatite even after 14 days of immersion.

Figure 57 shows SEM images of pure zein scaffolds after the conditioning in SBF up to 14 days. However, even after 14 days, no hydroxyapatite can be found on the surface of zein and therefore no bioactive behavior can be detected. This is confirmed by FTIR results, shown in Figure 58, which do not indicate specific changes in the spectra that could confirm the mineralization of the pure zein matrix. Only amide I, II and III can be

identified which are part of the zein structure, as already discussed above.

-1

1635 cm 1635 I Amide

1516 cm 1516 II Amide

-1

1238 cm 1238 III Amide

14 d

7 d

3 d

Absorbance [a.u.] Absorbance

1 d

0 d

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 58: FTIR spectra of pure zein after 0, 1, 3, 7 and 14 days of immersion in SBF showing no specific changes in the spectra indicating the non-bioactivity of pure zein. The identified peaks are discussed in the text.

After the reinforcement with 45S5 BG particles, the bioactive behavior clearly changed which can be seen in the SEM images presented in Figure 59. After 3 days, round-shaped crystals of hydroxyapatite are seen to be randomly distributed in the inner structure of the scaffold. With increasing immersion time in SBF, the hydroxyapatite layer becomes denser until the surface is homogeneously covered with the typical cauliflower structure of HA21,250 after 14 days.

Zein – As scaffold material

The formation of hydroxyapatite can be confirmed by FTIR. KBr pellets were prepared and analyzed in absorbance mode between 1800 cm-1 and 600 cm-1. Figure 60 shows the FTIR spectra of zein scaffolds reinforced with 30 wt.% of 45S5 BG after different time points of immersion in SBF. Amide peaks, visible at 1635 cm-1 and 1516 cm-1, belong to the zein structure145,326,327. The apatite formation, which was already visible in the SEM micrographs, is proven by the growing peaks in the area between 1460-1400 cm-1 and 1100-1000 cm-1 and can be assigned to C-O stretching253,254 and P-O stretching253, respectively. The broad

-1 -1 3- peak at 1235 cm and the small peak at 960 cm are attributed to the P-O stretching of the PO4 in the calcium phosphate layer253,333–335. Finally, the peak at 875 cm-1 indicates the formation of carbonated hydroxyapatite on the surface of zein-BG composites which can be attributed to the C-O bending253,254. The peak at 798 cm-1 is characteristic for the formation of a silica rich layer291.

Zein + 45S5 BG - 3 d Zein + 45S5 BG - 7 d Zein + 45S5 BG - 14 d

Figure 59: SEM images of zein composite scaffolds with 30 wt.% of 45S5 BG after the immersion in SBF for 3, 7 and 14 days (at different magnifications) showing the formation of hydroxyapatite on the surface.

Zein – As scaffold material

1100-1000 cm-1

P-O stretching

-1

1635 cm 1635 I Amide -1

1516 cm 1516 II Amide

1235 cm 1235 stretching P-O

-1

1460-1400 cm 1460-1400 stretching C-O

960 cm 960 stretching P-O

Si-O-Si

798 cm 798 

875 cm 875 bending C-O

14 d

7 d

3 d

Absorbance [a.u.] Absorbance

1 d

0 d

1800 1600 1400 1200 1000 800 600 Wavenumber [cm-1]

Figure 60: FTIR spectra of zein composite scaffolds with 30 wt.% of 45S5 bioactive glass after 0, 1, 3, 7 and 14 days of immersion in SBF showing a growing peak in the area between 1100 and 1000 cm-1 indicating the formation of hydroxyapatite. Further identified peaks are discussed in the text.

4.1.5 Mechanical characterization

In order to evaluate the mechanical properties of the obtained scaffolds, compression strength tests were performed. Uniaxial compression test were carried out with a preload of 0.1 N, a crosshead velocity of 1 mm/min and a maximum applied force of 1 kN (Zwick Z050, DE). Ten specimens of each sample group were subjected to compression until fracture. The strength was measured by the maximum stress immediately before the fracture point. Results showed higher compressive strength for pure zein (4.1 ± 0.8 MPa) compared to zein-BG composite scaffolds (2.2 ± 0.9 MPa). This behavior was already observed during handling of the scaffolds in which the composite scaffolds exhibited a more fragile and powdery structure than pure zein samples. It is assumed that the addition of bioactive glass hampers the plasticization process during leaching, as explained above. In addition, composite scaffolds show a lower compressive strength due to their higher porosity. However, compressive strength of natural cancellous bone ranges between 2 and 12 MPa51,306, therefore the measured values of pure zein and zein-BG composite scaffolds (combined with the enhanced bioactive behavior) match the requirements. Therefore, zein is of great interest for the potential application in bone tissue engineering.

Zein – As coating material

4.2 Zein as coating material on bioactive glass-based scaffolds – Manufacturing techniques, results and discussion

This section discusses the application of zein as coating material for 45S5 BG-based scaffolds which were produced with a highly interconnected porous structure by the foam replica technique described previously (section 3.2.1). As mentioned above, the low structural integrity of these brittle scaffolds is an important challenge for their application in bone tissue engineering. To overcome this problem, different kinds of coatings can be applied336. In this thesis, zein was investigated for the first time in terms of its potential as coating material for such kind of scaffolds. Zein-coated bioactive glass-based scaffold were thus produced via dip coating and characterized concerning microstructure, bioactivity and mechanical strength.

4.2.1 Sample preparation

Experiments were carried out under equal conditions to ensure the reproducibility. Only chemicals of analytical grade were used and are listed in appendix I.

Scaffold fabrication

For the fabrication of three-dimensional bioactive glass-based scaffolds the foam replica technique was applied, following the process described originally by Chen et al.29 The protocol in detail can be found in section 3.2.1 Scaffold and pellet fabrication. Shortly summarized, a slurry was prepared by dissolving PVA (0.3 wt.%) at 80 °C in ultrapure water. After cooling down to room temperature, KV 9062 (2 wt.%) and 45S5 BG powder (50 wt.%) were added and homogenously mixed. PU foams with 45 ppi served as sacrificial template (Figure 61, light microscope images taken with M50, Leica, DE) for the scaffold fabrication and were cut into cylindrical shape with dimensions of 7x12 mm (height x diameter).

Figure 61: Light microscope images of sacrificial polyurethane foams (45 ppi): top view (left) and cross-section (right).

The polymeric foam structure was then coated with bioactive glass slurry by dip coating. Therefore, samples

Zein – As coating material

were immersed into the BG slurry, retrieved after 5 minutes and squeezed to remove excess slurry. After drying at 60 °C for minimum 2 hours, the coating procedure was repeated to increase the thickness of the BG layer. Afterwards, the double-coated green bodies were dried again for at least 24 h at 60 °C. The sintering profile was chosen as represented in Figure 22. The sintering process revealed 45S5 bioactive glass-based scaffolds with a highly interconnected porous structure, shown in Figure 62.

Figure 62: Light microscope images of 45S5 BG-based scaffolds after the sintering process: top view (left) and cross-section (right).

Results of the characterization of the as-fabricated 45S5 BG-based scaffolds were presented in section 3.2.1.

Zein coating

A promising and simple technique for the coating of three-dimensional structures is the so-called dip coating. First, a solution of the coating material is prepared, then samples to be coated are immersed into the solution and remain there for a defined period of time. After the removal from the solution, a thin layer is formed on the surface of the sample. When the solvent evaporates, a thin polymer coating is left. Due to its simplicity, this method was selected for coating porous 45S5 BG-based scaffolds with zein. However, one of the main characteristics of zein is its insolubility in water associated with the high amount of non-polar amino acid residues (section 2.2.3 Structure and morphology). Several methods to solubilize zein are described in the literature337–339. For commercial use, aqueous alcohol solutions are commonly used35. Zein is soluble in aqueous ethanol solutions of 50-90 vol.%. Preliminary studies were carried out to find the right concentration of ethanol for the coating procedure. It was found that an aqueous zein solution with 80 vol.% of ethanol showed the best parameters in terms of zein solubility, viscosity and deposition behavior. Thus, different concentrations of zein (2 to 25 wt.%) were prepared in 80 vol.% of ethanol. Solutions with a low amount of zein (< 6 wt.%) led to poor quality coatings on the surface of BG-based scaffolds. Solutions with concentrations above 16 wt.% resulted in clogged pores. The most promising concentration was found to be 8 wt.%, which was used for further investigations, considering also results in the literature300. In summary, zein-coated 45S5 BG-based scaffolds were fabricated considering the following protocol. A zein solution was

Zein – As coating material

prepared by dissolving 8 wt.% of zein powder in 80 vol.% of ethanol. A magnetic stirrer was used to homogenously mix zein powder with the solvent. Scaffolds were immersed into the zein solution, retrieved after 5 minutes and left to dry at room temperature. SEM images in Figure 63 show the surface of 45S5 BG- based scaffolds after the coating with zein solution (8 wt.% of zein). For comparison, BG-based scaffolds without coating are shown in Figure 23. It is visible that zein exhibits a smooth surface applied as coating material. However, zein only partially covers the surface of the scaffolds with a thickness of  10 µm. The overall connected porosity is not significantly hampered. Nevertheless, the achieved inhomogeneous distribution of zein is desired in terms of keeping the high bioactivity of 45S5 BG-based scaffolds and is discussed further below (section 4.2.3 Evaluation of bioactivity).

Coating thickness

~ 10 µm Zein

Figure 63: SEM images of zein-coated BG-based scaffolds (8 wt.% of zein) at different magnifications showing the partially covered surface of the scaffold with zein.

Thermogravimetric analysis was carried out to give information about the amount of zein deposited on the scaffolds. For the TGA measurement, heating was performed in ambient atmosphere and with a heating rate of 5 K/min. Figure 64 shows the mass loss of uncoated and zein-coated BG-based scaffolds during TGA measurement. 45S5 bioactive glass exhibits a low mass loss ( 2 %) which can be ascribed to the evaporation of water during the heat treatment. Zein-coated scaffolds show a mass loss divided into three different stages which has been reported previously in the literature300. Stage I, in the range between room temperature and 250 °C, can be attributed to the loss of water. Stage II (250 - 450 °C) describes the thermal degradation of zein340 whereas the weight loss in stage III (450 - 600 °C) corresponds to the pyrolysis of residual organic components. Therefore, the amount of zein can be determined as  48 wt.%.

Zein – As coating material

I. -10

II. -20

-30 III.

Mass loss [%] loss Mass -40

-50

45S5 BG-based scaffolds without coating 45S5 BG-based scaffolds coated with zein -60 0 100 200 300 400 500 600 Temperature [°C]

Figure 64: Mass loss of uncoated and zein-coated 45S5 BG-based scaffolds during TGA ambient atmosphere (heating rate 5 K/min).

Crosslinking process

It is assumed that the swelling behavior of the zein coating on the surface of 45S5 BG-based scaffolds can be positively influenced by introducing a crosslinking step (results are discussed in section 4.2.2). For the crosslinking of zein, the literature mentions different reactants, e.g. citric acid341 or tetrahydrofuran342 (THF). Also NHS and EDC343 are described which were already used for the crosslinking of collagen within the framework of this thesis. Therefore, zein was crosslinked by EDC and NHS, referring to the protocol described in section 3.2.1 Crosslinking process. Shortly summarized, a 50 mM MES acid solution was prepared in 40 vol.% of ethanol. EDC and NHS were dissolved in the MES acid solution at a ratio of 1:1 (60mM) to obtain the crosslinking solution. Zein-coated scaffolds were immersed for 4 h at RT and afterwards washed in

0.1 M Na2HPO4 for 2 h to hydrolyze remaining O-acylisourea of the carbodiimide. Subsequently, crosslinked scaffolds (cl) were washed in UPW and left to dry at 60 °C. Uncrosslinked samples are labeled as uc. Figure 65 shows the FTIR spectra of zein-coated 45S5 BG-based scaffolds before and after the crosslinking process. For comparison, results of uncoated scaffolds are also included. Pure 45S5 BG shows typical peaks in the region between 1100 cm-1 and 1000 cm-1 and at 926 cm-1, 621 cm-1, 532 cm-1 and 455 cm-1 which were explained before. The zein coating reveals additional peaks at 1660 cm-1 and 1525 cm-1 which can be attributed to the amide I and amide II, respectively. After the crosslinking of zein, no change in the spectrum can be observed. However, this does not indicate the failure of the crosslinking process. Crosslinking with EDC and NHS forms amide-type crosslinks. It is assumed that the success of the crosslinking is not visible in the FTIR spectrum due to the overlapping with the existing amide peaks. To provide evidence for the crosslinked

Zein – As coating material

structure of zein, the degree of crosslinking was determined by the ninhydrin assay, as applied in section 3.1.1. Zein-coated scaffolds (50 ± 0.4 mg) with and without crosslinks were immersed in 5 ml of ninhydrin solution (2 wt.% dissolved in 99.8 vol.% of ethanol) and heated in a water bath at 100 °C for 20 minutes. The concentration of free amino groups is directly proportional to the intensity of the violet color of Ruhemann’s purple and was measured at 570 nm via UV-Vis spectroscopy (Specord 40). The degree of crosslinking was calculated according to equation [2]. Results revealed a zein structure with 10 ± 1 % of crosslinks. The relatively low degree of crosslinking may be attributed to the low amount of free amino groups in the zein structure. Also, the swelling behavior can affect the crosslinking process which is discussed next.

45S5 BG-based scaffold + Zein (cl) 45S5 BG-based scaffold + Zein (uc)

45S5 BG-based scaffold

-1

621 cm 621 Si-O-Sibending

1525 cm 1525 II Amide

-1

1660 cm 1660 I Amide

-1

532 cm 532 Si-O-Sibending

926 cm 926 Si-O-2NBOstretching

-1

1000 - 1100 cm 1100 - 1000 stretching Si-O-SiP-O stretching,

455 cm 455 Si-O-Sibending

Transmission[a.u.]

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 65: FTIR spectra of as-fabricated 45S5 BG-based scaffold, zein-coated 45S5 BG-based scaffolds with and without crosslinking showing the formation of additional peaks (amide I and II) after the zein coating confirming the presence of the protein. The identified peaks are discussed in the text.

4.2.2 Swelling properties and degradation behavior

To clearly understand the processes during crosslinking and to explain the degradation behavior of zein-coated BG-based scaffolds, the swelling kinetics is of great interest. Although zein is insoluble in water, water molecules can penetrate the zein structure which leads to swelling. In the worst case, the porous structure of the scaffold can be destroyed. In addition, the transportation of solvent molecules inside the structure of the scaffold influences cell behavior, growth and differentiation.

Zein – As coating material

Swelling properties

Swelling behavior of zein-coated scaffolds was investigated by simple sorption experiments, as described before (section 3.2.2). Samples with and without crosslinks were immersed in PBS solution up to 24 h at room temperature. Samples were removed and excess solution was dried with a filter paper (qualitative filter paper 410). Water uptake was calculated according to equation [7]. Results revealed that equilibrium was reached after 24 h of immersion and are presented in Figure 66. It is assumed that the swelling during the first hour is mainly attributed to the water absorption of the bioactive glass particles. After 24 hours, uncrosslinked zein shows a water uptake of ~ 115 % whereas crosslinked samples exhibit a water uptake of ~ 97 %. Introduced crosslinks can occupy or reduce binding sites for the water molecules which can describe the reduced swelling behavior. In addition, the hydrophobic character of zein influences the swelling properties. This can also explain the low degree of crosslinking. Due to the limited swelling behavior, EDC and NHS have difficulties penetrating the zein structure and therefore the crosslinking reaction is hindered.

Figure 66: Results of the swelling study on zein-coated scaffolds: relative water uptake of zein-coated BG-based scaffolds with and without crosslinking in PBS after 24 h.

Degradation behavior

An important parameter for tissue engineering applications of biomaterials is their degradability. To retrieve information about the stability of the zein-coated 45S5 BG-based scaffolds (n=4), they were immersed in PBS solution for different time points. The degradation behavior was determined by the weight loss of the samples and calculated according to equation [8]. Figure 67 shows the different curves representing the degradation rate up to 7 days of immersion in PBS. Zein-coated 45S5 BG-based scaffolds show a weight loss around 36 % which is independent from the crosslinking process. However, the weight loss is not attributed to the

Zein – As coating material

degradation of zein but mainly to the constant dissolution and leaching of the bioactive glass structure119. This is obvious as the uncoated BG-based scaffolds exhibit almost the same degradation rate (~ 34 %) as zein-coated samples. The slight weight loss of zein can be explained by its swelling and diffusion. To prove that the weight loss of the scaffolds is mainly caused by the reaction with water, reference samples of uncoated BG-based scaffolds were immersed in 96 vol.% of ethanol. During immersion in ethanol, the dissolution of bioactive glass is slowed down due to the low percentage of water. Therefore, it can be stated that zein-coated scaffolds show a high stability in water-based systems, as already reported before (section 4.1.3).

Zein-coated BG-based scaffold, uc - in PBS Zein-coated BG-based scaffold, cl - in PBS uncoated BG-based scaffold - in PBS

40 uncoated BG-based scaffold in 96 vol.% of ethanol

Degradation[%] rate 10

0 50 100 150 200 Time [h]

Figure 67: Degradation rate of zein-coated 45S5 BG-based scaffolds with and without crosslinking in PBS.

4.2.3 Evaluation of bioactivity

For the assessment of the bioactive behavior of zein-coated bioactive glass-based scaffolds, which is an important parameter for the bone-bonding ability, samples were immersed in SBF up to 10 days according to section 3.2.3. After different periods of time (1, 3, 7 and 10 d), samples were removed. The influence of the zein coating on hydroxyapatite formation on the surface of the samples was investigated by SEM observation and FTIR measurement. Figure 68 and 69 show the results for zein-coated 45S5 BG-based scaffolds without crosslinks. In the SEM images (Figure 68) the small round shape of calcium phosphate with a size of ~ 1 µm is visible after only 3 days. With longer immersion times, the layer of hydroxyapatite crystals on the surface becomes denser. After 10 days, the surface is homogenously covered and hydroxyapatite can be clearly identified by its typical cauliflower-like morphology21,250. The FTIR spectra in Figure 69 confirm the SEM observations. Peaks of amide I

Zein – As coating material

and II at 1660 cm-1 and 1525 cm-1, respectively, verify the presence of zein on the surface145,326,327. The broad peak in the area between 1100 and 1000 cm-1 and the small shoulder at 960 cm-1 can be attributed to the

253,301,302 2- stretching mode of the phosphate phase in the hydroxyapatite layer. C-O bending of the CO3 groups in the carbonated HA253,254,300 can be seen at 875 cm-1 whereas the peak at 795 cm-1 is related to the formation of a silica rich layer291. The crystallinity of the HA layer is shown by the growing double peak at 600 cm-1 and

-1 3- 253,254,300 570 cm and is assigned to the bending of the P-O of the PO4 group .

45S5 BG + 8 % Zein (uc) - 3 d 45S5 BG + 8 % Zein (uc) - 10 d

45S5 BG + 8 % Zein (uc) - 7 d 45S5 BG + 8 % Zein (uc) - 10 d

Figure 68: SEM images of zein-coated 45S5 BG-based scaffolds (uc) after the immersion in SBF for different time points showing the formation of hydroxyapatite.

Zein – As coating material

-1

-1 -1

10 d Si-O-Si

795 cm 795 

960 cm 960 stretching P-O C-O bending C-O

875 cm 875 570 / 600 cm-1

-1 -1

-1 P-O bending

1525 cm 1525 II Amide Amide I Amide

7 d cm 1660

1100-1000 cm 1100-1000 stretching P-O

3 d

1 d

0 d

Transmission[a.u.]

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 69: FTIR spectra of zein-coated 45S5 BG-based scaffolds (uc) after 0, 1, 3, 7 and 10 days of immersion in SBF showing additional peaks at 960 cm-1, 875 cm-1 and 795 cm-1, indicating the formation of HA. The identified peaks are discussed in the text.

Zein – As coating material

45S5 BG + 8 % Zein (cl) - 3 d 45S5 BG + 8 % Zein (cl) - 10 d

45S5 BG + 8 % Zein (cl) - 7 d 45S5 + 8 % Zein (cl) - 10 d

Figure 70: SEM images of zein-coated 45S5 BG-based scaffolds (cl) after the immersion in SBF for different time points showing the formation of hydroxyapatite.

Zein – As coating material

-1 -1

10 d -1

875 cm 875 bending C-O

Si-O-Si

-1

795 cm 795 

960 cm 960 stretching P-O -1

-1 570 / 600 cm -1 P-O bending

7 d

1525 cm 1525 II Amide

1660 cm 1660 I Amide

1100-1000 cm 1100-1000 stretching P-O

3 d

1 d

0 d

Transmission[a.u.]

2000 1800 1600 1400 1200 1000 800 600 400 Wavenumber [cm-1]

Figure 71: FTIR spectra of zein-coated 45S5 BG-based scaffolds (cl) after 0, 1, 3, 7 and 10 days of immersion in SBF showing additional peaks at 960 cm-1, 875 cm-1 and 795 cm-1, indicating the formation of HA. The identified peaks are discussed in the text.

Figure 70 and 71 show similar results obtained with the crosslinked zein-coated 45S5 BG-based scaffolds. It can be noted that the bioactive behavior was not influenced neither by the zein coating nor the crosslinking step. Furthermore, the bioactivity of the 45S5 BG-based scaffolds could be preserved due to the inhomogeneity of the zein coating. As explained in section 4.1.4, pure zein did not show any apatite formation after immersion in SBF for 14 days. A homogenous and dense coating of zein on the surface of bioactive glass-based scaffolds is expected to inhibit the continuous development of a HA layer on the surface. For comparison, the bioactive behavior of uncoated bioactive glass-based scaffolds was reported in section 3.2.3.

4.2.4 Mechanical characterization

To evaluate the potential for the application in bone tissue engineering, the compressive strength σ of zein- coated scaffolds was determined by uniaxial compression strength tests. Cylindrical shaped scaffolds (n=5) were compressed with a crosshead velocity of 1 mm/min (Zwick Z050, DE). The samples were preloaded at 0.1 N and a maximum force of 50 N was applied. The load was applied until a displacement of 4 mm was reached. Figure 72 shows exemplary stress-displacement curves of uncoated BG-based scaffolds as well as zein-coated scaffolds with and without crosslinking. In addition, the mean values of compressive strength σ can be seen.

Zein – As coating material

45S5 BG-based scaffolds, 8 % zein coating (cl) 0.3 0.3 45S5 BG-based scaffolds, 8 % zein coating (uc) 45S5 BG-based scaffolds, uncoated

II. 0.2 [MPa] 0.2 III. 

I. 0.1 0.1 45S5 BG-based scaffolds, uncoated 45S5 BG-based scaffolds, 8 % zein coating (uc)

Compressive stress [MPa] Compressive 45S5 BG-based scaffolds, 8 % zein coating (cl) Compressive strength Compressive

1 2 3

Displacement [mm]

Figure 72: Mechanical properties of 45S5 BG-based scaffolds with and without zein coating. Left: exemplary stress-displacement curve of uncoated and zein-coated BG-based scaffolds; right: mean values of compressive strength σ.

The uncoated 45S5 BG-based scaffolds exhibit a jagged process during compression, as already reported before. This behavior is attributed to the hollow nature of the struts and the microcracks on the surface. Low values of 0.04 ± 0.02 MPa are recorded as compressive strength. The stress-displacement curves of zein-coated scaffolds show the typical three regions for compressed foams which were already explained (section 3.2.4). The compressive strength constantly increases until a maximum is reached and the struts are fractured. As a result, the curves show a negative slope. During the progressive compression, the broken structure is densified and the compressive stress increases again. The crosslinking process does not show any significant effect on the mechanical strength of the samples. Uncrosslinked zein coating reaches values of 0.21± 0.02 MPa whereas crosslinked zein coating shows compressive strength of 0.19 ± 0.03 MPa. Regardless of the applied coating, the mechanical properties of the scaffolds are rather low compared to the compressive strength of cancellous bone in the range of 2 and 12 MPa51,306. However, the area under the stress-displacement curve increased in coated samples indicating higher fracture toughness imparted by the zein coating. Zein coating is still interesting for the application in bone tissue engineering not only due to its good cell compatibility which is described in detail in the next chapter. A comparison between collagen and zein coatings on 45S5 BG-based scaffolds is presented in section 5.3 Comparison of the investigated systems which also involves an analysis of relevant results in the literature.

In vitro studies – a static cultivation system

5 In vitro studies and comparison of the investigated systems

Within the current chapter 5, selected samples were tested in vitro to study their biocompatibility and interactions with cells. Therefore, two different cultivation systems were used. Different scaffolds were cultivated in static conditions with osteoblast-like cells (5.1) and investigated in terms of cell viability and relative proliferation. In addition, a perfusion bioreactor was chosen to mimic a dynamic cultivation system under physiological conditions (5.2). Samples were seeded with a bone marrow stromal cell line and tested regarding their osteoblastic differentiation behavior. Afterwards, the investigated systems were compared, not only amongst each other, but also with results reported in the literature which is presented in section 5.3.

5.1 In vitro studies – a static cultivation system

For cell related experiments, only chemicals suitable for cell culture were used which are listed detailed in appendix I. Statistic evaluation of the results was performed by One-Way ANOVA (Analysis of Variance, Origin 8.6) with significance levels at p ˂ 0.05, **p ˂ 0.01 and ***p ˂ 0.001.

5.1.1 Sample preparation

For the static cultivation, five different kinds of scaffolds were selected: (1) uncoated 45S5 bioactive glass-based scaffolds, (II+III) collagen-coated scaffolds with and without crosslinking as well as (IV+V) zein-coated scaffolds with and without crosslinking. 45S5 BG-based scaffolds without coating served as reference sample and were produced according to the process introduced in section 3.2.1. After the sintering process, scaffolds were sterilized with dry heat at 160 °C for 3.5 h344. Afterwards, the coating procedure (collagen and zein) was performed under sterile conditions. For the collagen coating, required solutions (APTS, acetone, NaOH, HEPES buffer and the crosslinking solution) were prepared as described by the protocol in section 3.2.1 and sterile filtered. 10x DMEM and the 0.5 % collagen solution were received sterile. For the zein coating, a 8 wt.% solution of zein in 80 vol.% of ethanol was prepared and stirred for 1 h which was sufficient for the sterilization process of zein. As already mentioned (section 2.2.1), 45S54 BG shows a high reactivity in aqueous media124 and its dissolution leads to a shift of the pH to an alkaline value. Na+ and Ca2+, which are part of the bioactive glass structure, are

+ + rapidly exchanged by H or H3O from the aqueous solution:

As H+ ions are replaced by cations in the solution, the pH of the solution increases124. This behavior can be

In vitro studies – a static cultivation system observed when scaffolds are immersed in DMEM. Hence, the alkaline pH value does not offer a healthy environment for the cultivation of cells. Therefore, scaffolds were immersed in DMEM without supplements prior to cell seeding. Samples were washed in PBS and the medium was changed several times. The change of pH was observed over time and is depicted in Figure 73.

8.6 45S5 BG-based scaffold Scaffold + collagen coating (uc) Scaffold + collagen coating (cl) Scaffold + zein coating (uc) 8.4 Scaffold + zein coating (cl) Pure DMEM

8.2

8.0 pH

7.8

7.6

0 50 100 150 200 250 300 350 400 Time [h]

Figure 73: pH change in DMEM solution when 45S5 BG-based scaffolds were immersed prior to cell seeding.

After 24 h of immersion, pure 45S5 BG-based scaffolds showed a pH around 8.4. Coated scaffolds (collagen as well as zein coating) exhibit lower pH values due to the reduced contact area of bioactive glass with the medium. With increasing immersion time, the pH was seen to decrease. After 16 days, all samples showed a stable pH value between 7.65 and 7.85 ± 0.09 and were then decided to be suitable for cell culture.

5.1.2 Cell culture and seeding

MG-63, a human osteosarcoma cell line, was used for the static cell experiments. MG-63 are adherent cells received from a malignant bone tumor and are commonly used as osteoblastic model cells345. They were cultured in low glucose DMEM supplemented with 10 vol.% of fetal bovine serum (FBS) and 1 vol.% of penicillin-streptomycin (PenStrep) in an incubator (37 °C with 5 % of CO2 and 95 % humidity). At confluency between 80 and 100 %, cells were collected by trypsinization. After trypsinization, cells were centrifuged, re-suspended in fresh medium and counted by trypan blue exclusion method in a Neubauer chamber. Scaffolds (I-V), which were placed in a 48-well plate, were seeded with 1 ml of cell suspension with 600.000 cells. After 3 days of incubation, scaffolds were turned and medium was changed twice a week. Samples seeded with MG-63 cells were cultured for 7, 14 and 21 days. Afterwards, cell viability and relative proliferation

In vitro studies – a static cultivation system

were investigated by a cell counting kit and cell proliferation ELISA (enzyme-linked immunosorbent assay), respectively. Cell morphology was analyzed by SEM observation.

5.1.3 Cell viability and relative proliferation

The cell viability (n=4) was measured after 7, 14 and 21 days as an indirect measurement of the viable cell number. For the measurement, a cell counting kit was used (Cell Counting Kit - 8, Sigma-Aldrich346). To make sure only adherent cells on the surface of the scaffolds were measured, samples were washed with PBS and moved to a new 48-well plate. A mastermix of 1 vol.% WST-8 in DMEM was prepared. In this case, DMEM without phenol red was used to exclude a change in color due to pH fluctuation. DMEM was supplemented with 10 vol.% FBS, 1 vol.% PenStrep and 200 mM L-Glutamine. Samples were covered with the mastermix of WST-8 and incubated for 4 h. During incubation, WST-8 (a water-soluble tetrazolium salt) is enzymatically reduced by cellular activity of dehydrogenase to an orange-colored formazan (Figure 74) which is dissolved in the culture medium. This compound can be detected colorimetrically by UV-Vis spectroscopy, so absorbance was measured at 450 nm with a microplate reader (PHOmo, Autobio Labtec Instruments).

WST-8 Formazan yellow orange

Figure 74: Schematically structure of WST-8 and formazan (created by ChemBioDraw Ultra 13.0)

Cells, which are metabolic active, metabolize tetrazolium salt to formazan in the cytosol. Thus, the amount of formazan is directly proportional to the number of living cells. Figure 75 shows the cell viability of seeded MG-63 cells on different coated scaffolds which shows increasing values for every sample with incubation time. After 21 days, the uncrosslinked collagen-coated BG-based scaffolds exhibit a significant lower cell viability (105 %) compared to pure bioactive glass-based scaffolds (130 %). It is supposed that cells attach to the collagen matrix but due to the high release rate of uncrosslinked collagen (section 3.2.2) a high cell number is lost in the beginning. On the other hand, the crosslinking process of collagen slows down the release rate, so a higher cell viability can be detected for crosslinked samples (120 %). Zein-coated samples revealed significant lower cell viability in uncrosslinked (89 %) as well as in crosslinked state (113 %). As described in section 4.2.2,

In vitro studies – a static cultivation system zein possesses a water uptake ability of 115 % and 97 % for uncrosslinked and crosslinked samples, respectively. Water uptake leads to swelling of the zein coating which blocks the pores of the scaffold caused by the missing bonding of zein to the BG surface. Blocked pores likely hinder the nutrient supply of the cells resulting in reduced cell viability. Crosslinking reduced the swelling but, compared to the reference samples of pure bioactive glass, the cell viability is still significantly low.

45S5 BG (Reference) * 45S5 BG + Coll (uc) ** * 45S5 BG + Coll (cl) * ** *** 45S5 BG + Zein (uc) 45S5 BG + Zein (cl) * *** *** *** ** 150 * ** *** * *** * *

100

Cellviability [%] 50

0 7d 14 d 21 d

Figure 75: Cell viability of MG-63 cells (absorbance at 450 nm) on different samples after 7, 14 and 21 days. Significance levels: *p ˂ 0.05, **p ˂ 0.01, ***p ˂ 0.001 (Bonferroni’s post-hoc test was used).

The relative proliferation (n=4) was determined by an enzyme-linked immunosorbent assay (ELISA). For the measurement, a cell proliferation kit was used (Cell Proliferation ELISA, BrdU, Roche347).

Thymidine Bromodeoxyuridine nucleoside synthetic nucleoside

Figure 76: Schematically structure of thymidine (left) and its synthetic analog Bromodeoxyuridine (right - created by ChemBioDraw Ultra 13.0).

In vitro studies – a static cultivation system

This kit is based on the incorporation of Bromodeoxyuridine (BrdU) during the DNA synthesis. BrdU is a synthetic nucleoside, an analog of thymidine (Figure 76), which is part of the DNA. During the proliferation of the cells, DNA is newly synthesized and thymidine is supplemented by BrdU. Scaffolds were washed in PBS and moved to a new 48-well plate to ensure only cells attached on the scaffolds are included in the measurement. According to the manufacture’s protocol347, a BrdU labeling solution was prepared. Scaffolds were incubated with this solution for 2 h at 37 °C. During the incubation, BrdU is incorporated into the DNA of proliferating cells. Afterwards, cells were fixed for 30 minutes at room temperature. This step of denaturation is required to bind an antibody to the incorporated BrdU which was done subsequent to the fixation for another 30 min. After several washing steps, a substrate solution was added which enables the immunohistochemical detection of the antibody. This results in a change of color. The reaction was stopped after 5 min with 1 M H2SO4 and absorbance was measured at 450 nm using a microplate reader (PHOmo, Autobio Labtec Instruments).

45S5 BG (Reference) 45S5 BG + Coll (uc) *** 45S5 BG + Coll (cl) ** 45S5 BG + Zein (uc) *** 45S5 BG + Zein (cl) 250 *** *** ***

*** *** 200

150

100

Relative[%] proliferation 50

0 7d 14 d 21 d

Figure 77: Relative proliferation of MG-63 cells (absorbance at 450 nm) on different samples after 7, 14 and 21 days. Significance levels: **p ˂ 0.01, ***p ˂ 0.001 (Bonferroni’s post-hoc test was used).

Figure 77 shows the relative proliferation of MG-63 cells on the investigated samples. After 7 days, all samples showed a similar proliferation except scaffolds coated with uncrosslinked collagen. As already explained above, uncrosslinked collagen exhibits a high release rate. Cells, which are attached to the collagen matrix, are lost which causes a reduced cell number compared to the other samples. However, remaining cells were found to recover which results in a similar proliferation to that on pure 45S5 BG-based scaffolds after 21 days of incubation. Compared to pure BG, only uncrosslinked and crosslinked zein show statistical significance after 21 days.

In vitro studies – a static cultivation system

5.1.4 Cell morphology

The morphology of seeded MG-63 cells was analyzed by SEM observations and is represented in Figure 78. For SEM analysis, samples required a specific preparation step to fix the cells on the surface of the scaffold. Therefore, samples were washed with PBS and fixed with fixative I and II for 1 h, respectively. The composition for both fixatives can be found in Table 12. After fixation, scaffolds were subsequently dehydrated in a graded ethanol series (30, 50, 70, 80, 90, 95 and 99.8 vol.%) and dried in a critical point dryer (EM CPD300, Leica, DE).

45S5 BG - 7 d 45S5 BG - 14 d 45S5 BG - 21 d

45S5 BG + Coll (uc) - 7 d 45S5 BG + Coll (uc) - 14 d 45S5 BG + Coll (uc) - 21 d

45S5 BG + Coll (cl) - 7 d 45S5 BG + Coll (cl) - 14 d 45S5 BG + Coll (cl) - 21 d

45S5 BG + Zein (uc) - 7 d 45S5 BG + Zein (uc) - 14 d 45S5 BG + Zein (uc) - 21 d

In vitro studies – a static cultivation system

45S5 BG + Zein (cl) - 7 d 45S5 BG + Zein (cl) - 14 d 45S5 BG + Zein (cl) - 21 d

Figure 78: SEM images of seeded MG-63 cells on different types of scaffolds for 7, 14 and 21 days (at different magnifications).

In Figure 78, SEM images of different types of scaffolds can be seen which represent the morphology of seeded MG-63 cells after different time periods. The SEM images confirm the results of cell viability and relative proliferation. All samples showed widely spread cells with their typical spindle-shaped morphology of MG-63 cells which indicates a high biocompatibility for all coatings. After 21 days of incubation, even multilayer growth of MG-63 cells can be observed. Due to the homogenous distribution of cells, it is assumed scaffolds exhibit suitable pore size and porosity which enables the infiltration of the cells. However, cells seem to be denser on the outside of the scaffolds which can be attributed to the static cultivation system and the associated gradient of nutrient and oxygen supply. Zein-coated 45S5 bioactive glass-based scaffolds even exhibit sections inside the scaffold where no cells could be found. This can be explained by the partially blocked pores due to the swelled zein coating which additionally influences the nutrient supply of the cells, as mentioned above.

In summary, it can be concluded that all samples show good cell-material interactions visible by the well- flattened surface of osteoblast-like cells. However, it must also be noted that the present cell study provides only basic results. To get a profound knowledge about the in vitro compatibility of the tested systems and also of the interactions of the cells with the different applied coatings, more cell biology studies are required. Different assays (e.g. life-dead-staining or ALP activity) are suggested to investigate the cell-material interactions. In addition, the synthesis of extracellular proteins (e.g. collagen, fibronectin and vitronectin) and the expression of integrin have to be analyzed to reason the biocompatibility of the investigated system.

Table 12: Composition of fixative I and II (used for the fixation of cells for SEM observations).

Reactant Fixative I Fixative II Sodium cacodylate trihydrate 0.2 M 0.2 M Glutaraldehyde 0.1 wt.% 0.3 wt% Paraformaldehyde 2 wt.% 3 wt.% Sucrose 5 wt.% -

In vitro studies – a dynamic cultivation system

5.2 In vitro studies – a dynamic cultivation system

In this section, a dynamic cultivation system is introduced. Static cell cultivation offers some limitations regarding sufficient nutrient and oxygen supply, especially in three-dimensional systems. These drawbacks can be overcome by introducing a dynamic cultivation approach. Cells are continuously provided with fresh medium. In addition, cells are stimulated by mechanical forces which are known to be essential for the development of new tissue7 (e.g. bone).

5.2.1 Evaluation of a bioreactor

For the establishment of a dynamic cultivation system, preliminary work was carried out which included the determination of the test parameters. Details are explained below. Only chemicals, which are suitable for cell culture, were used and are listed detailed in appendix I.

Assembly of the system

The dynamic cultivation was carried out with commercially available bioreactors made of polycarbonate (Minucells and Minutissue GmbH, DE). Additional equipment includes silicon tubing with several connectors, a peristaltic pump (Ismatec IPC 8, IDEX Health & Science, DE), a thermo plate with cover lid (ME 12501, MEDAX Nagel, DE), glass bottles with screw caps and sterile filters. The individual parts were carefully cleaned with ultrapure water in an ultrasonic bath and sterilized in an autoclave at 121 °C for 20 minutes. Afterwards, the system was assembled, as represented in Figure 79 under sterile conditions.

Figure 79: Set-up of the bioreactor assembled for this experiment. Left: schematically set-up and right: original set-up.

The culture medium was stored in a bottle. At the top of the screw cap three drill holes are present. One hole is for the sterile filter which ensures the gas exchange. The other two holes are for the inlet and outlet of the silicone tubes. The bioreactor, which was tightly sealed with two metal clamps, was placed on a thermo plate with lid and heated at 37 °C. The medium was moved by a peristaltic pump into the bioreactor and back to the

In vitro studies – a dynamic cultivation system

bottle. This modus operandi enables a continuous run of the bioreactor without changing the medium. Thus, the risk of infections can be minimized. The composition of the medium as well as selection of parameters will be explained next.

Selection of parameters

For the establishment of a dynamic cultivation system different parameters have to be taken into account348. One important parameter is the constant pH in the medium. Inside the body a pH between 7.2 and 7.4 is

348 available which is ruled by solved CO2 and the amount of NaHCO3 .

NaHCO3 is also part of the cell culture medium. By the partial pressure of CO2 inside a CO2 incubator the pH is stabilized. However, the applied bioreactor system was designed to work in ambient atmosphere which contents only 0.3 % CO2 compared to a CO2 incubator with 5 % CO2. Therefore, to maintain a physiological pH value during the experiments the system required a CO2-independent buffer, e.g. HEPES.

0 9.2 9.0 8.8 8.6 8.4

8.2 5 8.0 pH 7.8 7.5 7.6 10 12.5 15 17.5 20 22.5 25 7.4 optimal pH range 7.2 7.2 - 7.4 7.0 6.8 0 5 10 15 20 25 HEPES concentration [mg/ml]

Figure 80: Determination of HEPES concentration to maintain a pH between 7.2 and 7.4 inside the bioreactor chamber.

For the dynamic cultivation system, RPMI 1640 medium was used. To determine the required HEPES concentration for the cultivation without CO2 injection, aliquots with increasing HEPES concentration

In vitro studies – a dynamic cultivation system

(0-25 mg/ml) were incubated at 37 °C in ambient atmosphere overnight348. The pH was measured with a pH electrode and is represented in Figure 80. Results revealed a HEPES concentration of 12.5 mg/ml to keep the pH in an optimal range (between 7.2 and 7.4). Another important parameter for the dynamic cultivation is the flow rate. The literature reports low flow rates, i.e. between 1 and 5.4 ml/h349–351, as well as high flow rates in the range of 0.3 – 5 ml/min352–355 which have been applied to similar systems. For the present dynamic cultivation system, a flow rate of 3 ml/h was chosen, considering experiences reported in the literature. To get information about the fluid mechanics inside the bioreactor chamber, the fluid dynamics was calculated considering that turbulent and laminar flows as hydrodynamic forces can have significant effects on the formation of engineered tissue7. The Reynolds number (Re) indicates the flow characteristics. The literature reports that Re > 2320356 implies instability of the laminar flow which leads to turbulent flow conditions. For the calculation of Re, equation [10] was used

v  d  ρ Re pores medium [Eq. 10]

where v is the flow velocity, dpores the diameter of the pores, ρmedium the density of the medium (assumed as 1 g/cm3) and η the dynamic viscosity of the medium (0.78 x 10-3 kg/ms357). The determination of the Reynolds number presents an approximation as the scaffolds do not exhibit an ideal geometric system. Re was calculated inside the scaffold which is the narrowest point and therefore the highest Re can be obtained. Pores were assumed to be ideal tubes with diameter of 350 µm. The conversion of the flow rate [ml/h] to the flow velocity v [mm/min] was simplified by

flo rate v [Eq. 11] empty A

where vempty is the flow velocity inside the empty bioreactor and A is the cross-section area of the inside of the bioreactor (height 16.5 mm, width 15 mm). The flow velocity vscaffold, with samples inside the bioreactor chamber, was calculated as

v v empty [Eq. 12] scaffold where P is the porosity of the scaffolds which was assumed to be 0.9. Results of the calculations can be extracted from Table 13.

In vitro studies – a dynamic cultivation system

Table 13: Calculated values for the Reynolds number Re in the bioreactor system.

Flow rate [ml/h] vempty [mm/min] vscaffold [mm/min] Rescaffold 3 0.2 0.22 0.0016

As already mentioned, the calculation of the Reynolds number is an approximation. The system is idealized as the pores are described as perfect tubes. It is assumed that, due to the interconnected pore structure, turbulent flows inside the real system occur. As this phenomena is not included in the idealization of system, which was used for the determinations of the Reynolds number, the calculated Re is too low. However, the Reynolds number shows values far below the critical limit of 2320, thus the fluid dynamics mode can still be assumed to be laminar. A last important parameter for the dynamic cultivation system is a constant temperature of 37 °C inside the bioreactor chamber which is heated by a thermo plate. Because the bioreactor is made of polycarbonate with a bottom thickness of ~ 8 mm it is expected that a temperature of 37 °C on the outside does not imply the same value inside the bioreactor. Therefore, different temperatures were selected for the thermo plate and compared with the temperature inside the bioreactor chamber. Results are represented in Figure 81.

Temperature outside [°C] outside Temperature 38 Flow rate: 3 ml/h 37

34 35 36 37 38 Temperature inside [°C]

Figure 81: Temperature establishment inside the bioreactor chamber.

In summary, the dynamic cultivation system was applied with a flow rate of 3 ml/h. To ensure a constant temperature of 37 °C inside the bioreactor, the thermo plate was heated at 41.5 °C. Cell culture medium RPMI 1640 was supplemented with 12.5 mg/ml HEPES to keep the pH between 7.2 and 7.4 during the cultivation in ambient atmosphere.

In vitro studies – a dynamic cultivation system

Cell cultivation

ST-2, a stromal cell line from mouse, was used for the dynamic cultivation. ST-2 cells are adherent cells received from bone marrow. They were cultured in RPMI 1640 supplemented with 10 vol.% FBS and 1 vol.%

PenStrep at 37 °C with 5 % of CO2 and 95 % humidity. At confluency ~ 80 % cells were trypsinized, centrifuged, re-suspended in fresh medium and counted by trypan blue exclusion method in a Neubauer chamber. Based on the results achieved in section 5.1, only three different kinds of scaffolds were chosen for the dynamic cultivation, namely (I) pure 45S5 BG-based scaffolds, (II) crosslinked collagen-coated scaffolds and (III) crosslinked zein-coated scaffolds. Before samples were seeded with cells, they were pretreated in DMEM (section 5.1.1 Sample preparation). Afterwards, scaffolds (I-III) were placed in a 48-well plate and seeded with 1 ml cell suspension containing 600.000 cells. Before they were transferred to the bioreactor, samples were cultured for 3 days in the incubator to allow the cells to attach. ST-2 cells are known to differentiate to osteoblasts when specific osteogenic factors are added to the medium352,358,359. Because of this, ST-2 cells are often used to investigate the potential of scaffolds to influence the osteoblastic differentiation behavior. It is hypothesized that a dynamic cultivation improves the differentiation conditions. Therefore, after 3 days of static cultivation, samples were washed with PBS and medium was supplemented with 50 µg/ml ascorbic acid, 10 mM β-Glycerophosphate and 100 nM Dexamethasone for osteogenic stimulation. One group of scaffolds was moved to the bioreactor to investigate the hypothesis of improved differentiation behavior in a dynamic cultivation system. The other group was left inside the incubator for comparison. The flow chart (Figure 82) gives an overview of the different steps of the experiment:

Scaffolds seeded with 600.000 ST-2 cells

Static cultivation for 3 days RPMI 1640 supplemented with 10 vol.% FBS and 1 vol.% PenStrep

+ osteogenic factors

Bioreactor Incubator

Dynamic cultivation Static cultivation

RPMI 1640 supplemented with 10 vol.% FBS, RPMI 1640 supplemented with 10 vol.% FBS, 1 vol.% PenStrep, 12.5 mg/ml HEPES, 50 µg/ml ascorbic 1 vol.% PenStrep, 50 µg/ml ascorbic acid, acid, 10 mM β-Glycerophosphate, 100 nM dexamethasone 10 mM β-glycerophosphate, 100 nM dexamethasone

Figure 82: Overview of experiments which were carried out to investigate osteogenic differentiation behavior on scaffolds (type I, II and III – description is provided in the text) of ST-2 cells. For comparison, a dynamic system and static cultivation were used.

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The addition of osteogenic factors to the medium was set as “day 0”. Samples were cultured for 7 and 14 days. Cell viability of the different samples was measured by the WST-8 metabolism (as described before) and lactate dehydrogenase (LDH) activity. Osteoblastic cell differentiation was evaluated by assessment of the specific alkaline phosphatase (ALP) activity. SEM observation was used to analyze the morphology of attached ST-2 cells.

5.2.2 Cell viability and LDH activity

The cell viability (n=4) was measured before samples were supplemented ith osteogenic medium (“day ”). In addition, measurements were done after 7 and 14 days. The cell viability was assessed by the enzymatically reduction of WST-8, as described in section 5.1.3. Absorbance was measured at 450 nm, using a microplate reader (PHOmo, Autobio Labtec Instruments). Results are presented in Figure 83.

0 d 7 d static 7 d dynamic 250 *** 14 d static *** *** 14 d dynamic *** *** * *** *** *** ** ***

200

150

100 Cellviability [%]

0 45S5 BG (Reference) 45S5 BG + Coll (cl) 45S5 BG + Zein (cl)

Figure 83: Cell viability of ST-2 cells (absorbance at 450 nm) on different samples after 0, 7 and 14 days. Significance levels: *p ˂ 0.05, **p ˂ 0.01, ***p ˂ 0.001 (Bonferroni’s post-hoc test was used).

At day 0, all samples showed similar cell viability. Pure 45S5 BG-based scaffolds demonstrate a significant increase of cell viability with incubation time. Furthermore, samples, which were cultivated under dynamic conditions, indicate higher cell viability compared to cells cultured under static conditions. Samples coated with crosslinked collagen exhibit significant higher values of cell viability after 7 days of incubation. However, after 14 days, the cell viability decreases which implies a reduced cell proliferation. It is supposed that due to the presence of enzymes (supplemented with FBS) the collagen coating degrades which entails the loss of attached cells. This effect is supported by the dynamic cultivation conditions. Nevertheless, the reduced cell proliferation may be also attributed to the enhanced differentiation of the cells. Zein-coated scaffolds reveal significant lower

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In vitro studies – a dynamic cultivation system values comparing results of the static and dynamic cultivation. Due to the swelling of zein, which was already reported earlier (section 4.2.2), pores are partially blocked which hinders a sufficient nutrient supply for the cells. By a dynamic cultivation system, the nutrient supply can be enhanced which results in an increased cell viability. Lactate dehydrogenase (LDH) is an enzyme present inside of cells. It transforms lactate to pyruvate while NAD is transformed to NADH. This reaction is reversible360. The amount of LDH is proportional to the cell number. For the measurement of the LDH activity, a lactic dehydrogenase based assay kit was used (In Vitro Toxicology Assay Kit, Sigma-Aldrich361). Cells, which were attached to the scaffolds (n=3), were lysed after 14 days of incubation using a lysis buffer (0.1 wt.% Triton X, 20 mM TRIS, 1 mM MgCl2 and 0.1 mM ZnCl2) and centrifuged. According to the manufacture’s protocol (TOX-7, Sigma-Aldrich361), a mastermix was prepared by mixing each 20 µl of dye solution, cofactor preparation and substrate solution. 140 µl of the supernatant was mixed with 60 µl of the mastermix and incubated for 30 minutes at room temperature. During incubation, NADH is utilized to transform a tetrazolium salt to formazan. The reaction was stopped by adding 300 µl of 1 M HCl. Absorbance was measured at 490 nm and 690 nm (background absorbance) via UV-Vis spectroscopy (Specord 40, Analytik Jena).

45S5 BG (Reference) 45S5 BG + Coll (cl) 45S5 BG + Zein (cl)

** *** * ** 250

200

150

100 LDH activity [%] LDHactivity

0 Static Dynamic

Figure 84: LDH activity of ST-2 cells (absorbance 490 nm - 690 nm) on different samples after 14 days of incubation. Significance levels: *p ˂ 0.05, **p ˂ 0.01, ***p ˂ . ( onferroni’s post-hoc test was used).

The results of the LDH activity after 14 days of incubation are presented in Figure 84. Results show a slightly improved LDH activity when samples were cultured under dynamic conditions. Only collagen-coated scaffolds exhibit a reduced LDH activity which can be attributed to the degradation of the collagen coating as mentioned

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In vitro studies – a dynamic cultivation system

above. Because LDH activity depends linearly on the number of cells, results indicate that a higher cell number was present on the zein-coated scaffolds compared to pure 45S5 BG-based scaffolds and collagen-coated scaffolds. However, Figure 83 revealed reduced cell viability for zein-coated scaffolds. Even though a higher cell number is present, the cell viability is slowed down due to the partially blocked pores indicating reduced cell proliferation which may be also attributed to the enhanced differentiation of the cells

5.2.3 ALP activity

Alkaline phosphatase (ALP), an enzyme present in cells, is a marker for osteogenic differentiation362–365. ALP activity can be measured by enzymatically transformation of a colorless para-Nitrophenylphosphate (pNPP) to a yellow p-Nitrophenol (pNP) which can be detected colorimetrically. Therefore, cells on the scaffolds (n=3) were lysed using a lysis buffer (as described in section 5.2.2) and centrifuged. An ALP buffer solution was prepared by dissolving 0.1 M TRIS, 2 mM MgCl2 and 9 mM pNPP in ultrapure water at a pH between 9.8 and 10 (adjusted with HCl)366. 250 µl of the supernatant was incubated with 100 µl of the ALP buffer solution at 37 °C for 1 h. The reaction was stopped by the addition of 650 µl NaOH. Absorbance was measured at 405 nm and 690 nm (background absorbance). The enzymatically transformation of pNPP is related to the time of incubation. Therefore, the specific ALP activity was calculated regarding the reaction time and normalized to the total protein amount of the samples. The protein amount was measured by the Bradford reagent367. Coomassie Brilliant Blue, a triarylmethane dye, forms complexes with proteins which yields a blue color. 25 µl of the supernatant was incubated with 975 µl of the Bradford reagent. The absorbance was measured after 10 minutes of incubation at 595 nm.

45S5 BG (Reference) 45S5 BG + Coll (cl) 45S5 BG + Zein (cl) 1.4

1.2

1.0

0.8

0.6

0.4

0.2 ALP activity [nmol pNp/min*mg protein] pNp/min*mg [nmol ALP activity

0.0 Static Dynamic

Figure 85: Specific ALP activity of ST-2 cells on different samples after 14 days of incubation.

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In vitro studies – a dynamic cultivation system

Figure 85 represents the specific ALP activity of ST-2 cells on different kinds of scaffolds after 14 days of incubation. A slight decrease in ALP activity can be recognized comparing static and dynamic cultivation, even though no significant difference could be determined amongst the samples. Results might suggest an enhanced osteogenic differentiation for samples stimulated in dynamic conditions, as reported in the literature before. A decrease in ALP activity is described during the cell differentiation which is followed by an increase when the cell matrix is mineralized368–372. In this case, reduced ALP activity indicates an enhanced differentiation behavior. Nevertheless, to retrieve more information about the osteogenic differentiation, further experiments are needed. This includes for example the determination of intracellular calcium368,369,373 as well as PCR analysis (polymerase chain reaction) to identify the expression of ALP, collagen I or osteocalcin which are markers of the osteogenic differentiation363,365,369,370,372,374,375.

5.2.4 Cell morphology

The cell morphology was investigated by SEM observation. For the analysis, samples were fixed by fixative I and II, dehydrated in a graded ethanol series and afterwards dried in a critical point dryer. The protocol can be found in section 5.1.4. In Figure 86 and 87, the cell morphology of ST-2 cells after 14 days of static and dynamic cultivation, respectively, is shown.

45S5 BG - static 14 d 45S5 BG + Coll (cl) - static 14 d 45S5 BG + Zein (cl) - static 14 d

Figure 86: SEM images of seeded ST-2 cells on different types of scaffolds after 14 days of static cultivation.

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In vitro studies – a dynamic cultivation system

45S5 BG - dynamic 14 d 45S5 BG + Coll (cl) - dynamic 14 d 45S5 BG + Zein (cl) - dynamic 14 d

Figure 87: SEM images of seeded ST-2 cells on different types of scaffolds after 14 days of dynamic cultivation.

In both cultivation methods well attached cells on the scaffold surface can be seen. However, samples, which were cultured under static conditions, show areas inside the scaffold where no cells could be found. This problem was partially solved by dynamic cultivation, especially in case of zein-coated scaffolds which exhibit some blocked pores due to the swelling of zein. In addition, ST-2 cells cultured in the bioreactor system show multilayer growth and seem to express a more osteoblastic morphology compared to cells incubated without dynamic stimulation which indicates successful osteogenic differentiation. Summarizing the results, it can be concluded that the enhancement of the osteogenic differentiation by a dynamic cultivation is not clear yet. The present study supplies only basic results on the establishment of a bioreactor system. To be able to generate more fundamental knowledge about the differentiation behavior in dynamic conditions, more elaborated experiments have to be carried out as already explained above (e.g. determination of intracellular calcium as well as PCR analysis for the identification of expressed proteins in the extracellular matrix).

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Comparison of the investigated systems

5.3 Comparison of the investigated systems

The aim of this study was the fabrication and characterization of a series of 45S5 BG-based composite scaffolds for the possible application in bone tissue engineering. Therefore, scaffolds have to meet some demands which are described in the following section. The definition for an ideal scaffold for bone tissue engineering is still under discussion. The literature describes typical criteria which include requisites related to porosity, bioactivity and mechanical properties amongst others15,31,38,376. Salgado et al.38 define the following properties as essential: biocompatibility, porosity, pore size, surface properties, osteoinductivity, mechanical properties and biodegradability. The investigated systems are contrasted with some of the requirements to estimate their potential for bone tissue engineering applications.

5.3.1 Polymeric scaffolds with and without ceramic reinforcement

Within the framework of this thesis, different polymeric scaffolds were prepared and characterized. Briefly summarized, porous collagenous scaffolds were prepared by lyophilization (section 3.1). In addition, zein-based porous structures were created by the so-called salt leaching technique (section 4.1). For the enhancement of the bioactive behavior, both kinds of scaffolds were reinforced with 45S5 bioactive glass particles. The polymeric scaffolds were characterized in terms of microstructure, swelling properties, degradation behavior and bioactivity. Results can be found in the corresponding chapter (chapter 3 and 4). In this section, collagen- based scaffolds and zein-based porous structures are compared amongst each other and evaluated.

Pore size and porosity

To enable and support bone tissue ingrowth as well as vascularization, an interconnected porous structure of the scaffold is required377. Next to this, the pore size plays a key role for the nutrient supply of cells. Due to insufficient pore size, pores can be blocked which hinders the migration of cells. In the literature, the optimal mean pore size is controversially discussed as it also influences the mechanical properties of the scaffolds38. Some studies378–380 claim a pore size > 100 µm is not required for bone ingrowth whereas pores exceeding 300 µm are suggested by others381–383 to enable bone formation and vascularization. During this study, porous collagen scaffolds with pores in the range of 100 – 500 µm and 98 % porosity were produced via lyophilization. After the reinforcement of the collagenous structure with 45S5 bioactive glass particles, scaffolds still exhibited relatively high porosity (~ 95 %) and pore sizes between 50 and 150 µm were measured. Similar systems are described in the literature. A selection of relevant previous work is given in Table 14. Next to porous collagen structures, zein-based scaffolds were also produced in this study by salt leaching. Results showed zein structures with micropores > 1 µm on the surface and macropores of size ranging from

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Comparison of the investigated systems

100 to 400 µm (porosity ~ 75 %). After the reinforcement with bioactive particles, the overall porosity was determined to be ~ 84 %. Only few studies based on the fabrication of porous zein scaffolds have been published and can be found in Table 14. It can be stated that pore size and porosity are important parameters for cell migration and tissue ingrowth. However, it is difficult to set a specific value as the cell response not only depends on the overall porosity but on the entirety of the applied system, e.g. type of tissue and applied material. Therefore, a tunable pore size offers the possibility to tailor a given system according to the requirements. In case of collagen scaffolds, the pore size can be controlled by the freezing temperature while pore diameter of zein scaffolds depends on the size of the applied porogen.

Table 14: Pore sizes and porosity of collagen-based and zein-based scaffolds described in the literature.

Material Fabrication Pore size Porosity Application Reference Collagen- Murphy et Lyophilization 85 – 325 µm 99 % Bone tissue engineering glycosaminoglycan al.382,384 Cartilage tissue Collagen Lyophilization 150 ± 25 µm Oliveira et al.157 engineering Collagen + BG + Lyophilization 300 µm 70 – 85 % Bone tissue engineering Xu et al.385 phosphatidylserine Collagen + Lyophilization 150 – 250 µm ~ 98 % Tissue engineering Xu et al.166 poly-L-lactide Collagen + Lyophilization 100 – 200 µm Skin tissue engineering Ma et al.160 chitosan Collagen + Adipose tissue Davidenko et Lyophilization 100 – 220 µm 85 % hyaluronic acid engineering al.161 Collagen- Lyophilization ~ 200 µm 98 – 99.5 % Skin tissue engineering Liang et al.386 chondroitin sulfate Collagen + nano- Lyophilization ~ 99 % Bone tissue engineering Cunniffe et al.162 hydroxyapatite Zein + HA Porogen leaching > 75 % Bone tissue engineering Qu et al.320 Zein Porogen leaching 100 – 380 µm Bone tissue engineering Wang et al.321 Zein Porogen leaching 100 – 300 µm 75 – 79 % Bone tissue engineering Gong et al.156

Mechanical properties and biodegradability

For the application in bone tissue engineering, suitable mechanical properties of scaffolds are needed to ensure an easy handling without destroying the structure9,387,388. Furthermore, scaffolds have to withstand hydrostatic pressure during in vitro experiments. In vivo, scaffolds are exposed to permanent stress until the tissue regeneration process is completed. Hence, insufficient mechanical stability results in the fracture of the scaffolds whereas higher stiffness causes stress shielding which entails loss of bone tissue and loosening of the implant. In addition, the degradation rate should be in accordance with the formation of new tissue. To meet these demands, significant research is carried out9,10,15,38,389,390. Porous matrices based on neat natural polymers can

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Comparison of the investigated systems hardly reach compressive strength values close to those of cancellous bone (2 – 12 MPa51,306). To overcome this problem, composite materials of polymer matrices with ceramic reinforcements are designed to merge the positive characteristics from both material classes31,32,58,391,392. Thereby, the biodegradability and chemical versatility of polymers can be combined with bioactive ceramics and glasses to enhance the mechanical properties as well as the bioactive behavior (as discussed in the next section). As a matter of fact, nature is acting likewise by combining collagen and apatite in bone tissue. By the reinforcement of natural polymer matrices with ceramic fillers, the mechanical properties can be positively changed, as presented by Sarker et al.76 and Venkatesan et al.52 who reviewed collagen and alginate composite scaffolds, respectively, for bone tissue engineering applications. Also the mechanical properties of synthetic polymers can be enhanced by incorporating inorganic fillers, as reported by Maquet et al.393 who described the improvement of porous poly-

(D,L-lactide) (PDLLA) and poly(lactide-co-glycolide) (PLGA) scaffolds after the addition of 45S5 bioactive glass. This behavior is different to porous zein scaffolds which were produced during this study. Pure zein showed compressive strength of 4.1 ± 0.8 MPa whereas the compressive strength decreased to 2.2 ± 0.9 MPa after the reinforcement with 45S5 bioactive glass particles which is likely due to the reduced plasticization of the zein matrix (section 4.1.5). However, both values are in the range of data reported for the mechanical strength of bone tissue (2 – 12 MPa51,306). Besides this, the incorporation of bioactive glass stimulates the bioactive behavior of the natural derived polymer which is discussed next. In relation to collagen scaffolds, considering their hydrogel status, the mechanical properties of the developed porous collagen scaffolds seem to be more relevant for soft tissue engineering applications and are not presented here. However, by the reinforcement with 45S5 BG particles, the collagen matrix was stabilized which resulted in a decelerated degradation behavior and consistent geometry of the scaffolds during immersion in water-based solutions (section 3.1.3).

Bioactive behavior

Bioactivity describes the ability of a material to interact with living tissue, as reported in detail in section 2.2.1. In the case of bioactive glasses, the formation of a SiO2-rich layer followed by the precipitation of a CaO-P2O5 film, which crystallizes to form carbonated hydroxyapatite, was proposed by Hench et al. as the “mar er” of bioactivity124. For the evaluation of the biomineralization process, SBF studies were carried out in vitro, as described before (section 3.1.4). Pure collagen provides nucleation sites like –COOH which facilitate the interaction with Ca2+ and therefore the formation of a hydroxyapatite layer251,304. After the addition of bioactive glass particles to the collagen matrix, an accelerated growth of hydroxyapatite crystals was observed due to the combined effect of both materials. In contrast, pure zein did not show any bioactive characteristics after the immersion in SBF. A common method to induce and promote hydroxyapatite formation on a non-bioactive material is the addition of a bioactive ceramic392 like bioactive glass or hydroxyapatite. This principle was also applied during this study by mixing zein

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Comparison of the investigated systems

with 45S5 BG which resulted in a novel bioactive composite material, not reported in the literature up to now. Comparing the two kinds of developed polymeric scaffolds under the above mentioned aspects, both systems offer high potential for the application in bone tissue engineering. Next to this, zein-based scaffolds as well as collagen-based composites offer simplicity during manufacturing. However, porous collagen matrices exhibit very low mechanical properties and therefore they seem to be more relevant for soft tissue engineering. On the other hand, zein represents an option for bone tissue engineering. In spite of its natural origin, porous zein scaffolds reveal good compressive strength in the region of cancellous bone. Also the stability against enzymatic degradation indicates its suitability for biomedical applications. Nevertheless, the reduced plasticization of the zein matrix after bioactive glass reinforcement, which results in lower mechanical stability and faster enzymatic degradation, has to be adapted as the bioactive behavior is paramount in bone tissue engineering applications.

5.3.2 Biopolymer-coated 45S5 bioactive glass-based scaffolds

A significant drawback of 45S5 bioactive glass-based scaffolds is their high brittleness and relatively low fracture strength. As reviewed by Philippart et al.336 and Yunos et al.30, this problem can be overcome by toughening of scaffolds with different polymer coatings. During this study, two different biopolymers were chosen as coating material, namely collagen (section 3.2) and zein (section 4.2). Results of the characterization of the coated scaffolds can be found in the corresponding chapters (chapter 3 and 4). In this section, the two different coated scaffolds are compared and evaluated in terms of their potential as bone composite scaffolds.

Pore size and porosity

As already discussed, pore size and porosity are main parameters controlling cell migration and tissue ingrowth. 45S5 bioactive glass-based scaffolds were produced by the foam replica technique originally described by Chen et al.29. Pore size and porosity can be adjusted, in principle, by the applied polyurethane template foam. In case of polymer coatings on porous structures it has to be considered that the overall porosity should not be significantly hampered by clogging of the pores. As-fabricated bioactive glass-based scaffolds revealed a highly interconnected porosity (~ 94 %) after the sintering process with macropores in the range of 250-500 µm. Collagen coating on the porous bioactive glass-based structure was applied by a combined process of surface functionalization and collagen immersion whereas zein coatings were simply obtained by dip coating. Neither collagen nor zein coating affected the pore size significantly. Therefore, concerning the final scaffold porosity, both coating materials led to satisfactory results.

Mechanical properties

As indicated above, the application of polymer coatings aims primarily the improvement of the mechanical properties of porous scaffold structures by filling and bridging microcracks on the surface of the struts. The

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Comparison of the investigated systems literature mainly describes synthetic polymers as coating material, e.g. polycaprolactone (PCL)394–396 or poly(D,L-lactic acid) (PDLLA)310,397,398. But also natural derived proteins are attracting attention as polymer coating for scaffolds, e.g. gelatin399,400, silk401 or chitosan402. In this study, zein and collagen were applied which have not been largely discussed in literature247,300,403. As-fabricated 45S5 bioactive glass-based scaffolds showed compressive strength of 0.04 ± 0.02 MPa. After the applied collagen coating, samples exhibited values of 0.21 ± 0.03 MPa and 0.18 ± 0.02 MPa for uncrosslinked and crosslinked scaffolds, respectively. The zein coating revealed compressive strengths of 0.21 ± 0.02 MPa and 0.19 ± 0.03 MPa for uncrosslinked and crosslinked samples, respectively. Both coating materials showed similar results as the compressive strength was substantially improved. However, the received values are still low and are at the lower bound for spongy bone which are reported to be between 0.2 and 4 MPa when the relative density of bone tissue is ~ 0.129. Therefore, the coated scaffolds are only interesting for non-load bearing applications. Nevertheless, it has to be considered that in case of zein the coating is very inhomogeneous. It is assumed that the compressive strength could be further increased by a uniform and more homogenous distribution of the polymeric coating. In such case, however, the bioactivity may be lost because pure zein does not show any bioactive behavior (section 4.1.4). Hence, a compromise between bioactivity and mechanical stability has to be found by choosing the appropriate coating type and coating microstructure.

Bioactive behavior

As already mentioned, bioactivity is an important parameter to evaluate the interaction of a material with living tissue. As-fabricated 45S5 bioactive glass-based scaffolds are highly bioactive meaning that the applied coating should not negatively affect this characteristic. Because collagen offers nucleation sites for the formation of hydroxyapatite, the bioactivity of collagen-coated BG-based scaffolds was not influenced. Bioactivity studies were carried out in SBF (section 3.2.3) and revealed the formation of a significant amount of HA crystals which were well embedded in the collagen coating. Because pure zein does not show any indication of biomineralization, it is highly likely that the bioactivity is negatively affected. However, due to the inhomogeneous distribution of the zein coating inside the BG-based scaffold, the bioactive behavior was not influenced as parts of the BG-based scaffold struts were directly exposed to SBF (section 4.2.3). Therefore, with both types of coatings (collagen and zein), the bioactive behavior of as-fabricated 45S5 bioactive glass- based scaffolds could be preserved.

Biocompatibility and osteoinduction

Next to bioactivity, also biocompatibility plays an important role for tissue engineering applications. Biocompatibility describes the incorporation of a biomedical system into an organism without evoking an immune response which can be regulated by different parameters404. The biocompatibility can be evaluated in

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Comparison of the investigated systems

vitro by the interaction with cells. During this study, the biocompatibility was tested by seeding collagen-coated and zein-coated scaffolds with MG-63 cells, as described in detail in section 5.1. As already reported in the literature, 45S5 bioactive glass405–407, collagen408,409 and zein156,320,341 in separated form demonstrate all good biocompatible behavior. Therefore, similar results were expected with the developed composites which showed positive cell viability and well attached MG-63 cells on the surface of the scaffolds. Furthermore, collagen and zein are described in the literature to support osteoblastic differentiation7,319,320,410,411 which is part of the osteoinduction process412. This hypothesis was investigated during this study by seeding ST-2 cells in presence of osteogenic factors (details can be found in section 5.2). Although indications for the enhanced osteoblastic differentiation behavior were found on both systems (collagen and zein coating), extended experiments need to be carried out which are explained in the next chapter (future work).

It can be concluded, therefore, that natural derived polymers represent a very interesting approach as coating material for bioactive glass-based scaffolds. Next to mechanical properties also the cell response can be influenced. Even though the mechanical performance is not sufficient for load bearing applications, the composite scaffolds are interesting from the material’s point of vie . In addition, natural polymer coatings offer a high versatility and provide the possibility to incorporate biomolecules or drugs.

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Conclusion

6 Conclusion and future work

The present thesis has described two different approaches to design composite scaffolds for the application in bone tissue engineering. On the one hand, composites were developed based on the combination of collagen and 45S5 bioactive glass. On the other hand, 45S5 bioactive glass was also used in combination with zein. One novelty of this work was the introduction of a new coating method for porous structures with collagen. 45S5 bioactive glass-based, highly porous scaffold, which were fabricated by the foam replica technique, were coated with collagen by a combined coating procedure. First, the surface of the scaffolds was silanized to enable a chemical bonding between collagen and the bioactive glass surface. Afterwards, the functionalized samples were immersed in a collagen solution, left for gelling at 37 °C and dried at room temperature. To stabilize the collagen coating, an additional crosslinking step was introduced. The important advantage achieved applying this coating procedure is that layer thicknesses of a few micrometers can be realized while ensuring the macroporosity of the scaffold structure is not affected. The literature already describes collagen coatings of 45S5 BG-based scaffolds247. However, only a thin monolayer of collagen fibers could be deposited on the surface of the struts. One aim of polymer coatings is the improvement of mechanical performance of the scaffolds which could not be achieved with a monolayer coating. By the introduced coating procedure, values of compressive strength were enhanced by a factor of 5 (namely uncoated scaffolds 0.04 ± 0.02 MPa, crosslinked collagen- coated scaffolds 0.18 ± 0.03 MPa). These values are at the lower boundary of the compressive strength of bone and are mainly attributed to the high porosity and brittle nature of the scaffold. Although interesting results were achieved, major challenges remain. One problem is the hydrogel character of collagen in contact with water. Even if collagen coatings achieve high mechanical stability, the integrity of the scaffolds will be compromised after implantation. Therefore, collagen coatings are more relevant as vehicle for drug delivery than for high mechanical reinforcement effects. Future work in this field should focus on the encapsulation of biomolecules like antibiotics for anti-inflammation or for the treatment of osteomyelitis413 inside the collagen coating and the investigation of their release behavior. Furthermore, by modifying the degree of crosslinking, the release rate can be controlled. During the second approach, zein was combined with bioactive glass particles. By the so-called salt leaching technique novel bioactive composites with compressive strength of 2.2 ± 0.9 MPa, interconnected porosity of around 84 % and high stability against enzymatic degradation were developed. In the literature, the fabrication of stable and porous zein scaffolds is already described156,321. However, the significant drawback of pure zein for tissue engineering applications is the lack of bioactivity of this vegetal protein. By the combination with bioactive glass this problem was overcome, as reported in this work. Nevertheless, it has to be considered that by the presence of 45S5 bioactive glass the plasticization process of zein is influenced which results in lower compressive strength and reduced stability against enzymatic degradation compared to pure zein. An interesting

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Conclusion approach for further studies is the development of zein composites with enhanced bioactivity and improved mechanical properties at the same time. This may be realized by the optimization of the bioactive glass content. Also the size of the bioactive glass particles should be considered. Furthermore, zein can be combined with different polymers to increase the mechanical performance, e.g. PCL. Also other bioactive materials next to 45S5 bioactive glass can be incorporated to enhance the biomineralization process without affecting the plasticization process of zein. Because zein was already described as drug delivery vehicle326,414, another interesting approach is the encapsulation of drugs into the polymer matrix in order to develop bioactive scaffolds with drug delivery capability. Next to this, the crosslinking mechanism of zein still raises numerous questions which have not been properly addressed. In this study, zein-coated and collagen-coated bioactive glass-based scaffolds were also tested in vitro. During static cultivation with human osteoblast-like cells (MG-63), results revealed the expected biocompatibility of the investigated systems. However, the dynamic cultivation with a stromal cell line from mouse (ST-2), which was carried out to facilitate the osteogenic differentiation, did not provide clear results in terms of differentiation behavior. One reason for this observation is likely the limited investigated incubation time (14 days). To provide a clear statement of the differentiation behavior, extensive studies have to be performed which should also consider the optimization of the parameters of the bioreactor system (e.g. different flow rates), longer incubation periods and the determination of intracellular calcium as well as PCR analysis (polymerase chain reaction) to identify the expression of ALP, collagen I or osteocalcin (markers for osteogenic differentiation). Furthermore, elaborated cell studies have to be carried out with different cell types. Different assays (e.g. life- dead-staining or ALP activity) are suggested to investigate the cell-material interactions. In addition, the synthesis of extracellular proteins (e.g. collagen, fibronectin and vitronectin) and the expression of integrin have to be analyzed to reason the biocompatibility of the investigated system.

Overall, the results of this thesis have demonstrated the successful development of composite scaffolds based on natural derived proteins and 45S5 bioactive glass. The results provide a well-founded basis for further investigations in this field, as outlined above. Finally, it can be summarized that both zein and collagen show promising potential for the application in tissue engineering and further comparative studies, with and without combination with bioactive inorganic materials, are required to be able to conclusively determine if zein can replace collagen in applications as tissue engineering scaffolds.

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Appendix I - Chemicals and materials

Chemicals:

2-Propanol 2-propanol Technical, VWR Chemicals (Germany), CAS Number: 67-63-0 β-Glycerophosphate β-Glycerophosphate disodium salt hydrate, Sigma-Aldrich (Germany), CAS Number: 154804-51-0 Acetone Acetone Technical, VWR Chemicals (Germany), CAS Number: 67-64-1 APTS (3-Aminopropyl)triethoxysilane 99 %, Sigma-Aldrich (Germany), CAS Number: 919-30-2 Ascorbic acid L(+)-Ascorbic Acid, AppliChem (Germany), CAS Number: 50-81-7 Bioactive glass Commercially available bioactive glass with the 45S5 composition (Vitryxx®, Schott, Germany) Bradford Bradford Reagent (B6916), Sigma-Aldrich (Germany)

CaCl2  2 H2O Calcium chloride dihydrate, AnalaR NORMAPUR®, VWR Chemicals (Germany), CAS Number: 10035-04-8 Cell Counting Cell Counting Kit - 8 (96992), Sigma-Aldrich (Germany) Cell Proliferation Cell Proliferation ELISA, BrdU (11647229001), Roche (Germany) Collagen Collagen G1 from bovine calf skin, Matrix BioScience (Germany) Collagenase Collagenase from Clostridium histolyticum, lyophilized powder  125 CDU/mg solid, 0.5- 5.0 FALGPA units/mg solid, Sigma-Aldrich (Germany), CAS Number: 9001-12-1 DMEM 10x DMEM (F 0455), Biochrom (Germany) DMEM without phenol red (D5921), Sigma-Aldrich (Germany) Cell culture: DMEM (31885-023), Thermo Fisher Scientific (Germany) Dexamethasone Dexamethasone, Sigma-Aldrich (Germany), CAS Number: 50-02-2 EDC N-(3-Dimethylaminopropyl)-N’-ethylcarbodiimide hydrochloride  98.0 %, Sigma-Aldrich (Germany), CAS Number: 25952-53-8 EDTA Ethylenediaminetetraacetic acid, ACS reagent, 99.4-100.6 %, powder, Sigma-Aldrich (Germany), CAS Number: 60-00-4 Ethanol Ethanol denatured with about 1% methyl ethyl ketone, EMSURE®, Merck Millipore (Germany), CAS Number: 64-17-5 FBS Fetal Bovine Serum (F0804), Sigma-Aldrich (Germany) Folin Folin & Ciocalteu’s phenol reagent (F9252), Sigma-Aldrich (Germany) Glutaraldehyde Glutaraldehyde solution 50 %, AppliChem (Germany), CAS Number: 111-30-8 Glycerin Glycerin 87 % BioChemica, AppliChem (Germany), CAS Number: 56-81-5 Glycine Glycine  99 %, ReagentPlus®, Sigma-Aldrich (Germany), CAS Number: 56-40-6 HCl Hydrochloric acid, 37 %, VWR Chemicals (Germany), CAS Number: 7647-01-0 HEPES HEPES  99.5 %, Sigma-Aldrich (Germany), CAS Number 7365-45-9

H2SO4 Sulfuric acid 95-97 %, EMSURE® ISO, Merck Millipore (Germany), CAS Number: 7664- 93-9 KBr Potassium bromide for IR spectroscopy Uvasol®, Merck Millipore (Germany), CAS Number: 7758-02-3 KCl Potassium chloride, EMSURE®, Merck Millipore (Germany), CAS Number: 7447-40-7

114

Appendix I - Chemicals and materials

K2HPO4  3 H2O Potassium phosphate dibasic trihydrate, ReagentPlus®,  99.0 %, Sigma-Aldrich (Germany), CAS Number: 16788-57-1 KV 9062 Zschimmer & Schwarz GmbH & Co KG (Germany) LDH In Vitro Toxicology Assay Kit, Lactic Dehydrogenase based, (TOX7), Sigma-Aldrich (Germany) L-Glutamine L-Glutamine 200 mM solution, Sigma-Aldrich (Germany), CAS Number: 56-85-9 Lowry Lowry Reagent (L3540), Sigma-Aldrich (Germany) MES MES hydrate  99.5 %, Sigma-Aldrich (Germany), CAS Number: 1266615-59-1

MgCl2  6 H2O Magnesium chloride hexahydrate, ACS reagent, 99.0-102.0 %, Sigma-Aldrich (Germany), CAS Number: 7791-18-6 NaCl Sodium chloride, VWR Chemicals (Germany), CAS Number: 7647-14-5

NaHCO3 Sodium bicarbonate, ACS reagent,  99.7 %, Sigma-Aldrich (Germany), CAS Number: 144-55-8

NaN3 Sodium azide, Merck Millipore (Germany), CAS Number: 26628-22-8 NaOH Sodium hydroxide pellets pure, Merck Millipore (Germany), CAS Number: 1310-73-2

Na2HPO4 Sodium phosphate dibasic  99.0 %, Sigma-Aldrich (Germany), CAS Number: 7558-79-4

Na2SO4 Sodium sulfate anhydrous, AnalaR NORMAPUR®, VWR Chemicals (Germany), CAS Number: 7757-82-6 NHS N-Hydroxysuccinimide 98 %, Sigma-Aldrich (Germany), CAS Number: 6066-82-6 Ninhydrin Ninhydrin, Sigma-Aldrich (Germany), CAS Number: 485-47-2 Paraformaldehyde Paraformaldehyde 95 %, Sigma-Aldrich (Germany), CAS Number: 30525-89-4 PenStrep Penicillin-Streptomycin (15140-122), Thermo Fisher Scientific (Germany) PBS PBS tablets, each tablet prepares 100 ml of 10 mM PBS, VWR Chemicals (Germany) Cell culture: Dulbecco's phosphate-buffered saline (DPBS, 14190-094), Thermo Fisher Scientific (Germany) pNPP Phosphatase substrate, Sigma-Aldrich (Germany), CAS Number: 333338-18-4

PVA Polyvinyl alcohol, fully hydrolyzed, Mw  30000, Merck Millipore (Germany), CAS Number: 9002-89-5 RPMI RPMI 1640 Medium (61870-010), Thermo Fisher Scientific (Germany) Silver paste Conductive silver paste, Plano (Germany) Sodium cacodylate Sodium cacodylate trihydrate  98 %, Sigma-Aldrich (Germany), CAS Number: 6131-99-3 trihydrate Sucrose Sucrose  99.5%, Sigma-Aldrich (Germany), CAS Number: 57-50-1 Toluene Toluene AnalaR NORMAPUR®, VWR Chemicals (Germany), CAS Number: 108-88-3 TRIS Tris(hydroxymethyl)aminomethane AnalaR NORMAPUR®, VWR Chemicals (Germany), CAS Number: 77-86-1 TRIS-HCl Trizma® hydrochloride  99.0 %, Sigma- Aldrich (Germany), CAS Number: 1185-53-1 Trition X Triton™ X-100, Sigma-Aldrich (Germany), CAS Number: 9002-93-1 Trypan blue 0.4 % Trypan Blue solution, Sigma-Aldrich (Germany), CAS Number: 72-57-1 Trypsin 0.25 % Trypsin-EDTA (25200-056), Thermo Fisher Scientific (Germany) UPW Electrical conductivity of ≤ . 67 µS/cm, Purelab R 7, Veolia (Germany)

115

Appendix I - Chemicals and materials

Zein Zein, Sigma-Aldrich (Germany), CAS Number: 9010-66-6

ZnCl2 Zinc chloride, Sigma-Aldrich (Germany), CAS Number: 7646-85-7

Materials:

Bioreactor (incl. tube Minucells and Minutissue GmbH (Germany) connectors, screw caps and sterile filters) PU foam PU fully reticulated, 45 ppi, PL Bulpren S28133, Eurofoam Deutschland GmbH Schaumstoffe (Germany) Silicone tubing VWR Chemicals (Germany)

116

Appendix II - Characterization methods

To analyze the developed samples, different characterization methods were chosen. The operation mode of the applied equipment is shortly described and explained above.

Scanning electron microscopy

Scanning electron microscopy (SEM) is a common method to investigate the surface morphology of different samples. Therefore, a focused electron beam scans the surface of an object. By interactions of the electrons with atoms in the sample various signals are detected which produces an image. Samples were analysed by a high resolution scanning electron microscope (SEM, Auriga, Zeiss, Germany), operated at 2 kV and working distance ~ 5 mm. Scaffolds were fixed on a sample holder by conductive silver paste and sputtered with gold (Quorum Q150T S, United Kingdom) to avoid charging.

Contact angle

Contact angle measurements can be performed through the deposition of a water droplet on a flat surface. By establishing a tangent of the solid surface and the liquid drop, the contact angle  can be measured (Figure 88) and properties of the surface (e.g. surface energy or wettability) can be determined. Contact angles   90 ° indicate a surface with hydrophilic properties whereas angles   90 ° represent a hydrophobic character415,416. The experiments were carried out by a contact angle device DSA30 (Krüss GmbH, Germany).

Figure 88: Measurement of the contact angle.

X-ray photoelectron spectroscopy

X-ray photoelectron spectroscopy (XPS) is a technique for surface analysis and can provide information about elemental composition (except hydrogen and helium), chemical bondings and oxidation state417,418. The sample is exposed to electromagnetic radiation (X-ray) through which electrons in atoms near to the surface (5 – 30 Å) are excited. Photoelectrons are emitted of the material which can be detected energy dispersive. Figure 89 shows the schematic construction of an X-ray photoelectron spectroscope.

117

Appendix II - Characterization methods

Figure 89: Schematic construction of an X-ray photoelectron spectroscope.

The physical principle of photoemission spectroscopy can be explained through the photoelectric effect. In order to determine the characteristic binding energy EB of electrons in atoms they have to be excited. Sample and spectrometer exhibit the same fermi level due to electrical contacts (Figure 90A).

A) B) x 105 4.5

4.0 Na1s 3.5

O1s 3.0

2.5 Na KLL O O KLL

O KLL c/s 2.0

1.5 Na KLL

1.0 Ca2p3Ca2p1 Ca2p

0.5 Ca2s

C1s

P2s Si2s Si2p P2p 0.0 1200 1000 800 600 400 200 0 Binding energy [eV]

Figure 90: A) Schematic energy diagram. B) Due to characteristic peaks in the XPS spectrum, elements on the sample surface can be identified. Here, for instance, the XPS spectrum of an as-sintered pure bioactive glass sample produced during this study.

The work function  of the sample and of the spectrometer Sp are different. An excitation energy h between

100 and 1500 eV is applied, which is higher than , and a photoelectron with kinetic energy E'kin is emitted from the sample. Before entering the spectrometer, the photoelectron has to overcome the potential difference Sp - . If Sp is smaller than , the photoelectron is speeded up and the kinetic energy Ekin is measured in the spectrometer. The binding energy EB can be determined by

EB = h - Ekin - Sp [Eq. 13]

118

Appendix II - Characterization methods

The work function Sp of the spectrometer is a constant and can be assessed by the measurement of a material with known EB. The measured binding energy can be plotted against the detected number of electrons (counts per second). The emerging XPS spectrum shows characteristic peaks (Figure 90B) which are related to the electron configuration of the elements and so elements, which are present on the surface of the investigated material, can be identified. Because the detected number of electrons corresponds to the amount of element, also the atomic concentration of the elements can be determined. For the measurement, a multi-technique XPS (Phi- 5600) was used (Al-Kα radiation).

Fourier transform infrared spectroscopy

By Fourier transform infrared spectroscopy (FTIR), chemical bonds can be identified. Therefore, infrared radiation is used. Sample molecules selectively absorb the infrared radiation and are excited. The dipole moment of the sample molecules change which results in molecular vibrations. Frequency of vibration can be recorded by a detector. The recorded interferogram can be transferred into a spectrum by Fourier transform. Peaks correspond to a specific energy level and are related to a particular bond type. Samples were analyzed by a Nicolet 6700 FTIR spectrometer (Thermo Scientific, USA).

UV-Vis spectroscopy

Ultraviolet-visible (UV-Vis) spectroscopy is a simple technique and was used in framework of this thesis to determine the different concentrations of proteins and free amino groups in solutions colorimetrically.

Data

1.0 R = 0.98592 0.8

0.6

0.4 Absorbance 0.2

50 100 150 200 250 300 350 400 450 Concentration [µg/ml]

Figure 91: Schematic diagram of UV-Vis spectroscopy. By plotting absorbance against concentration a calibration curve can be drawn (regarding to manual).

119

Appendix II - Characterization methods

Therefore, solutions to be measured are filled into cuvettes. By electromagnetic radiation, valence electrons in the sample are excited. The adsorption of the energy is recorded by a detector (Figure 91). The absorbance is plotted against a known protein concentrations to establish a calibration curve which is necessary to ascertain the unknown concentration value. For the measurement of solutions with volume ~ 1 ml, cuvettes were placed a single beam UV-Vis spectrophotometer (Specord 40, Analytik Jena, DE). Solutions  150 µl were filled into a 96-well plate and analyzed by a microplate reader (PHOmo, Autobio Labtec Instruments).

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