Development of multifunctional wound dressings by using newly designed ion-doped borosilicate and borate bioactive glasses
Entwicklung multifunktionaler Wundauflagen unter Verwendung neu gestalteter ionendotierter bioaktiver Borosilikat- und Boratgläser
Der Technischen Fakultät der Friedrich-Alexander-Universität Erlangen-Nürnberg zur Erlangung des Grades
Doktor-Ingenieur
vorgelegt von
Katharina Schuhladen aus Donauwörth
Als Dissertation genehmigt von der Technischen Fakultät der Friedrich-Alexander-Universität Erlangen-Nürnberg
Tag der mündlichen Prüfung: 30.09.20
Vorsitzende des Promotionsorgans: Prof. Dr.-Ing. habil. Andreas Paul Fröba
Gutachter:
Prof. Dr.-Ing. habil Aldo Boccaccini
Prof. Valeria Cannillo Acknowledgement I
Acknowledgement
Firstly, I would like to thank Prof. Dr.-Ing. habil. Aldo R. Boccaccini for giving me the opportunity to be part of his research group. I want to express my sincere gratitude for his supervision and support in the past 3 years. Thanks to you, I have been able to develop, both scientifically and personally, in the recent years.
A big thank-you goes to all the members of the Institute of Biomaterials. I would like especially thank Heinz Mahler, Dr. Gerhard Frank, Dr. Julia Will and Bärbel Wust for supporting me in various tasks regarding the organization of labs and other institute duties. I want to particularly thank Dr. Liliana Liverani for all the fruitful discussions and the warm atmosphere in the office. Many thanks go to Dr. Rainer Detsch and Alina Grünewald for their support and suggestions regarding cell culture experiments. I also want to thank all Bachelor- and Master-students, visitors and project students, who I supervised during my time at the Institute of Biomaterials: Serena Miglietta, Theresa Reiter, Inês Almeida Barroso, Pratishtha Mukoo, Moritz Gerlich, Kristin Engel, Lukas Hofer, Swathi Naidu Vakamulla Raghu, Clara Steinert and Nadine Rembold. I hope you could learn something from me, I have definitely learned a lot from all of you. Moreover, I want to thank all my colleagues for supporting me during my time at the institute. A special thanks goes to Dr. Kai Zheng, Dr. Francesca Ciraldo, Dr. Svenja Heise and Dr. Jasmin Hum for the great time in the labs and offices. I also want to thank Vera Bednarzig, for always discussing and solving the challenges at hand, especially during the final stretch of my PhD research. Besides working in an already international group, I was also happy to meet visitors and guest from everywhere of the world. I want to express my gratitude especially to Prof. Josefina Ballare, with whom I had an amazing time in the office and hopefully an amazing time soon in Argentina.
I also want to thank the Johan-Gadolin scholarship program for giving me the opportunity to join the Johan Gadolin Process Chemistry Centre at the Åbo Akademi University in Turku. I would like to especially thank Prof. Leena Hupa for the chance to work in her facilities and to learn about bioactive glass science. I also want to thank Dr. Xiaoju Wang for her support during my visit. Many thanks to Dahiana Andrea Avila Salazar for the great time together, in the lab and during our travels. A big thank-you goes also to the organic chemistry team including Tor Laurén, Peter Backman, Linus Salvander, Luis Bezerra, Mia II Acknowledgement
Mäkinen and Jaana Paananen. Thanks for teaching me how to produce glass and for the great time in Turku!
Without several collaborations, it would not have been possible to write this thesis. Many thanks for Prof. Maciej Sitarz and Dr. Piotr Jelén (Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Cracow) for the support during the structural characterization of my bioactive glasses. I also want to thank Dr. Jochen Schmidt (Institute of Particle Technology, Department of Chemical and Biological Engineering, University of Erlangen) and Zuzana Neščáková (Department of Biomaterials, FunGlass, Alexander Dubček University of Trenčín) for their help with ICP-OES measurements. Special thanks go to Dr. Elisabeth Zinser, Prof. Alexander Steinkasserer and Lena Stich from the department of immune modulation of the Universitätsklinikum Erlangen for the great collaboration. Besides the collaborations connected to my PhD thesis, I was also able to join several collaborations and projects focusing on topics outside my thesis. I am really grateful for the opportunity given to me to work with you. It has certainly broadened the horizon of my research.
In the past three and a half years I have been lucky enough to work with a lot of great people. Even if I did not mention all of them by name (the list would definitely be too long), I would like to thank you all for the great time.
Finally, a big thank-you to my friends, who have always been a great support to me outside of everyday academic life. I owe a very important debt to my partner Alexander, without whom I would never have dared to do a PhD and without whom I would never have finished. My deepest gratitude goes to my family, which is continuously expanding, for their support, trust and love. Abstract III
Abstract
The development of novel, multifunctional wound dressings is essential in the goal of achieving new therapies able to increase the quality of the treatment of patients suffering from insufficient wound healing, especially in the case of chronic and infected wounds. Based on the high social, economic and personal impact of skin wounds, ideal wound dressings able to actively protect the wounded skin and promote the healing process are in demand. Among other properties, modern wound dressings should be easy to produce (cost-effective) and should combine antibacterial with angiogenic properties in addition of being mechanically robust and pliable for easy application. Wound dressings, which can be used to deliver topically drugs or molecules, are of great interest as they can face the challenges leading to an effective wound healing process.
Bioactive glasses (BGs) are a class of non-crystalline, inorganic materials able to release biologically active ions during implantation in the human body, which can induce specific cell responses such as accelerating haemostasis and angiogenesis. Research on BGs has mainly focused on their use in the field of bone replacing devices and bone tissue engineering. However, an increasing amount of studies is reporting promising results on the use of BGs in the regeneration of soft tissues and in wound healing. By designing new BG compositions with the capability to release ions that induce specific biological reactions, the applications of BGs are being further expanded in soft tissue repair. In this context, BGs have shown great potential for wound healing applications especially when BG compositions incorporating biologically active ions have been considered. Therefore, in the first part of this study, copper and/or zinc containing borosilicate and borate BGs were designed for wound healing applications. Copper was chosen to impart angiogenic properties and zinc to introduce antibacterial effects. The fabricated BGs were chemically and physically characterized and they showed excellent dissolution properties which were investigated by dissolution experiments carried out in simulated body fluid, lactic acid and TRIS solutions. Most importantly, especially the series of borate BGs containing copper and/or zinc showed promising results in an antibacterial study and were additionally tested in contact with immune cells. It was found that by using different compositions and concentrations of BGs, immune reactions can be actively (depending on the final application) dampened or induced, IV Abstract which can be achieved by the controlled release of therapeutic ions, in the present case, Zn and Cu.
In order to be useful for applications in wound dressings, BGs must be applied in desired shapes having flexible and pliable characteristics and suitable pore structure. Therefore, in the second part of this thesis, BGs in particulate form (particle size range of 5-20 µm) were combined with methylcellulose (MC) crosslinked with Manuka honey (MH) to achieve flexible and mechanically stable structures. Due to its great biocompatibility and thermosensitive behaviour, MC was chosen, while MH was selected as an innovative cross- linker. Besides providing crosslinking potential, which could be proven in this work by Fourier transform infrared spectroscopy (FTIR), MH was also chosen due to its favourable effects such as promoting epithelialization and antibacterial activity. By using three different fabrication methods, namely freeze drying, electrospinning and 3D printing, different types of wound dressings composed of MC and MH doped with BG particles were fabricated. Freeze dried MC-MH foams containing Cu-doped borate BG particles showed a high porosity (around 95%) and improved wettability. Moreover, the foams showed suitable mechanical properties (compressive strength: ~40 kPa) while acellular in vitro dissolution tests in simulated body fluid indicated the bioactivity of the composite foams. Most importantly, cell biology tests using fibroblasts, keratinocytes and endothelial cells indicated the material biocompatibility and the foams showed superior antibacterial effects against both E. coli and S. aureus bacteria. By additionally combining MC-MH and BG particles with the polymer poly(ɛ-caprolactone) (PCL), which is well-known as a biocompatible and biodegradable polyester, composite fibre mats were fabricated by electrospinning. PCL-MH blend fibre mats containing BG particles (particle size of 5-20 µm) were composed of nanofibers in the range of 100-300 nm and showed a contact angle of around 45°, which is favourable to promote cell attachment due to the hydrophilic character of the surface. Although the ultimate tensile strength of the fibres was slightly reduced by the addition of MH and BG (from around 5 to 3 MPa), the tensile strain was not affected (in the range of 50%). In accordance with the freeze-dried foams, cell biology tests using fibroblasts and keratinocytes indicated the biocompatibility of the electrospun fibre mats. However, tests using E. coli and S. aureus showed no antibacterial efficiency of the fibre mats, indicating the need of further optimization by e.g. increasing the amount of MH and BG. By using phosphate buffer saline as solvent, MC-MH inks containing BG particles (particle size range of 5-20 µm) were successfully produced and used for 3D printing. Mechanical characterization of the 3D printed MC-MH scaffolds showed that the mechanical properties can be tuned by adding different amounts of BG particles into the ink. Moreover, the degradation properties of the resulting printed BG containing MC-MH scaffolds were Abstract V improved in comparison to the too fast dissolving MC-MH scaffolds without the addition of BG. Cell biology tests using fibroblasts further showed the biocompatibility of the MC-MH-BG scaffolds, which indicates that MC-MH-BG inks are innovative for 3D printing and they should be further investigated for their use in the field of biofabrication of personalized wound dressings.
In summary, the use of ion-doped borate and borosilicate BGs is a promising strategy in the development of novel wound dressings that offer several therapeutic effects such as being antibacterial and angiogenic. In this thesis, BG particles were for the first time combined with MH and MC in different types of devices by using different fabrication methods, demonstrating the versatile nature of these material combinations. The successful application of MH as natural crosslinker in combination with ion-doped BGs, enable the fabrication of different wound dressings offering various advantages. Whereas freeze dried wound dressings are especially interesting in the treatment of infected, deep wounds, electrospun fibre mats are important for superficial wounds. Inks based on MC-MH in combination with BGs demonstrated an attractive biomaterial platform for 3D printing as they open a new field for further research focusing on biofabrication. Overall, the results obtained in this thesis showed the great potential of ion-doped BGs in combination with MC and MH for the fabrication of antibiotic-free wound dressings capable of preventing infections and with extra functionality to promote wound regeneration.
VI Abstract
Kurzzusammenfassung VII
Kurzzusammenfassung
Die Entwicklung neuartiger, multifunktionaler Wundauflagen ist entscheidend um neue Therapien zu entwickeln, mit denen die Qualität der Behandlung von Patienten mit unzureichender Wundheilung verbessert werden kann, insbesondere bei chronischen und infizierten Wunden. Ideale Wundauflagen sind aufgrund der hohen sozialen, wirtschaftlichen und persönlichen Auswirkungen von Hautwunden gefragt, die die verletzte Haut aktiv schützen und den Heilungsprozess fördern können. Moderne Wundauflagen sollten unter anderem einfach herzustellen (kostengünstig) sein und antibakterielle mit angiogenen Eigenschaften kombinieren sowie mechanisch robust und biegsam für eine einfache Anwendung sein. Wundauflagen, mit denen topisch Medikamente oder Moleküle abgegeben werden können, sind von großem Interesse, da sie sich den Herausforderungen stellen können, die zu einem wirksamen Wundheilungsprozess führen.
Bioaktive Gläser sind eine Klasse nichtkristalliner, anorganischer Materialien, die in der Lage sind, biologisch aktive Ionen während der Implantation in den menschlichen Körper wieder freizusetzen, was spezifische Zellreaktionen wie die Beschleunigung der Hämostase und Angiogenese auslösen kann. Die Forschung zu bioaktiven Gläsern konzentrierte sich hauptsächlich auf die Verwendung von bioaktiven Gläsern im Bereich des Knochenersatzes und im Knochen Tissue Engineering. Eine zunehmende Anzahl von Studien hat jedoch vielversprechende Ergebnisse zur Verwendung von bioaktiven Gläsern bei der Regeneration von Weichgeweben und bei der Wundheilung berichtet. Durch die Entwicklung neuer bioaktive Glas-Zusammensetzungen, die Ionen enthalten, welche spezifische biologische Reaktionen auslösen können, werden die Anwendungen von bioaktiven Gläsern bei der Reparatur von Weichgewebe weiter ausgebaut. In diesem Zusammenhang haben bioaktive Gläser ein großes Potenzial für Wundheilungsanwendungen gezeigt, insbesondere wenn bioaktive Glas-Zusammensetzungen mit biologisch aktiven Ionen in Betracht gezogen wurden. Daher wurden im ersten Teil dieser Studie kupfer- und/oder zinkhaltige Borosilikat- und Borat-bioaktive Gläser für Wundheilungsanwendungen entwickelt. Kupfer wurde aufgrund seiner angiogenen Eigenschaften ausgewählt und Zink, um antibakterielle Wirkungen hinzuzufügen. Die hergestellten bioaktiven Gläser wurden chemisch und physikalisch charakterisiert und zeigten ausgezeichnete Auflösungseigenschaften, die durch verschiedene Auflösungsexperimente in simulierter Körperflüssigkeits-, Milchsäure- und TRIS-Lösung VIII Kurzzusammenfassung untersucht wurden. Insbesondere die Serie von Kupfer- und / oder Zink enthaltenden Borat- bioaktiven Gläsern haben in einer antibakteriellen Studie vielversprechende Ergebnisse gezeigt und wurden zusätzlich in Kontakt mit Immunzellen getestet. Es wurde festgestellt, dass durch Verwendung unterschiedlicher Zusammensetzungen und Konzentrationen von bioaktiven Gläsern Immunreaktionen abhängig von der endgültigen Anwendung aktiv gedämpft oder induziert werden können.
Um für Anwendungen in Wundauflagen nützlich zu sein, müssen bioaktive Gläser in gewünschten Formen mit flexiblen und biegsamen Eigenschaften und einer geeigneten Porenstruktur verfügbar sein. Daher wurden im zweiten Teil dieser Arbeit bioaktive Gläser in Partikelform (Partikelgrößenbereich von 5 bis 20 um) mit Methylcellulose (MC) kombiniert, die mit Manuka-Honig (MH) vernetzt wurden, um flexible und mechanisch stabile Strukturen zu erhalten. Aufgrund seiner Biokompatibilität und seines wärmeempfindlichen Verhaltens wurde MC ausgewählt, während MH als innovativer Vernetzer ausgewählt wurde. Neben der Bereitstellung eines Vernetzungspotenzials, das in dieser Arbeit durch Fourier- Transformations-Infrarotspektroskopie (FTIR) nachgewiesen werden konnte, wurde MH auch aufgrund seiner günstigen Wirkungen wie der Förderung der Epithelialisierung und der antibakteriellen Aktivität ausgewählt. Unter Verwendung von drei verschiedenen Herstellungsverfahren, nämlich Gefriertrocknung, Elektrospinnen und 3D-Druck, wurden verschiedene Arten von Wundauflagen aus MC und MH hergestellt, die mit bioaktiven Glas- Partikeln dotiert wurden. Gefriergetrocknete MC-MH-Schäume, die Cu-dotierte Borat- bioaktive Glas-Partikel enthielten, zeigten eine hohe Porosität (etwa 95%) und eine verbesserte Benetzbarkeit. Darüber hinaus zeigten die Schäume geeignete mechanische Eigenschaften (Druckfestigkeit von etwa 40 kPa), während azelluläre in-vitro-Auflösungstests in simulierter Körperflüssigkeit die Bioaktivität der Komposit-Schäume zeigten. Am wichtigsten ist, dass zellbiologische Tests unter Verwendung von Fibroblasten, Keratinozyten und Endothelzellen die Material-Biokompatibilität und die Schäume überlegene antibakterielle Wirkungen gegen E. coli und S. aureus Bakterien zeigten. Durch zusätzliches Kombinieren von MC-MH- und bioaktiven Glas-Partikeln mit dem Polymer Polycaprolactone (PCL), das als biokompatibler und biologisch abbaubarer Polyester bekannt ist, wurden Verbundfasermatten durch Elektrospinnen hergestellt. Die Fasermatten auf der Basis von PCL und MC, die MH- und bioaktive Glas-Partikel enthielten, bestanden aus Nanofasern im Bereich von 100 bis 300 nm und zeigten einen Kontaktwinkel von etwa 45 °, was auf eine gute Zellanhaftung aufgrund des hydrophilen Charakters der Oberfläche hinweist. Obwohl die Zugfestigkeit der Fasern durch Zugabe von MH und bioaktiven Glas-Partikeln (von etwa 5 bis 3 MPa) leicht verringert wurde, wurde die Zugspannung nicht beeinflusst (im Bereich von Kurzzusammenfassung IX
50%). In Übereinstimmung mit den gefriergetrockneten Schäumen zeigten zellbiologische Tests unter Verwendung von Fibroblasten und Keratinozyten die Biokompatibilität der elektrogesponnenen Fasermatten. Tests unter Verwendung der Bakterien E. coli und S. aureus zeigten jedoch keine antibakterielle Wirksamkeit der Fasermatten, was auf die Notwendigkeit einer weiteren Optimierung durch z.B. erhöhen der Menge an MH und bioaktiven Glas erreicht werden kann. Unter Verwendung von Phosphatpuffer-Salzlösung als Lösungsmittel wurden MC-MH-Drucktinten, die bioaktive Glas-Partikel enthielten (Partikelgrößenbereich von 5 bis 20 um), erfolgreich hergestellt und für den 3D-Druck verwendet. Die mechanische Charakterisierung der 3D-gedruckten MC-MH-Gerüste zeigte, dass die mechanischen Eigenschaften durch Zugabe unterschiedlicher Mengen an bioaktiven Glas-Partikeln in die Tinte eingestellt werden können. Darüber hinaus wurden die Abbaueigenschaften des resultierenden gedruckten MC-MH-Gerüstes, welches bioaktives Glas enthielt, im Vergleich zu den zu schnell auflösenden MC-MH-Gerüsten ohne Zusatz von bioaktivem Glas verbessert. Zellbiologische Tests unter Verwendung von Fibroblasten zeigten ferner die Biokompatibilität der MC-MH-bioaktives Glas-Gerüste, was darauf hinweist, dass MC-MH-bioaktives Glas-Tinten für den 3D-Druck innovativ sind und für ihre Verwendung im Bereich der Biofabrikation von personalisierter Wundauflagen weiter untersucht werden sollten.
Zusammenfassend ist die Verwendung von ionendotierten Borat- und Borosilikat- bioaktiven Gläsern eine vielversprechende Strategie bei der Entwicklung neuartiger Wundauflagen, die verschiedene therapeutische Wirkungen bieten, z. B. antibakteriell und angiogen. In dieser Arbeit wurden bioaktive Glas-Partikel zum ersten Mal mit MH und MC in verschiedenen Arten von Vorrichtungen kombiniert, indem verschiedene Herstellungsmethoden verwendet wurden, um die Vielseitigkeit dieser Materialkombinationen zu demonstrieren. Die erfolgreiche Anwendung von MH als natürlicher Vernetzer in Kombination mit ionendotierten bioaktiven Gläsern ermöglicht die Herstellung verschiedener Wundauflagen mit verschiedenen Vorteilen. Während gefriergetrocknete Wundauflagen bei der Behandlung infizierter, tiefer Wunden besonders interessant sind, sind elektrogesponnene Fasermatten für oberflächliche Wunden geeignet. Drucktinten auf Basis von MC-MH in Kombination mit bioaktiven Gläsern zeigten großartige Ergebnisse und eröffnen ein neues Feld für weitere Forschungen mit Schwerpunkt auf der Biofabrikation. Insgesamt zeigten die in dieser Arbeit erhaltenen Ergebnisse das große Potenzial von ionendotierten bioaktiven Gläsern in Kombination mit, mit MH vernetztem, MC, für die Herstellung von antibiotikafreien Wundauflagen, die Infektionen verhindern können und über zusätzliche Funktionen zur Förderung der Wundregeneration verfügen.
X Kurzzusammenfassung
Table of Contents
Table of Contents
Acknowledgement ...... I
Abstract ...... III
Kurzzusammenfassung ...... VII
1 Introduction...... 1 1.1 Motivation: Demand for soft tissue regeneration techniques ...... 1 1.2 Aim and objectives of the work ...... 2
2 Fundamentals ...... 5 2.1 Wound healing ...... 5 2.1.1 Skin...... 5 2.1.2 Wounds and wound healing process ...... 6 2.1.3 Treatment of wounds ...... 8 2.2 Bioactive glass ...... 10 2.2.1 Structure of bioactive glasses ...... 10 2.2.2 Mechanisms of bioactivity ...... 13 2.2.3 Ion doping ...... 17 2.2.4 Applications ...... 18 2.3 Wound dressing fabrication methods ...... 20 2.4 Cellulose ...... 23 2.4.1 Methylcellulose ...... 24 2.4.2 Crosslinking methods ...... 25 2.5 Manuka honey ...... 26
3 Borate and borosilicate glasses doped with copper and/or zinc ...... 29 3.1 Introduction ...... 29 3.2 Experimental Procedure ...... 29 3.2.1 Glass melting and glass powder preparation ...... 29 3.2.2 Characterization methods ...... 30 3.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX) ...... 30 3.2.2.2 Nuclear Magnetic Resonance (NMR) ...... 30 3.2.2.3 X-ray diffraction (XRD) ...... 31 3.2.2.4 Differential thermal analysis (DTA) ...... 31 3.2.2.5 Hot stage microscopy (HSM) ...... 31 3.2.2.6 FTIR ...... 31 3.2.2.7 Static in vitro dissolution ...... 32 3.2.2.8 Continuous in vitro dissolution ...... 32 3.2.3 Antibacterial test...... 33 3.2.4 Cell culture ...... 33 3.3 Results ...... 35 3.3.1 Structural characterization of the fabricated BGs ...... 35 3.3.2 Thermal characterization of the fabricated BGs ...... 40 3.3.3 Acellular dissolution behavior of BGs in different solutions ...... 42 3.3.3.1 Dissolution in TRIS under static conditions ...... 42 3.3.3.2 Dissolution in SBF under static and dynamic conditions...... 45 3.3.3.3 Dissolution in lactic acid under static conditions ...... 50 Table of Contents
3.3.4 Antibacterial efficiency of the BGs ...... 53 3.3.5 Dose-depending effect of BGs on dendritic cells ...... 54 3.4 Discussion ...... 60 3.4.1 Influence of network former on BG characteristics ...... 60 3.4.2 Influence of incorporation of copper/zinc on BG characteristics ...... 63
4 Fabrication of bioactive glass containing wound dressings by freeze drying ...... 67 4.1 Introduction ...... 67 4.2 Experimental procedure ...... 67 4.2.1 Production of foams ...... 67 4.2.2 Characterization ...... 68 4.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX) ...... 68 4.2.2.2 Porosity ...... 68 4.2.2.3 Fourier transform infrared spectroscopy FTIR ...... 68 4.2.2.4 Mechanical testing ...... 69 4.2.2.5 Contact angle measurement ...... 69 4.2.2.6 X-ray diffraction (XRD) analysis...... 69 4.2.3 In vitro dissolution in SBF ...... 70 4.2.4 Antibacterial efficiency of the foams ...... 70 4.2.5 In vitro cell tests ...... 71 4.2.5.1 Indirect cytotoxic test using mouse embryotic fibroblast (MEF) cells ...... 71 4.2.5.2 Direct cell test using human dermal fibroblast (hDF) ...... 71 4.2.5.3 In vitro scratch test using mouse embryotic fibroblast (MEF) and human keratinocytes-like (Hacat) cells72 4.2.5.4 Co-culture of human dermal fibroblasts (hDFs) and human umbilical vein endothelial cells (HUVECs) 72 4.3 Results ...... 73 4.3.1 Crosslinking efficiency of MC with MH ...... 73 4.3.2 Characterization of the foams ...... 74 4.3.3 Acellular bioactivity using SBF ...... 78 4.3.4 Antibacterial efficiency ...... 81 4.3.5 Compatibility with different skin cells ...... 82 4.3.5.1 Conditioned CCM containing dissolution products used for cell experiments...... 82 4.3.5.2 Indirect cell tests using MEF cells ...... 84 4.3.5.3 Direct cell test using human dermal fibroblast (hDF) ...... 86 4.3.5.4 In vitro scratch test using mouse embryotic fibroblast (MEF) cells and human keratinocytes- like (Hacat) cells ...... 88 4.3.5.5 Co-culture of human dermal fibroblasts (hDFs) and human umbilical vein endothelial cells (HUVECs) 89 4.4 Discussion ...... 92 4.4.1 Practical properties ...... 92 4.4.2 Physical properties ...... 93 4.4.3 Biological properties ...... 94
5 Fabrication of bioactive glass containing wound dressings by electrospinning ...... 97 5.1 Introduction ...... 97 5.2 Experimental procedure ...... 97 5.2.1 Production of electrospun fibre mats ...... 97 5.2.2 Characterization ...... 98 5.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX) ...... 98 5.2.2.2 Fourier transform infrared spectroscopy FTIR ...... 98 5.2.2.3 Mechanical testing ...... 98 5.2.2.4 Contact angle measurement ...... 99 Table of Contents
5.2.2.5 X-ray diffraction (XRD) measurement ...... 99 5.2.3 In vitro dissolution in SBF ...... 99 5.2.4 Antibacterial efficiency of the fibre mats ...... 99 5.2.5 In vitro cell tests ...... 99 5.2.5.1 Direct cell tests using human dermal fibroblasts (hDFs) ...... 100 5.2.5.2 In vitro scratch test using human dermal fibroblast (MEF) and human keratinocytes-like (Hacat) cells100 5.3 Results ...... 101 5.3.1 Production and characterization of MC-MH-B3 fibre mats ...... 101 5.3.2 Characterization of optimized fibre mats ...... 102 5.3.3 Acellular bioactivity using SBF ...... 105 5.3.4 Antibacterial efficiency ...... 109 5.3.5 Compatibility with different skin cells ...... 110 5.3.5.1 Conditioned CCM containing ionic dissolution products ...... 110 5.3.5.2 Direct cell tests using human dermal fibroblasts (hDFs) ...... 111 5.3.5.3 In vitro scratch test (migration test) using human keratinocytes-like (Hacat) cells...... 113 5.4 Discussion ...... 114 5.4.1 Practical properties ...... 114 5.4.2 Physical properties ...... 115 5.4.3 Biological properties ...... 117
6 Fabrication of bioactive glass containing wound dressings by 3D printing ...... 119 6.1 Introduction ...... 119 6.2 Experimental procedure ...... 119 6.2.1 Scaffold production ...... 119 6.2.2 Characterization ...... 121 6.2.2.1 Optical microscope images ...... 121 6.2.2.2 Fourier transform infrared spectroscopy (FTIR) ...... 121 6.2.2.3 Mechanical testing ...... 121 6.2.2.4 Contact angle measurement ...... 121 6.2.3 Swelling behaviour and change of weight of the 3D printed scaffolds ...... 121 6.2.3.1 Swelling and change of weight under “dry” conditions in DW ...... 122 6.2.3.2 Swelling and change of weight under “wet” conditions ...... 122 6.2.4 In vitro direct cell tests using hDFs ...... 122 6.3 Results ...... 123 6.3.1 Optimization of printing ink and printing parameters ...... 123 6.3.2 Characterization of optimized 3D printed scaffolds ...... 125 6.3.3 Swelling behaviour and change of weight of 3D printed scaffolds under different conditions ...... 130 6.3.3.1 Swelling and change of weight under wet conditions ...... 131 6.3.3.2 Swelling and change of weight under dry conditions ...... 132 6.3.4 Preliminary cell compatibility of 3D printed scaffolds ...... 133 6.4 Discussion ...... 134 6.4.1 MC as printing material ...... 134 6.4.2 Effect of addition of MH and B3 BG ...... 136 6.4.3 Suitability of MC-MH based hydrogels containing B3 BG as bioink for biofabrication ..... 138
7 Summary and Outlook ...... 139 7.1 Concluding remarks ...... 139 7.2 Future directions ...... 144
Table of Contents
1 Introduction 1
1 Introduction
1.1 Motivation: Demand for soft tissue regeneration techniques
The main task of the largest organ of the human body, the skin, is to act as biological, physical and chemical barrier for the underlying organs [1]. The skin is also the organ most exposed to wounds in form of injury, scratches and burns [2], [3]. Undoubtedly, the proper function of the skin is vital for a healthy life and, in turn, dysfunctional wound healing results in incremental pain, immobility and discomfort for the patient as well as to a socioeconomic burden worldwide [1], [4]. It is estimated that around 2-4% of the total health budgets worldwide is spent for wound treatments, mainly in the form of wound dressing changes, nursing time, hospitalization and related infections [5]. Furthermore, approximately around 2% of the world population will suffer a chronic wound in their lives [6], [7]. Under normal circumstances, healing of a wound is a timely and orderly process, resulting in a restoration of the integrity of the skin (acute wounds). Chronic wounds in contrast are defined by the failing of the wound healing process, in most cases due to diseases (e.g. diabetes) or bacterial infections [8], [9]. Based on the increasing number of patients suffering from diabetes, on the appearance of multi-resistant bacteria and on the aging of the world population, it is estimated that the number of people suffering chronic wounds and the related costs will further increase [7], [10].
Considering the highly economic, social and personal impact that skin wounds have on individuals and on national systems, there is a high interest in the development and fabrication of novel wound dressings. More than 3000 types of wound dressings are up to date available on the market and can be used to treat different types of wounds to achieve faster healing [11]. Whereas originally wound dressings were used to just passively cover and protect the wound [12], nowadays wound dressings are developed to actively protect and additionally promote the healing process [13]. In order to treat a chronic wound, these dressing have to address several issues, which might occur during the (impaired) wound regeneration process. For instance, ideal wound dressings should promote haemostasis in order to diminish blood loss, which could lead to impaired wound healing [14], [15]. In addition, a disturbed process of angiogenesis (the formation of new blood vessels) is one of 2 1 Introduction the main reasons for the insufficient healing of acute wounds, leading to chronic wounds [16]. Blood vessels supply the wound with oxygen and nutrients which are needed during the regeneration process [4]. Therefore, an ideal wound dressing should promote the angiogenesis process (e.g. by release of the growth factors) [1]. Another common reason for chronic wounds is the appearance of infections [16], [17]. Especially large wounds can open the way for pathogenic bacteria, where they can grow and multiply, subsequently entering the human body [18]–[20]. After reaching a critical level of bacteria, the wound healing process will be disturbed leading to a chronic and painful wound [21]. Infected wounds are traditionally treated by using antibiotics [3]. However, due the decreasing number of new developed antibiotics and rapid spread of antibiotic resistance bacteria, alternative (antibiotic-free) antibacterial treatments must be developed [22].
1.2 Aim and objectives of the work
In the last few decades, numerous biomaterials have been increasingly developed in order to improve the wound healing process by positively interacting with the biological system [23]. Within the field of biomaterials, bioactive glasses (BGs) are an example of bioreactive and biodegradable biomaterials, which are able to release biologically active ions during degradation leading to a specific cell response [24], [25]. Although BGs are well-known for their use in bone regeneration, an increasing amount of research is focusing on the use of BGs in soft tissue repair and wound healing [26], [27]. Besides being able to degrade and release ions, BGs offer the great advantage of being easily doped with special therapeutic active ions which can be locally released to achieve specific biological effects [28]. In this research project, based on the well-known 13-93 BG composition, new BGs containing additionally copper and zinc were developed to introduce angiogenic (copper) and antibacterial properties (copper and zinc) [29]–[31]. In terms of infections, ion-doped BGs offer a great advantage: at least according to the current knowledge, bacteria are not able to develop resistance against the antibacterial ions released by BG, which shows the great potential of BGs in comparison to antibiotics [32], [33]. Therefore, the first research question, on which this thesis focuses on, is:
Are the developed ion-doped BGs suitable for applications in wound healing?
A suitable modern wound dressing should fulfil several requirements, including being easy to handle, sterilizable and cost-effective [34], [35]. Additionally, modern wound dressings should be ideally able to keep a moist environment, allow the migration and 1 Introduction 3 proliferation of relevant skin cells and be antibacterial [9], [12], [36]. In order to fulfil these requirements, BGs have to be fabricated into desired shapes and showing flexible characteristics by using a suitable fabrication method. Convenient fabrication techniques in this regard are freeze drying, electrospinning and 3D printing. Moreover, BGs (in particulate form) have to be combined with a suitable polymer to achieve a flexible (bendable) structure. Therefore, within this thesis, two more research questions were handled:
Can suitable wound dressings be fabricated by combining BGs with methyl cellulose (MC) and Manuka honey (MH)?
and
Are freeze drying, electrospinning and 3D printing suitable fabrication methods to produce modern wound dressings incorporating BG in particle form?
MC was selected as “biopolymer partner” of BGs due to its biocompatibility and thermosensitive behaviour, while MH was selected based on its favourable antibacterial and anti-inflammatory properties. A graphical summary of the thesis is given in Figure 1, providing information about the materials used and the fabrication methods employed. In order to face the different research questions, the research was divided in different tasks, which are presented in the chapters of the thesis. Chapter 2 covers the Fundaments including the state-of-the-art in the field of wound healing, BGs and fabrication methods for wound dressings. Chapter 3 focuses on the fabrication as well as the basic characterization of the ion- doped BGs investigated in this project. A special focus was given to the dissolution of the BGs and the resulting biological properties using immune cells. Further, three fabrication methods were investigated; namely freeze drying (Chapter 4), electrospinning (Chapter 5) and 3D printing (Chapter 6). Wound dressings fabricated by these techniques were chemically, physically and biologically evaluated. Based on the results obtained in this thesis, the last chapter “Final conclusion and future directions” (Chapter 7) discusses the findings of the different experimental tasks carried out and future research directions in the field of biomaterials for wound healing applications are suggested inspired by the key results achieved in this project.
4 1 Introduction
Figure 1: Graphical abstract summarizing the work carried out in this project 2 Fundamentals 5
2 Fundamentals
2.1 Wound healing
2.1.1 Skin
Skin (Figure 2), the largest organ of the human body, comprises about 8% of the total human body mass [9]. It covers the entire external body surface area and has a total area of around 2 m2 [9], [37]. The main function of the skin is to protect the body against chemical, mechanical, thermal and osmotic damage as well as against microbial invasion [34]. Moreover, the skin is responsible for the synthesis of vitamin D3, has to prevent water loss and is involved in the immunological surveillance [2], [38]. In order to fulfil all these tasks, the skin needs to be at the same time tough, robust and flexible and additionally needs to enable the formation of a self-renewing and self-repairing interface between the internal body and its environment [9], [39].
Figure 2: Cross-section of human skin showing the different layers: epidermis, dermis and subcutaneous tissue (inspired by refs [9], [40], [41]).
According to Figure 2, the human skin can be divided in three different layers: the epidermis, dermis and subcutaneous tissue layer [9], [40], [41]. The first layer, the epidermis, 6 2 Fundamentals is composed of mainly (up to 90%) keratinocytes and additionally of other cell types e.g. merkel cells, melanocytes and Langerhans cells [40]. Therefore, this layer is totally cellular, but although it is quite thin it has a sufficient thickness to work as a vital barrier [9], [34]. The epidermis is known to prevent the entry of microorganisms, to maintain body hydration and, by constant recycling of the basal cell layer, to maintain homeostasis [9], [34], [40]. The second layer, the dermis, is around 1-4 cm thick and is composed of collagen, elastin, hair roots, nerve cells, blood vessels, lymphatic vessels, sweat glands, glycosaminoglycan and mesenchymal stem cells [34], [40], [42]. The majority of cells found in the “real skin layer” are fibroblasts, which are able to produce remodelling enzymes (e.g. proteases, collagenases) [9], [34]. The main function of fibroblasts is to provide a suitable structural toughness to the skin and, at the same, a suitable flexibility in order to support the extensive lymphatic system, vasculature and nerve bundles [9], [40]. The last layer is composed of mainly fat and a layer of loose connective tissue [34]. The subcutaneous layer functions as a support and anchor for the dermal and epidermal layers and is well vascularized [9], [34], [40]. In this layer, adipocytes, blood vessels, macrophages, nerves and fibroblasts can be found [40].
2.1.2 Wounds and wound healing process
The skin is the organ in the human body which is most exposed to injury, scratches and burns due to external physical, mechanical, biological and chemical agents [2], [3]. Moreover, the functions of the skin can be impaired due to illness, e.g. diabetes, cancer [2]. A wound is in generally defined as a “disturbance in the normal skin anatomy and function” [8] resulting in the “loss of continuity of epithelium with or without the loss of underlying connective tissue” [8] due to a mechanical trauma, reduced blood circulation, burns, surgical procedure and/or aging [36]. This damage or even loss of the function and integrity of the skin can lead to significant disability or even to death [36]. Depending on the depth, wounds of the skin can be classified into different categories: superficial (only the epidermis is affected), partial thickness (the epidermis and dermis are affected) and full-thickness/deep (also subcutaneous layer and sometimes even bone are affected) wounds [9], [43]. In any case, the wound healing process starts immediately to restore the function of the skin. This healing process can be divided into different phases, as schematically shown in Figure 3. 2 Fundamentals 7
Figure 3: The different phases of wound healing. Overlap occurs between the different phases. The beginning and ending of each phase are approximate (inspired by refs [44], [45]).
Immediately after the injury occurs, the first wound healing phase starts: the homeostasis phase [8]. During this phase, several processes take place. First, bleeding occurs in order to ensure the removal of toxic waste followed by blood clotting [9]. During blood clotting, among others fibrin, fibronectin, etc. supply a scaffold-like matrix, which ensures the migration of relevant cells (fibroblasts, keratinocytes, endothelial cells, etc.) and causes the accumulation of relevant growth factors [40], [43]. This phase is normally completed within a few hours [8]. Subsequently, after the first phase is completed, the inflammatory phase starts. In the early phase neutrophils arrive and produce interleukin-1 (IL-1), IL-6 and tumor necrosis factor (TNF-alpha). These proteins activate the inflammatory response. Moreover, neutrophils stimulate the secretion of vascular endothelial growth factor (VEGF) and IL-8 in order to repair blood vessels. In the late phase of the inflammatory phase, monocytes differentiate into macrophages, which also produce several growth factors to stimulate the expansion and migration of relevant cells [40]. The inflammatory phase is normally completed within 24-72 hours after the injury occurs, but could also take up to 7 days [8], [40]. However, the adaptive immune system is active at least for 4 days involving T-cells and B-cells [40].
As shown in Figure 3, the proliferation phase starts before the inflammatory phase ends, in normal conditions 3 days after the injury occurs [44]. The proliferation phase can be further subdivided into three steps: re-epithelialization followed by angiogenesis and then the formation of granulation tissue. During the first step, cytokines and growth factors (e.g. nerve growth factor, epidermal growth factor, insulin-like growth factor and keratinocyte growth factors) activate the re-epithelialization, which in turn causes the expansion of epithelial cells, fibroblasts, keratinocytes and stem cells. As the name is suggesting, during the angiogenesis 8 2 Fundamentals phase, growth factors (e.g. fibroblast growth factor beta (FGF-beta), platelet derived growth factor (PDGF), insulin and VEGF) activate the growth of endothelial cells, leading to the formation of new blood vessels. In the last phase of the proliferation phase, macrophages, granulocytes and mainly fibroblasts are involved. Fibroblasts produce collagen as well as other important extracellular matrix (ECM) molecules. The ECM is working as a scaffold structure for cell adhesion and also organizes the growth and differentiation of relevant cells (including fibroblasts). At the end of the proliferation phase, fibroblasts either differentiate into myofibroblasts, which form a scar, or undergo apoptosis [40].
In the last phase of the wound healing process, the remodelling phase (Figure 3), more collagen and elastin are produced by the wound scar. Fibroblasts mature still into myofibroblasts and keratinocytes also undergo apoptosis. More importantly, inflammatory cells (e.g. T-cells and macrophages) play a crucial role in the ending of the immune response as well as in the ending of the wound healing process [40].
If the wound healing process occurs timely and orderly, it results in a sustained restoration of the functional and anatomic integrity of the skin. They are named acute wounds [8], [9]. In case that the regular wound healing fails, an acute wound turns into a chronic wound [8]. Possible reasons for the disturbance of the wound healing process could be specific diseases, such as diabetes and tumours [9]. Moreover, contaminations could have negative effects on the ECM, the release of growth factors and the granulation tissue, and are therefore the main reason for the development of chronic wounds [2], [3]. In this case, the wound healing process takes a long period of time, often exceeding 12 weeks and reoccurrence is also not uncommon [9]. To avoid chronic wounds and also to treat acute/chronic wounds, the main objective of research focusing on the regeneration of skin is to design tools to support the human body during the wound healing process in order to heal a wound as fast as possible with a minimum of pain, scarring and discomfort for the patient [9].
2.1.3 Treatment of wounds
Treatment of wounds normally starts with the cleaning and disinfecting of the wound. Depending on the size, a wound can be then closed by a wound dressing [12]. Wounds, which are larger than 1cm in diameter, mostly full-thickness wounds (see section 2.1.2), need a wound dressing in order to avoid extensive scar formation [35]. In former times, wound dressings like gaze, bandages, lint, cotton wool and plasters were applied to just cover and 2 Fundamentals 9 protect the wound [12], lacking in supporting the wound healing process. Therefore, the current gold standard is autologous skin transplantations, which could be obtained by harvesting (using a dermatome) a skin layer including the epidermis and a thin layer of the dermis. Then, the harvested skin layer can be applied on a (full-thickness) wound and heals normally within one week. However, donor sites for harvesting the skin layer are extremely limited [35]. Another possibility would be to use a skin layer harvested from a cadaveric skin (allografts). Both, autografts and allografts are offering ethical and safety issues as the possibility of agent transmission and rejection [35], [43]. Moreover, problems including contraction, high costs, scarring and delayed vascularization are associated with the current gold standard [34], [36].
Therefore, current research is focusing on the development of new wound dressings fulfilling a range of requirements in order to support the repair and regeneration process of the skin more effectively. Besides being easy to handle and apply to the wound, covering and protecting the wound bed and being sterile (or sterilizable), wound dressings need to be cost effective [34], [35]. In order to create a perfect environment for wound healing, a wound dressing needs to keep a moist environment to allow migration and proliferation of cells, provide a proper oxygenation, reduce water/blood loss and perform a thermal insulation of the wound [9], [12], [36]. Moreover, the wound dressing should allow the removal of waste and debris [12]. In terms of chronic wounds, a huge advantage of an ideal wound dressing would be to work as an barrier for infection and bacterial invasion [9], [36]. Finally, a wound dressing should be readily available and cost-effective [35].
Tissue engineering products are an alternative to the current gold standard, which may fulfil all the requirements. Wound dressings, based on the concept of tissue engineering, are commonly biodegradable scaffolds from synthetic and/or natural origin fabricated by various techniques. On top of these scaffolds, skin cells (autologous or allogenic origin) are grown in vitro for a few weeks and then, these dressings (with cells on top) are transplanted into the wounds [43]. These modern dressings can support the wound healing process by different ways, as by modifying the physical parameters, by giving an extracellular support for the cells or by delivering bioactive molecules. These molecules can in turn improve the wound healing process among others by positively influencing skin cells, by promoting angiogenesis and/or by being antibacterial [12]. Research focusing on the development of wound dressings based on the tissue engineering concept has to accomplish two main tasks: the generation of suitable biodegradable materials (e.g. polymers as main matrix, ceramics as fillers) and then to fabricate a suitable wound dressing out of these materials [34]. 10 2 Fundamentals
2.2 Bioactive glass
In order to face the first challenge mentioned above, namely the generation of suitable biodegradable materials, bioactive glasses (BGs) were chosen in this project. Since their invention in 1969 by Prof. Larry Hench [25], [46]–[48], BGs have gained a lot of attention in the field of orthopaedic, dental applications and bone tissue engineering [49]–[53]. The original idea of Prof. Hench was to develop a material that could bond to bone [46]. In contrast to the common inert materials used in orthopaedic applications (e.g. stainless steel), BGs interact with human tissues resulting in a favourable reaction of the human body in the presence of BGs [54]. In order to understand the reaction occurring, basic knowledge about the BGs structure and the resulting properties is necessary.
2.2.1 Structure of bioactive glasses
Glasses in general have an amorphous structure without a long-range order (in contrast to crystalline ceramics). Moreover, a glass transition range, in which a system changes from a super cooled liquid into a solid glass structure, is also characteristic for glasses. Due to the ability of glasses to smoothly decrease their viscosity with increasing temperature, different shapes and forms can be easily fabricated [54]. In general, glasses are composed of three different components, the network formers, network modifiers and intermediate oxides, as shown schematically in Figure 4.
Figure 4: Typical glass structure including the three components: network formers, network modifiers and intermediate oxygens (including bridging and non-bridging oxygens) [47], [54]–[56] 2 Fundamentals 11
Network formers are able to form a glass structure without the need of additional components. Typical network formers for BGs are silica, boron trioxide and phosphorus pentoxide or a mixture of these [54]. By introducing network modifiers, the glass structure can be altered. Network modifiers turn bridging oxygens atoms, which are mainly characterized by covalent bindings, into non-bridging oxygen atoms, which are mainly characterized by ionic bindings. Typical examples for network modifiers are oxides of alkali or alkaline-earth metals (e.g. Ca, K, Na, Mg) having a high coordination number [57]–[59]. As already mentioned, the intermediate oxides are present in the form of bridging and non- bridging oxygens [60], [61].
A) B)
Figure 5: Basic building units of bioactive glasses based on A) silica and B) boron trioxide (inspired by refs [54], [62])
The glass chemical composition, especially in the case of BG, determines the properties of the BG, which in turn determines the performance of the BGs and BGs containing implants in the human body. In order to design a suitable BG, first the right network former needs to be chosen. Originally, BGs were based on silica, as the first composition developed by Prof. Hench in 1969 [25], [47]. The so-called 45S5 BG composition is composed of (in wt.%) 45 SiO2-24.5
Na2O-24.5 CaO-6 P2O5 [25], [47]. As shown in Figure 5 A, this silicate glass, as well as every
BGs based on silica, is built by SiO4 tetrahedrons, which can connect with each other through the above described bridging atoms. This tetrahedrons, also described as Q4 units (4 describes that the tetrahedron is connected to four bridging oxygen atoms), turn into Q3, Q2 and Q1 units due to the addition of network modifiers. In contrast to conventional glasses, which are mainly built of Q4 and Q3 units, BGs are showing a much disrupted structure composed of mainly Q2 units [54], [59], [63]. The proportion of the different Qn units inside a glass matrix can be also described by the network connectivity. The network connectivity can be calculated based on the amount of network modifiers inside a glass composition, which gives an idea about the number of bridging atoms per network former. In both cases, based on calculations or measurement of Qn units, the network connectivity can be used to give an assumption about the polymerization degree of the glass network and in turn about the glass properties 12 2 Fundamentals
(e.g. bioactivity, crystallization tendency, glass transition temperature). BGs normally have a network connectivity between 2-3 [64]–[67].
Compared to silicate BGs, BGs based on boron trioxide offer a lower chemical durability leading to higher dissolution rates, which is more favourable for wound healing applications. Therefore, in this work BGs were fabricated based on silica and boron trioxide
(or a mixture of both). In contrast to SiO4 tetrahedrons, borate BGs are formed by BO3 trihedrons or chains of BO3 triangles, whereas due to the threefold coordination number, borate BGs cannot form a full three dimensional network (leading to the low chemical durability) (Figure 5B) [68], [69]. The addition of network modifiers does not (in contrast to silicate BGs) result in a transformation of bridging to non-bridging oxygen atoms. To a certain extent, the addition leads rather to a formation of BO-4 tetrahedrons balanced with M+. By exceeding a certain amount of modifiers, the BO-4 groups change back to negatively charged
BO3 groups and then the amount of non-bridging oxygens increases [70]–[73]. Consequently, the addition of network modifiers leads to first a higher network connectivity, reaching a peak and then, by adding a higher amount of network modifier to a lower network connectivity [68], [71]. This phenomenon is widely accepted as the so-called borate anomaly [70].
BGs are mainly produced by two techniques, the melt-quenching and the sol-gel method. By polycondensation of organic precursors, BGs can be produced by the sol-gel method. However, in this thesis the scope lies on BGs produced by melt-quenching technique. In this method, suitable precursors of BGs oxides (for instance in the form of high purity carbonates as granules or powders) are melted in a metal or ceramic crucible (e.g. platinum) at high temperatures, between 1000-1500°C, depending on the composition. By performing the cooling rapidly, crystallization can be inhibited and a glass can be obtained [54], [74], [75]. In general, the tendency to crystallize of BGs is higher than that of conventional glasses due to the lower network connectivity compared to conventional glasses. Therefore, in order to be able to process BGs at high temperatures (to obtain fibres, etc.), the so-called hot working range is an important property of BGs. The hot-working range is defined by the glass transition temperature and the onset of crystallization [54]. In order to develop BGs with a large hot working range, the ratio of alkali and alkaline earth metal cations needs to be low.
This could be for example obtained for the 13-93 BG composition (in wt.% (5.5 Na2O, 11.1
K2O, 4.6 MgO, 18.5 CaO, 3.7 P2O5, 56.6 SiO2), having a higher hot working range than 45S5 BG [54], [75]. However, this goes along with the increase of network connectivity, which in turn can lead to a decrease of solubility and therefore bioactivity [54]. 2 Fundamentals 13
2.2.2 Mechanisms of bioactivity
To explain why BGs are named “bioactive” and how BGs provide their favourable properties, the term bioactivity must be first defined. In general, bioactive materials are materials designed to induce a specific (positive) biological reaction during implantation [75]. The underlying mechanism, by which BGs lead to favourable biological reactions, is their ability to dissolve by surface reactions in contact with aqueous media. Due to the above described BG structure (section 2.2.1), BG exhibit a more open glass network structure compared to conventional glasses. BGs exhibit the ability to dissolve during their immersion in an aqueous solution since water molecules can penetrate easily into the network [54]. The ability of BG to dissolve is strictly related to network connectivity and can be therefore easily manipulated by changing the composition of BG [75]. During the dissolution of BGs, different processes occur determining the bioactivity of BGs, key steps in the bioactive behaviour are the release of ions and the formation of an apatite surface layer in contact with biological fluids [54].
Hench et al. described the proposed stages occurring during the formation of an apatite surface layer on the initial 45S5 composition (Figure 6) [76]. The first five stages happen during in vitro (without growth factor adsorption and cell attachment) as well as in vivo dissolution of BGs, whereas the remaining steps (6-8) occur during placement in a living tissue [75]. As shown in Figure 6, the first 5 steps occur within the first 24 hours and lead to an increase of pH. By measuring the increase of pH and the release of ions, the dissolution mechanism and simultaneous apatite formation can be observed [77]. Although the glass composition has a great influence on the rate of apatite formation, these steps are valid for all silicate based BGs [47]. A rate-limiting step for the dissolution of some silicate BGs was found to be the silica-rich layer, which hinders the transport of ions leading to the deceleration or stopping of the apatite formation [51].
In case of borate BGs, the formation of an apatite layer is similar as described for silicate BGs in Figure 6. The main difference is shown in Figure 7. In contrast to silicate BGs, borate BGs form an apatite layer without the formation of silanol bonds and hydrated silica gel [78]. For borate BGs, the rate-limiting step is therefore not the silica-rich layer, it is rather the ion diffusion to the interface [79]. By comparing silicate and borate BGs (of equivalent composition, just by replacing SiO2 by B2O3 in mol.%), it could be found that the pH increased faster and reached a plateau earlier for borate BGs (after around 50h for borate BGs in contrast to around 500h for silicate BGs) [56]. Moreover, the end pH was lower for borate BGs
(9.6) than for silicate BGs (11.5), most probably due to stronger acidic nature of B(OH)3 than 14 2 Fundamentals
Si(OH)4 [80]. This proves the lower chemical durability of borate BG compared to silicate BGs, described in section 2.2.1.
Log Time No Process in h Rapid cation exchange, in particular Na2+ and Ca2+ from the BG are
1 exchanged with H+ or H3O+ ions from the solution (e.g. blood) 1 PH increase and formation of silanol bonds on the surface High pH leads to the attack of Si-O-Si bonds by hydroxyl ions 2 Soluble silica is lost in the solution and more silanol bonds are formed Polycondensation and polymerization of silanol groups 3 Formation of hydrated silica gel
Diffusion of Ca2+ and PO43- through silica rich layer and (if present) 4 2 from the solution Formation of amorphous CaP-rich film on top of silica-rich layer Crystallization of CaP-rich film and formation of crystalline apatite (could be a mixture of hydroxycarbonate apatite and fluorapatite, 5 10 depending on BG composition) Growth factor adsorption and cell attachment on top of apatite layer Macrophage adhesion and preparation of implant site for tissue 6 repair 20 Osteoblast precursor cell arrival 7 Osteoblasts differentiation and proliferation Osteoprogenitor cells colonize surface of BG leading to production 100+ of various growth factors stimulating cell division, mitosis, and 8 production of extracellular matrix proteins Onset of ECM mineralization
Figure 6: Schematic representation of the steps leading to the formation of an apatite layer on 45S5 BGs [47], [75], [76]
2 Fundamentals 15
A)
B)
Figure 7: Comparison of apatite formation of A) silicate and B) borate BGs (inspired by refs [47], [78], [79])
By replacing partly silica with boron trioxide, borosilicate BGs are developed [81].
These BGs have the advantage that by varying the relative contents of SiO2 and B2O3, the dissolution rate and therefore the overall bioactivity of the BGs can be tailored [82]. As shown in Figure 8, borosilicate BGs form, similar to silicate BGs, a silica-rich layer during dissolution. However, this layer is porous, which leads to the growing of the apatite layer inwards. Therefore, in the case of borosilicate BGs, the overall dissolution reaction is controlled by the silica-rich layer (similar to silicate BGs) or by the ion diffusion (similar to borate BGs) [81].
Figure 8: Proposed steps of dissolution and apatite formation of borosilicate BGs (inspired by ref [81])
As mentioned above, there are two mechanisms responsible for the bioactivity of BGs. Besides the mechanism of apatite formation, ions released from BGs affect cell proliferation, mineralization as well as gene expression. Especially ions acting as modifiers (see section 2.2.1), can be easily released through the glass network by ion exchange with the surrounding body fluid and therefore are available in the human body. The local release of ions from BGs offers the advantage of having the therapeutic effect at site resulting in an optimized effect 16 2 Fundamentals with minimized side effects. Moreover, ions released from BGs are normally continually released, which offers a sustained therapeutic effect and additionally the release rate can be easily tailored by changing the network connectivity [54]. Depending of the composition of the BG, different ions can provide various biological effects, as summarized for selected ions in Table 1.
Table 1: Biological effects of various ions released from BGs
Ion Therapeutic effect Favor osteoblast differentiation and proliferation [28], [54], [75] Favor apatite precipitation [54] and ECM mineralization [75] Acceleration of blood-clotting [75], [83] and stimulation of angiogenesis Calcium [75], [84] Mediates interaction and activation of different factors of the coagulation process [12] At high levels, increases keratinocyte differentiation [12] Stimulates new bone formation [28], [75] and impeded apatite crystallization [54] Magnesium Promotes bone cell adhesion and stability [75] Stimulates of migration and proliferation of microvascular cells [84] Stimulates RNA synthesis in fibroblasts [54] Stimulates vascularization and angiogenesis [83] Boron Dietary boron stimulates bone formation [28] Regulates immune and inflammatory reactions and macrophage polarization [12] Key role in formation and calcification of bone tissue [28], [75] Silicon Stimulates of collagen type I production [75], [83] Promotes of neovascularization [75], [83], [84] Stimulates the expression of matrix Gla protein for bone [28], [83] Phosphorus Stimulates migration and tube formation in the human umbilical vein endothelial cell (HUVEC) model [84]
However, the above described dissolution and therefore bioactivity mechanisms can also lead to some undesired effects. The increase of pH due to the ion exchange in the early stage of dissolution can have a negative impact on the metabolism and function of cells. Moreover, the high pH can have an influence on cell respiration, affecting diffusion of gases 2 Fundamentals 17 and nutrients to the cells. However, this negative impact of the pH is normally not observed in vivo [85]. On the other hand, the raise of pH and the resulting raise of osmotic pressure are also believed to be the reason of the antibacterial effect of BGs [86]. Besides being antibacterial, the described two mechanisms of bioactivity lead to several advantageous properties, e.g. having good biocompatibility, bonding to hard and soft tissue, activate osteogenic genes and foster cell growth [87].
2.2.3 Ion doping
In order to further extend the functionality of BGs, it is a common approach to introduce biologically active ions which provide extra effects on cells [84], [88]–[90]. Glasses, in contrast to crystals, are not dependent on a specific stoichiometry. Therefore, the relatively large flexibility (see section 2.2.1) in the composition of BGs allows the addition of various kinds and amounts of ions [54]. Typical examples are silver, gallium, cobalt, lithium, manganese, etc. [75], [84], [88], [89]. However, since zinc and copper will be used in this work, a more detailed overview about these two ions in biomedical applications will be given here.
Zinc is a well-known essential trace element. It is needed in more than 300 enzymes, where it is involved in catalytic reactions and/or in the stabilization of protein structures [91]. Zn is known to be essential in wound healing [54], to stimulate bone formation [75], [83] and angiogenesis and to enhance nerve regeneration [83]. Moreover, it has some anti- inflammatory effects [75] and is crucial in the immune function of the human body [12]. Most interestingly, zinc exhibit some bactericidal behaviour [54], [92], probably due to the fact that zinc binds to membranes of bacteria, therefore prolonging the lag phase leading to an increased cell division time [93]. Copper, similar to zinc, offers also antibacterial properties. The mechanism behind the bactericidal behaviour is believed to be DNA degradation, protein oxidation, generation of reactive oxygen species and lipid peroxidation [75]. Additionally, copper is well-known for its stimulation effects on angiogenesis through mimicking hypoxia conditions [75], [94]. Moreover, Cu ions are known to upregulate VEGF gene expression as well as activating other proangiogenic factors (e.g. bFGF, TNF-alpha, IL-1) [84], leading to stimulation of the proliferation of endothelial cells [94]. Besides providing antibacterial and angiogenic effects, copper is known to accelerate wound healing in vivo [75] and to increase the differentiation of mesenchymal stem cells [28], [83].
Copper and Zinc, as well as other therapeutically active ions, offer advantages compared to synthetic drugs. In contrast to antibiotics, bacteria are not able to develop a 18 2 Fundamentals resistance against the antibacterial properties of BG [33]. In a study where bacteria were exposed recurrently to BGs, even minor concentrations of BGs were still effective against the bacteria [32]. Additionally, BGs are incorporated into the wound dressing and therefore applied directly on the site where the therapeutic effect is needed. Therefore smaller amounts with less systemic side effects are needed compared to oral applications of drugs [95]. Moreover, it is also a common strategy to incorporate growth factors into a wound dressing [96]. However, there exists several problems with the degradation and bioavailability of growth factors, possible leading to multiple side effects [9], [95].
Besides providing positive effects, ions can also have some cytotoxic effects. Therefore, the kind as well as the amount of the selected ion needs to be carefully chosen. This is of great importance in the case of ions offering additional therapeutic effects (e.g. Cu, Zn). In a certain concentration range, these ions lead to the required biological effect, therefore a specific release rate as a function of time needs to be achieved by choosing the optimal amount of ions incorporated into the BG. In case of higher doses (outside the therapeutic window), the ion released could lead to cytotoxic effects, whereas at lower doses no therapeutic effect is achieved [28], [62], [97]. It is also important to consider the possible synergistic effects which could occur by combining different ions inside the BG network, possibly leading to more efficient bioactive behavior or to a toxic effect. Since ions do not act individually but in combinations, it is also important to consider the time-dependent ion concentration ratio, e.g. of the added ions to the basic ions forming the glass (e.g. Si, B, P).
2.2.4 Applications
As already mentioned, BGs were mainly investigated for their use in contact with hard tissue. Therefore it is not surprising that commercial products made out of BG are mainly available in the field of bone substituting materials and regeneration (e.g. NovaBone®, PerioGlas®, BoneAlive®) as well as dental applications (e.g. NovaMin®, BioMin®) [75], [98], [99]. However, borate BG nanofibers are commercially available, trademarked as DermaFuseTM/MirragenTM, in order to treat long-term venous stasis ulcers in diabetic patients [82], [100], [101]. Additionally, wound care films dressings containing Ag-doped phosphate BG particles (antimicrobial Arglaes® film) are available on the market and used to control infections by the sustained release of silver [100]. Besides these two products, some patents proposing the use of BG for wound healing have been filed [102]. For instance, one patent focuses on the use of BG in combination with chitosan as a haemostatic agent [103], whereas another patent is focusing on hydrocolloid dressing made out of salicylic acid, a cellulose 2 Fundamentals 19 derivative, BG particles and water in order to promote the formation of granulation tissue [104]. Besides patents and commercial products, a huge number of researches are focusing on the use of BG in wound healing, treatment of burns and soft tissue applications [26], [27], [77], [83].
One recent study focusing on ion-doped borate BG (similar to the BG that will be used in this study) was done by Liu et al. [105], who examined the effect of ion doping on the properties of 13-93 B3 glass fibres, one doped with CuO and ZnO and the other one doped with CuO, ZnO, Fe2O3 and SrO. The investigation of the reaction of glass fibres in SBF showed that the ion doping had little effect on the degradation and conversion of the BG into amorphous calcium phosphate (ACP). However, the crystallization of ACP to hydroxyapatite was inhibited, which could be beneficial for wound healing. Further, Zn and Fe ions were not released, instead they remained mainly in the ACP or HA product unlike Cu and Sr [105]. In another study performed by Zhao et al [29], the effect of copper doping on the properties of borate bioactive glass fibres as well as their potential use in form of fibres for wound dressings in vitro and in vivo was examined. First, they studied the release of copper from the doped microfibers and its effect on the response of cells. The results showed that the ionic dissolution products promoted HUVEC migration, tubule formation and secretion of VEGF. Second, they studied the potential of the microfibers to heal full-thickness skin defects. The copper doped fibres were seen to stimulate angiogenesis to a higher extent than undoped fibres. However, no significant difference was found between collagen deposition, maturity and orientation of defects treated with doped and undoped fibres [29]. By combining borosilicate BG doped with Copper with poly (lactic-co-glycolic acid) (PLGA), Hu et al. [94] fabricated a composite dressing for wound healing applications. Moreover, by additionally loading vitamin E (known to have anti-inflammatory properties) into the wound dressing, the biological function could be further improved. In vitro tests showed that the multifunctional dressing successfully induced VEGF secretion and the migration and tube formation in HUVECs, thus having an angiogenic effect. Additionally, in vivo tests showed improved collagen remodelling and vessel sprouting [94]. These selected studies pointed out the promising results, which could be achieved by using borate BG for wound healing applications, which could be further enhanced by combining BG with other materials to create a tuneable wound dressing with specific properties (e.g. mechanical properties, degradation rate, bioactivity, etc.) suitable for treating different types of wounds, including chronic wounds. 20 2 Fundamentals
2.3 Wound dressing fabrication methods
By using BG as a material of choice for developing wound dressings, the first task (as mentioned in section 2.1.3) is solved. In order to solve the second tasks, the choice of a suitable fabrication method, several different techniques are available, which are suitable to produce wound dressings. Hydrogels, which are able to form a hydrophilic, insoluble and swelling wound dressing, are well known to maintain a moist environment and to improve the transmission of moist vapour and oxygen [9], [106]. Moreover they are known to reduce pain, they are not adherent and have a cooling effect. However, they are not suitable for wounds producing a lot of exudate since they would absorb too much leading to maceration. This fluid accumulation could lead to bacterial proliferation and therefore to infectious wounds [9], [107]. On the other hand, modern wound healing films are semi-permeable adhesive sheets, which provide a suitable barrier to prevent bacterial infection, however such dressings are also not suitable for exudative wounds [9]. By using the solvent casting/particulate leaching technique, where a porogen is incorporated during the film production and then leached out, porous films are obtained. Although the method is simple and inexpensive, limitations are given in the resulting shapes and in the possible presence of residual solvents [108]. In this work, freeze drying, electrospinning and 3D printing were the fabrication method chosen, due to their ability to be applied for deep, chronic and exudative wounds.
Freeze Drying or lyophilisation
To produce foams by freeze drying or lyophilisation, a polymer solution needs to be frozen [9]. During the freezing process, solvent ice crystals are formed, which are surrounded by polymer aggregates. During the second step (drying), the surrounded pressure is reduced to a lower level than the mean pressure of the frozen solvent and the solvent is triggered to undergo sublimation without passing the liquid phase (Figure 9). After complete drying/sublimation of the solvent, a polymer foam with an interconnected porous structure can be obtained [109]. Freeze drying offers the advantage that different pore sizes can be fabricated by changing several parameters (e.g. concentration of polymer solution, choice of solvent, freezing temperature) [108]. Freeze dried foams can be used as wound dressings since they are known to keep a moist environment and the temperature [2]. Moreover, during the application of freeze-dried foams, they turn into a gel, which can absorb excess exudates. This makes them excellent candidates for drug delivery. In contrast to hydrogels, foams which turn into gels can allow the diffusion of the contained drug without being diluted and having the risk to lose some of the incorporated drugs [9]. However, such foams have weak mechanical properties and are not suitable for dry wounds [2], [9]. 2 Fundamentals 21
Figure 9: The phase diagram of a solvent showing exemplary how a solvent pass from the solid phase into the gas phase by sublimation during freeze drying (inspired by ref [110])
Electrospinning
By setting up a high intensity electric field between an injection needle and a collector, which is commonly grounded, fibres can be produced by the electrospinning process (Figure 10) [111]. A polymeric solution, which is filled in a needle or small capillary connected to a pump, is gradually expelled through the injection needle, leading to the formation of a semi spherical polymer solution globule of liquid at the end of the needle. By rising up the intensity of the electric field, the hemispherical exterior of the formed solution globule extends, leading to the formation of a cone-shape (Taylor cone, Figure 10) [112]. When the electric field approaches a crucial value (repulsive electrical forces exceed surface tension forces), a charged jet of the polymer solution can be expelled from the Taylor cone. The solvent of the polymer solution evaporates during time of flight and the obtained polymer fibres can be collected at the target [113], [114]. The collected fibres form a fibre mat with three dimensional architecture, which has a high porosity and a well interconnected open-pore structure [9]. The resulting mats can be easily tuned by changing the electrospinning parameters and solution properties (e.g. solvent, viscosity, voltage, flow rate, temperature, distance to collector, etc.) [112], [115]. By electrospinning, a wide range of polymers can be used and fibres in the micro- to nanoscale exhibiting high surface-to-volume ratio can be produced. Depending on the chosen polymer, toxic solvents might be needed leading to possible residuals inside the obtained fibre mats [113]. Wound dressings made out of electrospun fibre mats can mimic the structure of the ECM of the skin, leading to an optimal environment for the adhesion and proliferation of skin cells [115]–[117]. Moreover, 22 2 Fundamentals electrospun fibre mats are known to have the ability to absorb high amounts of exudates, to provide breathability and permeability for nutrients and gas, such fibrous mats and can also prevent bacteria contamination [2], [9], [115].
Figure 10: Schematic overview of the electrospinning process, showing the explained Taylor cone (inspired by ref [118])
3D Printing
3D printing in general describes a method based on layer-by-layer processing of a liquid or solid material to form gradually increasing 3d structures. Each newly formed layer should adhere during the printing process to the previous layer. By using computer-aided design models, different (and complex) 3D structures can be dictated to the 3D printer [109], [119]. Pore size, geometry and interconnectivity can be therefore controlled and even patient specific structures are possible [109]. By controlling the parameters (e.g. print speed, print head temperature, pressure, layer height, etc.), complex scaffolds in the micro- and macroscale can be created [109], [119]. Commonly used 3D printing techniques include inkjetting, laser- assisted deposition and extrusion. Here, extrusion based 3D printing was used (Figure 11). By applying pneumatic or mechanical pressure to a syringe, a solution/gel can be extruded through the nozzle to form continuous filaments. In a special field of 3D printing, the bioprinting or biofabrication, living cells, ECM components and other biomaterials are incorporated in the polymer based bioink and deposited by extrusion in order to build cell laden tissue constructs [109]. Bioprinting offers, beside the high printing resolution, the incorporation of high cell densities inside the 3D printed construct. However, the development of printable bioinks is still challenging and the shear stress effect on cell behaviour within the printing nozzles needs to be taken into account [109], [120]. 2 Fundamentals 23
Figure 11: Schematic principle of extrusion based 3D printing using mechanical (left) or pneumatic pressure (right) (inspired by ref [121])
As mentioned above, producing wound dressings by freeze drying, electrospinning or 3D printing requires a polymer solution. Therefore, besides choosing BG due to its excellent properties, as described in section 2.2, a suitable polymer must be chosen in order to fabricate multifunctional composite wound dressings. Polymers can be divided based on their origin, whereas typical representatives of naturally derived polymers are polysaccharides and proteins. Synthetic polymers, on the other hand, include polyesters, poly(caprolactone) (PCL), poly(lactic acid) as well as some types of polyurethane [108], [122]. Natural polymers are known for their biocompatibility and biodegradability, but their availability is limited, the properties are not always reproducible and exhibit poor mechanical properties [122]. In contrast, synthetic polymers have controllable chemical, mechanical and structural properties. However, due to the challenges associated with synthetic polymers, e.g. their possible cytotoxic behaviour [108], a natural polymer was chosen here.
2.4 Cellulose
Cellulose is the most extensively occurring natural polymer, which can be obtained from several renewable sources [123]–[125]. Since cellulose is the major substance in plant cell walls [126], it can be obtained from a variety of plants, e.g. hemp, cotton or flax. However, due to its high availability, the main commercial source is wood. Besides using plants as source, it is also possible to use cellulose produced by seaweed, fungi and bacteria [123]. Cellulose is easily accessible at low cost [126] and, depending on the source, it is available in different forms showing a variety surface chemistries, degrees of crystallinity and crystal structures [127]. Cellulose, as shown in Figure 12, belongs to the class of homopolysaccharides and consists of beta-D-glucopyranose units, which are linked by beta- 24 2 Fundamentals
1,4 glycosidic bonds [123], [126]. Cellulose is hydrophilic, but cellulose fibres contain due to intramolecular hydrogen bonds and the beta-1,4 configuration ordered crystalline regions leading to its insolubility in water and rigid behaviour [128], [129]. Cellulose is known to accelerate wound healing through the maintenance and release of several growth factors (e.g. phosphodiesterase growth factor, epidermal growth factor and basic fibroblast growth factor) and to encourage the proliferation and relocation of dermal fibroblasts. Moreover, cellulose is known to disrupt the proliferation of bacteria [126]. Therefore, several studies have been carried out using cellulose as wound dressing material and clinical studies proved their high potential for the treatment of chronic wounds [130]–[134]. Moreover, a commercial cellulose- based dressing (Aquacell® Hydrofiber, ConvaTec, USA) is available on the market [135].
Figure 12: Chemical structure of cellulose [135]. Hydroxyl groups are marked in grey.
2.4.1 Methylcellulose
In order to develop new biomaterials with specific properties, polymers are usually chemically modified. Famous examples for cellulose derivatives are carboxymethylcellulose, hydroxypropylmethylcellulose, hydroxyethylcellulose and methylcellulose (MC) [129]. MC is the simplest cellulose derivative, where, as shown in Figure 13, hydroxyl groups are substituted by methyl groups [129]. To describe and distinguish different MCs, the degree of substitution (DS), the number of CH3-groups divided by the number of glucose units, is an important property. The DS can vary from 0 to 3, whereas MCs with DC between 1.3 and 2.5 are water soluble [129]. If the DS is lower than 1.3, too many sufficient hydrogen bonds remain resulting in non-solubility in water. However, if the DS is too high, MC is hydrophobic and therefore also insoluble in water [104]. MC can be fabricated by etherification of cellulose, which means by a reaction between cellulose, alkali and an etherifying agent. The synthesis starts by swelling cellulose fibres in an alkaline medium. The resulting alkali-cellulose reacts then with the etherifying agent, e.g. iodomethane, chloromethane or dimethyl sulphate. The etherification is followed by purification (removal of by-products by washing in hot water), 2 Fundamentals 25 drying and pulverization. By adding toluene, acetone or isopropanol, it is possible to obtain MCs with different DS’s [102].
Figure 13: Chemical structure of methylcellulose [128], [129], [136]. Methyl groups, which are substituting hydroxyl groups (grey), are marked in red.
MC is an interesting polymer for the fabrication of hydrogels due to its thermal behaviour. At temperatures below the critical solution temperature (LCST), water molecules surround the hydrophobic methoxyl groups resulting in the formation of a cage [128], [137]. Due to the formation of this cage, MC is lipophilic and hydrophilic at the same time and can be, in contrast to cellulose, dissolved in water. At temperatures below the LCST, the viscosity is quite constant at different temperatures [129]. At higher temperatures (above 30°C), the formed cages distort and break, resulting in the exposure of the hydrophobic groups [128]. With ongoing increasing temperatures, the viscosity strongly increases resulting in the formation of a gel [129]. By adding salting-in or salting out ions (in form of e.g. salts, sugar and alcohol), the LCST can be increased and decreased, respectively [137].
MC is approved by the U.S. Food and Drug Administration (FDA) for several applications due to its high biocompatibility [138], [139]. Therefore, MC has been used in pharmaceuticals as thickener, binder, stabiliser and emulsifier [129], [140]. Moreover, MC is an alternative to gelatine for the fabrication of capsules for nutrients [129]. Additionally, it is also used in other fields due to its adsorption properties (stabilization of paint), adhesion properties (printing), moisture retention (films) and thermothickening ability (cosmetics) [141].
2.4.2 Crosslinking methods
The properties of hydrogels are altered by cross-linking. Cross-linking leads in general to an increased elasticity (however, at some point the hydrogel will become too rigid), reduction of solubility, increased viscosity and increased mechanical properties [142]. Cross- 26 2 Fundamentals linking can be defined by the creation of chemical or physical bonds, which connect functional groups of two polymer chains to each other through covalent bonding or supramolecular interactions (e.g. hydrogen bonding, ionic bonding, etc.) [143]. Cross-linking can be performed by radiation, for instance by an electron beam or gamma radiation [144], by enzymatically cross-linking or by using small molecule crosslinkers such as glutaraldehyde, genipin or citric acid [143].
A)
B)
Figure 14: Crosslinking can be performed by A) chemical crosslinking using a synthetic or natural crosslinker and B) by ionic crosslinking (inspired by ref [144])
Due to the presence of hydrophilic groups, polysaccharides and therefore also cellulose/methylcellulose can be crosslinked [144]. Most commonly, single polysaccharides are chemically cross-linked by chemical or natural crosslinkers or by ionic cross-linking (Figure 14) [144]. Common examples for chemical cross-linkers are epichlorohydrin for cellulose, genipin for gelatine and glutaraldehyde for chitosan [144]–[146] as well as Ca2+ for ionic cross-linking of alginate [147]. A suitable crosslinker needs to be non-cytotoxic [143]. However, several well-known chemical cross-linkers, e.g. glutaraldehyde or formaldehyde, are known to be (relatively) highly cytotoxic [148]. Therefore, appropriate cross-linking agents of natural origin are preferable due to their higher biocompatibility and less cytotoxicity.
2.5 Manuka honey
For the treatment of infected wounds, it is clinical practice to use wound dressings additionally loaded with antibiotics. This offers the significant advantage that the antibiotics are delivered locally, exactly where they are needed [3]. However, due to the possible negative effects of antibiotics (especially in high doses), the decreasing number of new developed antibiotics and the rapid spread of antibiotic resistance of several bacteria [3], [22], 2 Fundamentals 27 phytotherapeutic agents, which are extracted from plants, are a great alternative to treat infections [149]. It is believed that, in contrast to antibiotics, phytotherapeutics retain their effectiveness against bacteria, since even with continuous exposure of the herbal extracts, resistance is unlikely to occur [22], [150]. Besides providing an antibacterial effect, phytotherapeutic agents are believed to offer a broad spectrum of different positive biological effects with less negative side effects [151].
Honey is known since ancient times for its beneficial effects on human health, especially in the treatment of wounds [152]. Nowadays, the acceleration effect of honey on the regeneration of different wounds, e.g. burns, infected wounds and surgical wounds is proven by clinical studies [153]–[159]. Honey is a sweet and viscous sugar-solution, having yellowish- brown colour, which is gathered by bees and other insects from the nectar of different flowers [160]. It contains more than 200 substances, a mixture of around 80-85% carbohydrates (responsible for the sweetness), 15-17% water, 0.1-0.4 % proteins and 0.2% ash. Additionally, it contains small quantities of vitamins (e.g. vitamin B6, niacin, thiamine), enzymes, amino acids, organic acids and minerals (e.g. calcium, copper, iron) [152], [160]. There is a wide variety of honeys, which differ in their biological properties, chemical compositions and medicinal value depending on their origin (floral source, bee species and geographical setting) [3], [160]. Honey is known to promote wound healing since it provides local nutrients to the wound, reduces inflammation and stimulates angiogenesis, granulation and wound epithelialization [161], [162]. Additionally, it was shown that a treatment with honey leads to thinner scars, which has both, a regenerative and aesthetic effect [160]. Moreover, honey was shown to have low to no allergic reactions and side effects, which could be also proven in direct contact with the skin cells (fibroblasts and keratinocytes) [163]. Besides the mentioned effects, honey is known to be antibacterial [164], antidiabetic [165], antimutagenic [166], antitumoral [167]–[169], antiviral [170] and antifungal [171]. Additionally, studies showed that honey is immunomodulatory by having pro- and anti-inflammatory effects and being antioxidant [172]–[176].
Several components and properties of honey are responsible for its beneficial effects on human health. The high sugar content inside the honey leads to high osmolality, whereas the low water content, the high viscosity together with the low pH forms a protective barrier against a variety of pathogens [3], [177]. The pH of honey is acidic, normally in the range of 3.4 to 6.1. It prevents biofilm formation and supports macrophages to eradicate bacteria [3], [38]. On the other hand, the mentioned osmotic pressure due to the high sugar content leads to the dehydration of microbes as well as to hindrance of microbial development [3]. Hydrogen peroxide, which can be produced by various honeys, has some bacteriostatic and bactericidal 28 2 Fundamentals properties. It is able to react with the cell walls of bacteria and intracellular lipids, nucleic acids and proteins. Whereas it is in low concentrations beneficial for the proliferation of endothelial cells and fibroblasts and for angiogenesis, it could be dangerous for cells and tissues in high concentrations [3], [152], [177]. Besides hydrogen peroxide, also other components of honey, e.g. antioxidants, lysozymes, flavonoids, phenolic acids, bee peptides and methylglyoxal, having some antibacterial effects [3]. This, as well as other components of honey, are also known to stimulate/inhibit the release of various cytokines (e.g. IL-1beta, IL-6) [152]. In summary, due to its components and the related effects, honey provides perfect conditions for wound healing by preventing a secondary dressing to stick to the wound bed, therefore decreasing the pain during dressing changes [177].
Due to variations between different honeys, the use of honeys with certified activities is recommended. The most popular example is Manuka honey (MH), some products containing MH are commercially available (e.g. Actilite®, MediHoney® and Algivon®) [3], [178]. MH is obtained from the flowers of the Manuka tree (Leptospermum scoparium), which grows in New Zealand [3] [152]. Besides the already mentioned actions of honey in general, MH has a high methylglyoxal content [3]. Therefore, MH is characterized by its unique Manuka factor (UMF), which is related to the concentration of methylglyoxal in MH. Since methylglyoxal derives from dihydroxyacetone and the high levels of dihydroxyacetone available in the Manuka nectar, the activity of MH can be certified based on its UMF [152]. Additionally to the already mentioned effects of honey, MH is known to impede the growth of multi-resistance S. aureus [3].
Besides using MH due to its beneficial effects on wound healing, in this study an innovative aspect of MH was explored, namely its use to crosslink MC. As reported in literature, glyoxal can be used to crosslink cellulose (as well as polyvinyl alcohol, starch and gelatine) [142], [179]. Therefore, it is hypothesized that the methylglyoxal inside the MH is able to form a chemical crosslink with the functional groups of MC. Therefore in this work, MC-MH was chosen as the platform to deliver BG in the development of novel wound dressings. 3 Borate and borosilicate glasses doped with copper and/or zinc 29
3 Borate and borosilicate glasses doped with copper and/or zinc1
3.1 Introduction
Based on the fact that the dissolution behaviour and the consequent ability to form ACP are crucial for the success of BGs, this Chapter focuses on the fabrication of different BGs designed in their chemical composition to exhibit bioactive and antibacterial properties and on their characterization. Besides the chemical, thermal and physical characterization of the produced BGs, the main focus will be on their dissolution behaviour, on their dose-depending effect on immune cells and on their antibacterial efficiency.
3.2 Experimental Procedure
3.2.1 Glass melting and glass powder preparation
Borate and borosilicate BGs based on the well-known silicate 13-93 composition were fabricated and additionally doped with copper and/or zinc (composition given in Table 2) by the melt-derived technique. The analytical grade reagents H3BO3, CaCO3, (CaHPO4)(2(H2O)),
K2CO3, Na2CO3, MgO, Cu(NO3)2x2.5(H2O), Zn(NO3)2 (all from Sigma-Aldrich) and Belgian quartz sand (obtained from Åbo Akademi, Finland) were used. The reagents were homogenously mixed and then melted in a platinum crucible according to their composition. In case of borate BGs, melting was carried out at 1050 °C for 2h, in case of borosilicate BGs, melting occurred at 1100 °C for 3h and in case of silicate BGs, melting was performed at 1360 °C for 3h. Subsequently after melting, all BG compositions were cast and annealed for 1h at 520 °C. After cooling down, the BG blocks were manually crushed. All BGs were melted twice to ensure homogeneity. Depending on the characterization technique used, BGs were further manually crushed and sieved to obtain size range fractions of < 45 µm, 45-90 µm, 90- 300 µm and 300-500 µm [180].
1 Some of the results presented in this chapter were previously published by the authors (as first author) and are hereby adapted from reference [180], [191] 30 3 Borate and borosilicate glasses doped with copper and/or zinc
Table 2: Nominal compositions of the fabricated BGs in wt.%
Name Na2O K2O MgO CaO B2O3 P2O5 SiO2 ZnO CuO 13-93 5.5 11.1 4.6 18.5 - 3.7 56.6 - - B3 5.5 11.1 4.6 18.5 56.6 3.7 - - - B3-Cu 5.5 11.1 4.6 15.5 56.6 3.7 - - 3 B3-Zn 5.5 11.1 4.6 17.5 56.6 3.7 - 1 - B3-Cu-Zn 5.5 11.1 4.6 14.5 56.6 3.7 - 1 3 BS 5.5 11.1 4.6 18.5 36.6 3.7 20 - - BS-Cu 5.5 11.1 4.6 15.5 36.6 3.7 20 - 3 BS-Zn 5.5 11.1 4.6 17.5 36.6 3.7 20 1 - BS-Cu-Zn 5.5 11.1 4.6 14.5 36.6 3.7 20 1 3
3.2.2 Characterization methods
3.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX)
SEM (Leo 1530, Oberkochen, Germany) combined with energy-dispersive X-ray analysis (EDX, Thermo Scientific UltraDry, Madison, WI) was used to examine the surface morphology of different BG powders with particle size range of 500-300 µm. The SEM was operated at 1 kV and working distance of 2 mm. BG particles were fixed on a sample holder by carbon tape.
3.2.2.2 Nuclear Magnetic Resonance (NMR) 2
29Si, 11B and 31P solid state Magic-Angle-Spinning NMR (MAS-NMR) spectra of the different fabricated BGs (particle size of below 45 µm) were measured on an APOLLO console (Tecmag Inc., USA) at a magnetic field of 7.05 T. The magnetic field was produced by a superconducting magnet (300 MHz/89mm, Magnex). To spin the sample at 8 kHz, a high- speed MAS probe (Bruker HP-WB) equipped with the KEL-F cap and 4 mm zirconia rotor was used. The frequency scale in ppm was referenced to TMS for 29Si, to H3BO3 for 11B and to H3PO3 for 31P. In order to be able to precisely analyse the spectra measured by NMR, the Bruker OPUS 7.2 software was used to decompose the spectra. With a starting ratio of 0.5 set of
2 NMR measurements were carried out at Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Cracow by Prof. Sitarz and Dr. Jeleń 3 Borate and borosilicate glasses doped with copper and/or zinc 31
Gaussian-Lorentzian peaks, the Levenberg Marquardt iteration algorithm [181] was used. The highest value of local residue root-mean-square deviation error was recorded at approx. 2.
3.2.2.3 X-ray diffraction (XRD)
The phase compositions of the BGs (particle range <45 µm) were characterized by powder X-ray diffraction (XRD) measurements (Miniflex 600 HR, Rigaku, Japan) with Cu Kα radiation. Data were collected with a step size of 0.02ᵒover a 2θ range from 20ᵒ to 60ᵒ. Measurements were carried out directly using the BG powders and were performed in triplicates.
3.2.2.4 Differential thermal analysis (DTA) 3
By differential thermal analysis (Netzsch STA 449F1), the different BG transition temperatures (Tg) and the crystallization peak temperatures (Tp) were measured. Following parameters were used: heating rate of 20 °C/min up to 1400°C for 13-93 BG, up to 1150°C for borosilicate BG and up to 1100°C for borate BG. 50 mg of each BG powder (particle size range 45–90 µm) were measured in a platinum pan in nitrogen atmosphere.
3.2.2.5 Hot stage microscopy (HSM) 4
Using HSM (Misura 3.0, Expert System), the sintering behaviour of the silicate, borosilicate and borate BGs was studied [182]. Following parameters were used: heating rate of 5 °C/min to fusion temperature (Tf) + 50 °C. A cylindrical sample with a height of 3 mm and a diameter of 2 mm was pressed from the powdered BG sample (particle size range < 45 µm) and imaged after every 5 °C increase in temperature. Then, the sintering curve was plotted using the height as a function of temperature and the onset of crystallization temperature Txi was determined. Isotropic densification behaviour was assumed [183]–[185].
3.2.2.6 FTIR
FTIR (Nicolet iS50, Thermo Fisher Scientific) was performed on ground BG particles (particle size range 300-500 µm) before and after immersion in simulated body fluid (SBF). The conditions for this analysis were as follow: spectral range 4000 to 400 cm-1; window material, CsI: 16 scans and resolution 4 cm-1.
3 Facility available at: Johan Gadolin Process Chemistry Centre, Åbo Akademi University, Turku, Finland 4 Facility available at: Johan Gadolin Process Chemistry Centre, Åbo Akademi University, Turku, Finland 32 3 Borate and borosilicate glasses doped with copper and/or zinc
3.2.2.7 Static in vitro dissolution
To evaluate the dissolution behaviour and the resulting bioactivity of the different BG compositions, static in vitro dissolution tests were performed using three different dissolution media. To evaluate the ion release, 50 mM Tris-buffered (TRIS) solution and to mimic the in vivo degradation of the BGs in a possible acidic environment during wound healing, 0.5 M lactic acid (LA, Sigma Aldrich) solution (pH=2) was used. Further, in order to prove the bioactivity, SBF was used and prepared according to the protocol of Kokubo [186]. To adjust the pH of SBF and Tris-buffered solution to 7.4, 1 M HCL was used.
1.5 g/L of the different BGs powders (particle size range 300-500 µm) were immersed in each solution (in accordance to the literature [187]) and placed in an incubating orbital shaker agitated at 100 rpm at 37 °C. After 3h, 8h, 24h, 48h and 72h of immersion, the pH (FiveEasy Plus pH meter FP20) was measured. The filtered dissolution solutions were collected for further analysis using inductively coupled plasma atomic emission spectroscopy (ICP-OES, PerkinElmer Optima 5300 DV, Shelton, CT) in order to measure the amounts of ions in the different dissolution media. All samples were measured in triplicates. After immersion in SBF, BG particles were additionally collected and stored until they were examined by EDX, XRD and FTIR (see above).
3.2.2.8 Continuous in vitro dissolution 5
The dissolution kinetics of the BG particles (particle size range 300-500 µm) was additionally measured with a continuous flow-through-cell at 37 °C (Figure 15). The used dissolution medium was SBF, prepared as described in section 3.2.2.7. The cell containing the BGs particles was connected to a peristaltic pump with a pumping rate of 0.2 ml/min. The tubing, the cell and the used filters (pore size 0.3 µm) were made out of Teflon. In order to avoid contaminations, new filters were used for each BG composition. Moreover, the tubing and the cell were cleaned with 0.1 M HCL and ultra-pure water after each measurement. Each cell contained 280 ± 5 mg of the BGs particles [188]. Used dissolution media (3ml) were collected after 3h, 8h, 24h and 48h, the pH of the collected dissolution media was measured and the liquids were stored in the fridge for further analysis with ICP-OES (PerkinElmer Optima 5300 DV, Shelton, CT). All samples were measured in triplicates.
5 Facility available at: Johan Gadolin Process Chemistry Centre, Åbo Akademi University, Turku, Finland 3 Borate and borosilicate glasses doped with copper and/or zinc 33
Figure 15: Schematic structure of the continuous flow-through-cell experiment
3.2.3 Antibacterial test
In order to evaluate the antibacterial potential of the fabricated BGs, an agar diffusion test using Staphylococcus aureus (Gram-positive) and Escherichia coli (Gram-negative) bacteria (obtained from the Microbiology Department of the University of Erlangen- Nuremberg) was performed. Based on a previously developed protocol [189], bacteria were expanded and agar plates were prepared by covering a petri dish of 10 cm diameter with a uniform layer of agar (LB Agar (Lennox), Lab M Ltd.). Samples were placed on the agar in form of pellets with a diameter of 10 mm (0.3 g of milled BG (particle size 5-20 µm) were pressed using a hydraulic press (PE-010, Mauthe Maschinenbau) with a load of 1t) and incubated at 37 °C in an orbital shaker. After incubation for 24h, the inhibition zones around the BG pellets were measured using ImageJ (calculated from 10 measurements). All samples were done in triplicates for S. aureus and E. coli, respectively.
3.2.4 Cell culture
Conditioned cell culture medium (CCM) used for these experiments was prepared as followed. For fresh CCM, RPMI 1640 (Lonza, Veviers, Belgium) was used and additionally supplemented with streptomycin (100 mg/ml, Sigma), penicillin (100 U/ml, Sigma), L- glutamine (2 mM, Sigma), 2-mercaptoethanol (50 mM, C. Roth), 10% heat-inactivated fetal calf serum (FCS, Merck) and 10 mM 2-[4-(2-hydroxyethyl)piperazin-1-yl]ethanesulfonic acid (HEPES). In order to prepare conditioned CCM, 1 g of BG powder (particle size in the range of 300-500 µm) was added to 10 ml of CCM to form a 10% suspension. After 24h of incubation at
37°C in a humidified atmosphere of 95% air and 5% CO2, the remaining BG particles were removed and the collected conditioned CCM was labelled as 10%-CCM. In order to verify the influence of pH increase during the dissolution of BGs, conditioned CCM was also prepared 34 3 Borate and borosilicate glasses doped with copper and/or zinc without adding HEPES. In both cases, during the 24h of incubation time, the pH (FiveEasy Plus pH meter FP20) was recorded after 3h, 6h, 12h and 24h. To examine the concentration of IDPs in the 10%-CCM, ICP-OES (Thermo Scientific iCAP 6500) was used. All measurements were done in triplicates. The 10%-CCM was then diluted using fresh CCM in order to obtain 1% and 0.1%-CCM following a previously established protocol [190]. For cell culture experiments, pure CCM and CCM containing sodium hydroxide to adjust the pH to 8.5 (pH-CCM) were used [191].
Cell culture experiments using bone marrow derived dendritic cells (DCs) from mice were performed as described previously [191]6. Briefly, DCs were generated from C57BL/6 and BALB/C mice, after being approved by the animal ethical committee of the government of Unterfranken, Würzburg. The obtained immature cells were then cultured in contact with 10%-, 1%- and 0.1%-CCM and additionally, as control, in contact with fresh CCM and pH-CCM (pH 8.5 adjusted with NaOH) for 48h. Then using LPS (Escherichia coli 0127:B8, 100 ng/ml, Sigma), the DCs were matured for 12h and the percentage of living cells was determined by staining using 7-AAD (7-Aminoactinomycin D) and counting (BD FACS Canto II cytometer) the cells. By analysing the DC specific surface markers (CD11c (clone N418), MHCII (clone M5/114.15.2), CD86 (clone GL-1), CD80 (clone 16-10A1) and CD83 (clone Michel-19), all from BD Biosciences) using flow cytometry (FCS Express 5 Flow Cytometry Software), the effect of BGs on the DCs was further examined. Additionally, the effect of the cultured DCs in the presence of different conditioned media on DC-mediated allogeneic T-cell stimulation was evaluated by co-culturing the DCs incubated in conditioned CCM with allogeneic murine splenic cells (T-cells) for 72h. Then, the cell proliferation was assessed pulsing the cells with 3H-thymidine (1 µC/well) (PerkinElmer) for 16 h. Moreover, the quantity of interleukin IL-6, TNF-α, IL-1α, GM-CSF and interferon- γ (IFN-γ) in the supernatants of DC-T cell co-cultures was determined as reported previously [191].
6 This cell culture study was carried out at the Department of Immune Modulation, Universitätsklinikum Erlangen by Dr. Elisabeth Zinser and Ms. Lena Stich 3 Borate and borosilicate glasses doped with copper and/or zinc 35
3.3 Results
3.3.1 Structural characterization of the fabricated BGs
BS-Cu-Zn BS-Zn BS-Cu
BS
B3-Cu-Zn
Intensity (a.u.) Intensity B3-Zn B3-Cu
B3 13-93
20 25 30 35 40 45 50 55 60 65 2Q (°) A)
-1 -1 -1 700 cm 900-1100 cm 1300-1500 cm -1 1150-1300 cm BS-Cu-Zn BS-Zn BS-Cu BS B3-Cu-Zn
B3-Zn
Absorbance (a.u.) Absorbance B3-Cu
B3
-1 470 cm 13-93
400 600 800 1000 1200 1400 1600 1800 2000 -1 B) Wavenumbers (cm )
Figure 16: A) XRD patterns and B) FTIR spectra of the fabricated silicate, borate and borosilicate BGs based on the 13-93 composition. XRD patterns of the different BGs prove their amorphous structure. Identified peaks in FTIR spectra are discussed in the text.
Silicate, borate and borosilicate BGs were successfully produced by the melt- quenching route. By using XRD analysis, the amorphous structure of the BGs could be confirmed, although it should be mentioned that the XRD pattern is quite noisy and small peaks might be not visible (Figure 16A). To further characterize the BGs, FTIR measurements were performed (Figure 16B). The main absorption bands, which are typical for silicate BGs, can be found in the 13-93 silicate BGs from 900-1100 cm-1 and at 470 cm-1, which are attributed to Si-O-Si stretching and bending mode, respectively [192]. Moreover, the bands at 36 3 Borate and borosilicate glasses doped with copper and/or zinc
760-780 cm-1 can be attributed to the symmetrical stretching vibrations of the Si-O(Si) and Si- O(P) bridges [193]–[195]. In the produced borate BGs (B3, B3-Cu, B3-Zn and B3-Cu-Zn), absorption bands from 700 cm-1 as well as 900-1100 cm-1 related to B-O stretching mode of tetrahedral BO4 units were detected. Moreover, absorption bands from 1150-1300 cm-1 and
1200-1500 cm-1 related to B-O stretching of trigonal BO3 units can be found [78], [196]. Both main absorption bands mentioned for silicate and borate BGs can be found in the produced borosilicate BGs (BS, BS-Cu, BS-Zn and BS-Cu-Zn). This result is a clear sign that the structural network in the borosilicate BGs is formed by silica (SiO2) and boron trioxide (B2O3). Furthermore, no influence of the ion doping can be found on the measured absorption spectra of the borate and borosilicate BGs doped with copper and/or zinc.
A)
B)
Figure 17: The bold line shows the original A) 29Si NMR and B) 31P NMR spectra of the 13-93 BG, whereas the remaining thin black lines are the single peaks based on the decomposition process (dashed, bold line). 3 Borate and borosilicate glasses doped with copper and/or zinc 37
To characterize the structure of silicate and borate BGs, NMR measurements were additionally carried out7. The 29Si NMR spectrum of the 13-93 silicate BG (Figure 17 A) indicates the presence of mainly Q3Si (-99 and -94 ppm) and Q2Si (-88 and -81 ppm) units [197]–[199]. In addition, the low intensity peak at -106 ppm could be attributed to Q4Si units. However, this peak is rather low and could be also noise. According to Figure 17 B), phosphorus, in rather low amounts, is present in the form of Q1P units (-6 ppm), probably in the form of Si-O-P bonds or as pyrophosphate (P2O7)4-) [54]. The peak at +4 ppm can be assigned to orthophosphates ((PO4)3-) [200]–[203].
A)
B)
Figure 18: The bold line shows the original A) 11B NMR and B) 31P NMR spectrum of the B3 BG, whereas the remaining thin black lines are the single peaks based on the decomposition process (dashed, bold line).
7 NMR measurements were carried out at Faculty of Materials Science and Ceramics, AGH University of Science and Technology, Cracow by Prof. Sitarz and Dr. Jeleń 38 3 Borate and borosilicate glasses doped with copper and/or zinc
As shown in Figure 18 A, according to the 11B NMR spectrum of the B3 borate BG, boron is mainly in the form of tetrahedral and triangular coordination’s present. The dominant peak at -19 ppm can be attributed to BO4 units, whereas the peaks at -8, -13 and -25 ppm can be assigned to BO3 units. In accordance to the 31P NMR spectrum of the 13-93 BG, the 31P NMR spectrum of the B3 BG (Figure 18 B) also shows a peak at 3 ppm and -2 ppm. These peaks also indicate the presence of mainly orthophosphates.
Figure 19: Comparison of 11B NMR spectra of B3 BG doped with copper and/zinc. The bold line shows the original of the B3, B3-Cu, B3-Zn and B3-Cu-Zn BGs, whereas the remaining thin black lines are the single peaks based on the decomposition process (dashed, bold line).
By comparing the 11B NMR spectra of B3 BG with B3 BG-based compositions doped additionally with copper and/or zinc, differences can be observed (Figure 19). All spectra contain the already discussed peaks attributed either to BO4 units (around -19 ppm) or to BO3 units (around -8, -14 and -23 ppm). However, it seems that the intensities of the peaks are different, indicating that the addition of copper and/or zinc has a significant influence on the coordination of boron within the glass network. Especially in the case of copper containing B3-BG, the peaks at -8, -14 and -23 ppm seems to be increased in comparison to the peak at -19 ppm. Although it is not clear (due to the boron anomaly) [70] whether this leads to a disruption or to a strengthening of the network, the addition of ions seems to lead to an increase of BO3 units. Moreover, the effect seems to be greater for copper ions (observable in 3 Borate and borosilicate glasses doped with copper and/or zinc 39 both, B3-Cu and B3-Cu-Zn BG). However, it should be also mentioned that a greater amount of copper (3 wt.%) was incorporated compared to zinc (1 wt.%).
Figure 20: Comparison of 31P NMR spectra of B3 BG doped with copper and/zinc. The bold line shows the original spectra of the B3, B3-Cu, B3-Zn and B3-Cu-Zn BGs, whereas the remaining thin black lines are the single peaks based on the decomposition process (dashed, bold line).
To further examine the effect of the addition of copper and/or zinc on the B3-based BGs, the measured 31P NMR spectra of B3, B3-Cu, B3-Zn and B3-Cu-Zn BGs can be compared (Figure 20). In accordance with the already discussed B3 BG, all ion-doped B3 BG showed the same peaks attributed to orthophosphates (around 3 ppm) or to Q1P units (around -3 ppm). In comparison to the undoped B3 BG, an increase of the peak at around -2 ppm in comparison to the peak at 3 ppm can be observed in the spectra of the ion-doped B3 BG. This is an indication that more phosphorus is present in the form of Si-O-P bonds or as pyrophosphate (P2O7)4-), which might lead to an increased network stability. However, since this data is just qualitative and therefore the results obtained by NMR measurements are just indications, the effect of ion-doping on the borate BGs was further studied by thermal measurements and dissolution experiments. Moreover, due to the complex structure of borosilicate BGs containing three different possible network formers (silica, boron trioxide and phosphorus pentoxide), the structure of the produced borosilicate BGs was not studied here. 40 3 Borate and borosilicate glasses doped with copper and/or zinc
3.3.2 Thermal characterization of the fabricated BGs
Additionally to chemical and structural characterization, the fabricated BGs were thermally characterized using DTA and HSM (Table 3). In Figure 21, the thermal spectra of the silicate (13-93), the borate (B3) and the borosilicate (BS) BGs are shown. The glass transition temperatures (Tg) of each BG can be taken from the endothermic peak: 603 °C for 13-93, 520 °C for B3 and 538 °C for BS. From the exothermic peak that gives the temperature range at which crystals form, the crystallization peak temperature (Tp) can be taken: 1059 °C for 13-93, 726 °C for B3 and 773 °C for BS. Due to the fact that no significant differences between the undoped borate/borosilicate BGs and BGs doped with Cu and/or Zn could be measured, Figure 16 B shows only the results of 13-93, B3 and BS BGs, exceptional B3-Cu. In case of B3- Cu (as well as B3-Cu-Zn, not shown here), the addition of copper leads to a broader exothermic peak, which can be explained by the above mentioned effect copper has on the boron and phosphorus coordination within the B3-Cu and B3-Cu-Zn network (section 3.3.1).
BS
B3-Cu
DSC (a.u.) B3
13-93
0 200 400 600 800 1000 1200 1400 1600 Temperature (°C)
Figure 21: DTA curves of produced silicate (13-93), borate (B3) and borosilicate (BS) BGs as well as borate BGs doped with copper (B3-Cu)
Additionally to DTA measurements, HSM was performed in order to evaluate the sintering behaviour of the fabricated BGs (Figure 22). As shown exemplary in Figure 22 A, the temperature Tsi, at which the sample shrinks to a height of 95%, as well as the onset of crystallization temperature Txi (starting point of the plateau) can be obtained by HSM measurements [188]. All the obtained temperature values by HSM and DTA are summarized 3 Borate and borosilicate glasses doped with copper and/or zinc 41 in Table 3. Based on these values, the hot-working range can be calculated using the following formula (section 2.2.1) [188]:
∆푇 = 푇푥푖 − 푇푔 (1)
As summarized in Table 3, the widest hot working range could be measured for the silicate 13-93 BG, followed by the borosilicate BGs (BS). Consequently, the borate BGs (B3) offer the shortest hot working range.
110
100 13-93 Tsi 90
80
70
60
Txi Height (%) Height
50
40
30
20 400 500 600 700 800 900 1000 1100 1200 A) Temperature (°C)
110 B3 100 B3-Cu B3-Zn 90 B3-Cu-Zn BS 80 BS-Cu 70 BS-Zn BS-Cu-Zn
60 Height (%) Height 50
40
30
20
500 550 600 650 700 750 B) Temperature (°C)
Figure 22: Sintering curves of A) silicate 13-93 and B) of borate (B3) and borosilicate (BS) BGs measured by HSM
42 3 Borate and borosilicate glasses doped with copper and/or zinc
Table 3: An overview of the measured thermal properties of the fabricated BGs by DTA and HSM
Tg in °C Tp in °C Tsi in °C Txi in °C ΔT = Txi-Tg Glass (DSC) (DSC) (HSM) (HSM) in °C 13-93 604 1059 686 900 296 B3 520 726 568 603 83 B3-Cu 510 769 559 597 87 B3-Zn 516 762 561 610 94 B3-Cu-Zn 504 776 544 590 86 13-93-BS 538 773 586 645 107 BS-Cu 521 746 539 623 102 BS-Zn 533 767 574 628 95 BS-Cu-Z 521 730 574 620 99
In the literature [204]–[207], Tg of 590-620 °C and Tp of 920-1040 °C were reported for silicate 13-93 BG. Therefore, the here measured Tg and Tp of 13-93 BG are in good agreement with the literature. Additionally, a similar hot working range, around 250 °C for the 13-93 composition has been reported [77]. Besides the well reported thermal properties of the silicate 13-93 BG composition, there is less information available about the borate (and borosilicate) versions of the 13-93 composition.
3.3.3 Acellular dissolution behavior of BGs in different solutions
As mentioned in section 2.2.2, the bioactivity of a BG strongly depends on its dissolution behaviour. Therefore, the dissolution behaviours of the different BGs fabricated in this work were examined in contact with different dissolution media by measuring the change of pH and the release of ions at different immersion times.
3.3.3.1 Dissolution in TRIS under static conditions
TRIS was chosen as dissolution medium, since this solution does not contain any ions present in the BGs and therefore the ion release from BGs can be easily detected. As mentioned in section 2.2.2, during the dissolution of BGs, an increase of pH occurs. According to Figure 23, the silicate 13-93 BG leads to the lowest (and slowest) increase of pH, reaching a final pH of 7.65 ± 0.01. In contrast, both borate and borosilicate BGs lead to a faster increase of pH during dissolution, reaching a final pH in the range of 7.78 to 7.85. The findings fit well to 3 Borate and borosilicate glasses doped with copper and/or zinc 43 the already mentioned study [79], which found out that due to the lower chemical durability borate BGs dissolve faster, leading therefore to a faster increase of pH [79]. No significant influence of the ion-doping on the increase of pH could be found.
13-93 7.9 B3 B3-Cu B3-Zn 7.8 B3-Cu-Zn BS BS-Cu 7.7 BS-Zn pH BS-Cu-Zn 7.6
7.5
7.4 0 20 40 60 80 100 120 Time (h)
Figure 23: Change of pH versus immersion time of fabricated BGs in TRIS under static conditions
As shown in Figure 24, during dissolution the BG components are released in form of ions into the medium as a function of immersion time. As already shown by pH measurements, the silicate 13-93 BG dissolves the slowest and show, similar to the pH increase, an almost linear release of ions. Here, a special look should be taken on the release of the network former Si, since, based on the release of Si, an assumption can be made about the total dissolution behaviour of the silicate BG. Even after 96h of immersion time, only around 5 wt.% of the total amount of silicon was released from the 13-93 glass. In contrast, around 20% of Si+B (the network formers) and 30% of B was released from the borosilicate and borate BGs, respectively. The increasing release rate with increasing amount of boron trioxide is in accordance to the observation made by pH measurements (Figure 23). For all BGs, the release of network modifiers (Mg, Na and K) is higher compared to the release of network former. However, the release profile of the network modifier and the network former are for silicate (linear release), borosilicate (logarithmical) and borate (logarithmical) BGs similar, which is a sign that by adjusting the amount of boron trioxide inside the glass network, the release of all components of the glass can be tailored. Phosphorus seems to be the exception in terms of the release profile. As mentioned in section 2.2.2, P as well as Ca are involved in the precipitation of CaP-rich species. Here, an almost linear amount of P in the TRIS solution was measured after varying immersion times and additionally, in contrast to the remaining network modifiers, a slower release of Ca was detected. Both ions prove the occurrence of precipitation and therefore the bioactivity of the fabricated BGs according to Hench et al. [25]. 44 3 Borate and borosilicate glasses doped with copper and/or zinc
70 300 B 60 Si 250
50 200 40
150 30 B3 B3-Cu B3-Zn 20 100 13-93 B3-Cu-Zn BS BS 10 BS-Cu 50 BS-Cu
BS-Zn BS-Zn Cumulative concentration of B (mg/L) B of concentration Cumulative Cumulative concentration of Si (mg/L) Si of concentration Cumulative 0 BS-Cu-Zn BS-Cu-Zn 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
160 140 Ca 140 K 120 120 100 100 80 13-93 13-93 B3 80 B3 60 B3-Cu B3-Cu B3-Zn 60 B3-Zn 40 B3-Cu-Zn B3-Cu-Zn BS 40 BS 20 BS-Cu BS-Cu
BS-Zn 20 BS-Zn Cumulative concentration of K (mg/L) K of concentration Cumulative Cumulative concentration of Ca (mg/L) Ca of concentration Cumulative 0 BS-Cu-Zn BS-Cu-Zn 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
70 60 60 Na Mg 50 50 40 40 13-93 13-93 B3 30 B3 30 B3-Cu B3-Cu B3-Zn 20 B3-Zn 20 B3-Cu-Zn B3-Cu-Zn BS 10 BS BS-Cu BS-Cu 10
BS-Zn BS-Zn Cumulative concentration of Mg (mg/L) Mg of concentration Cumulative Cumulative concentration of Na (mg/L) Na of concentration Cumulative BS-Cu-Zn 0 BS-Cu-Zn 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
6 40
5 P 35 Cu
30 4 25 13-93 3 B3 20 B3-Cu B3-Zn 15 2 B3-Cu-Zn BS 10 B3-Cu 1 BS-Cu B3-Cu-Zn
BS-Zn 5 BS-Cu Cumulative concentration of P (mg/L) P of concentration Cumulative
BS-Cu-Zn (mg/L) Cu of concentrations Cumulative BS-Cu-Zn 0 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
Figure 24: Ion release during dissolution of the different investigated BGs in TRIS, measured by ICP-OES 3 Borate and borosilicate glasses doped with copper and/or zinc 45
The addition of copper to the borosilicate BGs (BS-Cu and BS-Cu-Zn) was shown to increase the release of Si. In contrast, the incorporation of Zn seems, to a certain extent, to inhibit the release of Si. In case of ion-doping of borate BGs, no significant influence on the network former could be found. However, the addition of copper (B3-Cu and B3-Cu-Zn) seems to lead to a slower release of Ca (or a higher rate of precipitation). For all Zn-doped borate and borosilicate BGs, the release of Zn could not be measured, probably due to the fact that the released amount was below the detection limit of ICP-OES. In contrast, copper released from borate (B3-Cu and B3-Cu-Zn) and borosilicate (BS-Cu and BS-Cu-Zn) BGs reached final amounts of 35 ± 1.9 mg/L and 30 ± 0.9 mg/L after 96h of immersion time, respectively.
3.3.3.2 Dissolution in SBF under static and dynamic conditions
7.9 Static SBF Dynamic SBF 7.9
7.8 7.8 7.7 13-93 13-93 B3 7.7 pH pH 7.6 B3 B3-Cu pH B3-Cu B3-Zn 7.6 B3-Zn 7.5 B3-Cu-Zn B3-Cu-Zn BS BS BS-Cu 7.4 7.5 BS-Cu BS-Zn BS-Zn BS-Cu-Zn BS-Cu-Zn 7.3 7.4 0 20 40 60 80 100 120 0 10 20 30 40 50 60 Time (h) Time (h)
Figure 25: Change of pH versus immersion time of BGs in SBF under static (left) and dynamic (right) conditions
Based on the suggestion of the Technical Committee 4 (TC04) of the International Commission of Glass (ICG) [77], SBF solution was used in addition TRIS to test the fabricated BGs. Moreover, the dissolution tests using SBF were done under static conditions and under dynamic conditions, in order to represent more closely the dynamic situation in the human body (where due to the blood flow a dynamic environment occurs) [208]. As shown in Figure 25, the final pH of all BGs is in the range of 7.65-7.70 after 96h of immersion time in static conditions in SBF. In contrast, under dynamic conditions, the highest pH could be measured at the beginning of the immersion time (around 7.8 for borate/borosilicate BGs). Just the silicate 13-93 BG seems not to react in the first 3h and shows just a minor increase of pH over the complete immersion time. Here, it is important to mention that the pH during the dynamic dissolution experiments was measured at room temperature, with a reference pH of 7.64 ± 0.01 for pure SBF. After 3h, borosilicate and silicate BGs led to a nearly linear decrease of pH, reaching a final pH of 7.71 ± 0.04 for borate BGs (B3) and a final pH of 7.70 ± 0.07 for 46 3 Borate and borosilicate glasses doped with copper and/or zinc borosilicate BGs (BS). Additionally, as already found during dissolution experiments using TRIS, no measurable significant influence of ion-doping on the pH value was found.
150
60 125 Si B 50 100
40 75
30 B3 50 B3-Cu B3-Zn 20 13-93 25 B3-Cu-Zn BS BS 10 BS-Cu BS-Cu 0
BS-Zn BS-Zn Cumulative concentration of B (mg/L) B of concentration Cumulative Cumulative concentration of Si (mg/L) Si of concentration Cumulative 0 BS-Cu-Zn BS-Cu-Zn -25 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
140 55 Ca Mg
120 50
100 45 13-93 13-93 B3 B3 80 B3-Cu B3-Cu 40 B3-Zn B3-Zn B3-Cu-Zn B3-Cu-Zn 60 BS BS BS-Cu 35 BS-Cu
40 BS-Zn BS-Zn Cumulative concentration of Ca (mg/L) Ca of concentration Cumulative BS-Cu-Zn (mg/L) Mg of concentration Cumulative BS-Cu-Zn 30 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
15.0 40 P Cu 35 12.5
30 10.0
25 13-93 13-93 B3 7.5 B3 20 B3-Cu B3-Cu B3-Zn B3-Zn 5.0 B3-Cu-Zn 15 B3-Cu-Zn BS BS BS-Cu 2.5 BS-Cu 10
BS-Zn BS-Zn Cumulative concentration of P (mg/L) P of concentration Cumulative BS-Cu-Zn (mg/L) Cu of concentration Cumulative BS-Cu-Zn 5 0.0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
Figure 26: Release of ions during dissolution of the investigated BGs in SBF under static conditions (Note the different size of the y-axis)
Since the amount of sodium and potassium is quite high in SBF, only the released amounts of Si, B, Ca, Mg, P and Cu are shown in Figure 26. In comparison to TRIS, SBF seems to accelerate the dissolution of the different BGs. This is especially obvious for borate and borosilicate BGs, whereas only 6 wt.% of B and 12 wt.% of B+Si were released respectively. However, the acceleration effect was only minor for the silicate 13-93 BG. In contrast to TRIS, 3 Borate and borosilicate glasses doped with copper and/or zinc 47 in addition to Ca and P, also Mg seems to be involved in the precipitation (observable by the less amount of released Mg at 48h). Additionally, also the release of copper (30-35 mg/L) is less compared to the measured release of Cu during dissolution in TRIS.
Table 4: Content of the cations (mol.%) in the different BG surfaces before and after dissolution in SBF under static conditions according to EDX analysis
Glass Na Mg Si P K Ca B Cu Zn 13-93 5.0 2.5 25.6 1.4 9.2 14.2 48h 3.0 1.7 29.9 2.4 6.1 11.1 96h 2.5 1.4 32.2 2.7 4.5 9.4 13-93-B3 4.5 2.1 1.2 6.1 8.2 21.3 48h 0.7 1.2 5.4 1.6 19.2 17.3 96h 0.7 1.1 8.8 1.4 27.6 11.1 B3-Cu 4.3 2.2 1.4 7.9 9.7 19.1 2.3 48h 1.1 1.7 8.1 2.0 21.8 13.0 1.6 96h 0.7 1.3 5.2 1.7 15.9 10.8 1.1 B3-Zn 3.6 1.4 0.9 5.2 6.8 23.1 0.4 48h 3.1 1.3 2.3 4.8 16.1 15.0 1.2 96h 0.9 1.3 10.1 1.4 28.5 8.0 3.9 B3-Cu-Zn 3.2 1.3 0.7 4.6 5.1 24.1 1.3 0.3 48h 2.8 1.6 6.6 1.8 16.1 16.0 1.1 2.2 96h 2.4 1.4 5.6 1.3 20.9 10.1 0.9 2.7 13-93-BS 4.0 1.9 6.1 1.1 6.5 9.8 16.8 48h 0.6 1.0 8.2 4.0 1.3 11.7 16.5 96h 0.6 1.1 10.4 5.2 1.1 14.0 13.0 BS-Cu 4.7 2.2 7.1 1.4 6.9 8.7 15.1 1.7 48h 0.8 1.6 9.6 5.2 1.4 13.7 12.5 2.5 96h 1.0 1.3 8.8 9.2 1.3 18.7 8.2 1.1 BS-Zn 4.5 1.9 6.3 0.9 6.1 8.6 17.2 0.4 48h 0.7 1.4 12.3 3.3 1.8 8.0 14.1 1.5 96h 0.0 1.0 10.5 5.3 1.3 15.5 11.8 1.6 BS-Cu-Zn 3.9 1.7 5.3 1.0 5.1 6.5 18.7 1.5 0.6 48h 0.7 1.5 11.1 4.2 1.9 11.9 12.1 2.7 1.5 96h 1.0 1.2 14.2 5.8 1.6 10.3 9.4 2.8 2.1
48 3 Borate and borosilicate glasses doped with copper and/or zinc
Since SBF is a powerful tool to measure the bioactivity of BGs and is widely used [78], [79], [187], [209]–[211], remaining BG particles after 48h and 96h in SBF under static conditions were further analysed using EDX, FTIR and XRD. The results obtained by EDX (Table 4) correlate well with the results obtained by ICP-OES measurements (Figure 26). Here it is important to mention that due to the insensitivity of EDX for low atomic number elements, it is hard to measure the exact amount of B inside the borate and borosilicate BGs. However, as expected, a clear decrease of B (as well Mg, Na and K) can be observed. The increase of Si can be attributed to the formation of a silica-rich layer in silicate and borosilicate glasses (according to section 2.2.2 and literature [68], [78], [79], [81], [212]). Most interestingly, an increase of Zn could be measured for all zinc-containing BGs. Therefore, it seems that Zn, instead of being released in the dissolution media, is incorporated in the precipitation of CaP-rich species (increase of P and slow decrease of Ca).
By using FTIR, the formation of CaP-rich species on the surface of the BG particles after immersion in SBF for 96h was examined (Figure 27 A). Peaks at 1000-1100 cm-1 and at 1400- 1460 cm-1 corresponding to P-O and C-O (both stretching), respectively, could be found. Additionally, the peak at 875 cm-1, corresponding to C-O bending, and the peak at 500-600 cm-1, corresponding to bending vibrations of P-O groups, proved the formation of the CaP-rich species [213], [214]. In order to verify if the formed CaP-rich species crystallize to hydroxyapatite, XRD measurements were done on the remaining BG particles after 96h immersion in SBF (Figure 27 B). No sufficiently prominent peaks could be detected, probably due to the fact that the newly-formed hydroxyapatite is poorly crystallized or because the XRD was not sensitive enough to the minor amounts of crystalline hydroxyapatite formed, or a combination of both reasons.
In comparison to the dissolution of the BGs in SBF under static conditions and in accordance with the results obtained by pH measurements, the ion release profiles are completely different under dynamic conditions. Due to the fact that higher amounts of BG were used in the dynamic dissolution study in comparison to the static dissolution studies and the measured amount of ions is cumulative under static conditions, the ion released concentrations cannot be directly compared. As expected, the measured amounts of ions in the collected dissolution media after being in contact with the BGs were initially the highest, followed by a nearly constant release. A constant release and therefore supply with ions is favourable for the regeneration of the tissues. Moreover, no significant influence of doping the borate and borosilicate BGs with copper and/or zinc could be found. 3 Borate and borosilicate glasses doped with copper and/or zinc 49
-1 -1 -1 500-600 cm 1000-1100 cm 1400-1460 cm -1 875 cm BS-Cu-Zn BS-Zn BS-Cu BS B3-Cu-Zn
B3-Zn
Absorbance (a.u.) Absorbance B3-Cu
B3
19-93
400 600 800 1000 1200 1400 1600 1800 2000 A) Wavenumbers (cm-1)
BS-Cu-Zn
BS-Zn
BS-Cu
BS
B3-Cu-Zn
Intensity (a.u.) Intensity B3-Zn B3-Cu
B3 13-93 20 25 30 35 40 45 50 55 60 65 70 B) 2Q (°)
Figure 27: A) FTIR spectra and B) XRD patterns of the silicate, borate and borosilicate BGs after immersion in SBF under static conditions for 96h. Identified peaks are discussed in the text.
In contrast to dissolution experiments in static conditions, it was not possible to collect the remaining BG particles after dissolution experiments using dynamic conditions. Therefore, the remaining BG particles could not be further evaluated. In future studies, the setup of the dynamic experiment should be improved in order to be able to collect the BG particles.
50 3 Borate and borosilicate glasses doped with copper and/or zinc
50 Si 13-93 160 B 45 B3 BS B3-Cu 40 BS-Cu 140 B3-Zn BS-Zn B3-Cu-Zn 35 120 BS-Cu-Zn BS 30 100 BS-Cu BS-Zn 25 80 BS-Cu-Zn 20 60 15
40 Concentration of B (mg/L) B of Concentration Concentration of Si (mg/L) Si of Concentration 10
5 20
0 0 0 5 10 15 20 25 30 35 40 45 50 0 5 10 15 20 25 30 35 40 45 50 Time (h) Time (h)
60 13-93 200 13-93 Mg B3 Ca B3 B3-Cu 190 55 B3-Cu B3-Zn 180 B3-Zn B3-Cu-Zn 170 B3-Cu-Zn 50 BS BS BS-Cu 160 BS-Cu BS-Zn 150 BS-Zn 45 BS-Cu-Zn 140 BS-Cu-Zn
130 40
120 Concentration of Mg (mg/L) Mg of Concentration Concentration of Ca (mg/L) Ca of Concentration 110 35
100
90 30 0 5 10 15 20 25 30 35 40 45 50 0 5 10 15 20 25 30 35 40 45 50 Time (h) Time (h)
80 13-93 20.0 P B3 Cu B3-Cu 70 B3-Cu 17.5 B3-Cu-Zn B3-Zn BS-Cu 60 B3-Cu-Zn 15.0 BS-Cu-Zn BS 50 BS-Cu 12.5 BS-Zn BS-Cu-Zn 10.0 40
7.5 30
5.0 Concentration of P (mg/L) P of Concentration
20 (mg/L) Cu of Concentration 2.5 10 0.0 0 5 10 15 20 25 30 35 40 45 50 0 5 10 15 20 25 30 35 40 45 50 Time (h) Time (h)
Figure 28: Release of ions during dissolution of the investigated fabricated BGs in SBF under dynamic conditions (Note the different scales of the y-axis)
3.3.3.3 Dissolution in lactic acid under static conditions
During bacterial infections, a local pH decrease usually occurs. Therefore, the fabricated BGs were additionally tested in LA solution with a pH of 2. As shown in Figure 29, during dissolution an increase of pH depending on the used BG can be observed. The silicate 3 Borate and borosilicate glasses doped with copper and/or zinc 51
13-93 BG led to a comparable minor increase of pH (up to 2.14 ± 0.01), proving, in accordance with the dissolution experiments in SBF and TRIS, the lower reactivity of the silicate BG. In contrast, both borate and borosilicate BGs led to a higher pH increase, up to 2.29 ± 0.02. However, the trend during the dissolution was different. Borosilicate BG led first to an initial burst increase, followed by a decrease at 24h and then a slow further increase of pH was recorded.
2.35 13-93 B3 2.30 B3-Cu B3-Zn B3-Cu-Zn 2.25 BS BS-Cu 2.20
BS-Zn pH BS-Cu-Zn 2.15
2.10
2.05
2.00 0 20 40 60 80 100 120 Time (h)
Figure 29: Change of pH versus immersion time for different BGs in LA under static conditions
The results obtained by ICP-OES measurements support the observation made by pH measurements. As shown in Figure 30, almost no Si could be released from the silicate BG, whereas around half of the total Si content was released from the borosilicate BGs. In addition, B, K and Na were almost completely released from the borate and borosilicate BGs within the first 8h of immersion time in LA solution. Therefore, it could be observed that borate and borosilicate BGs dissolve quickly in an acidic environment, whereas the silicate BG almost does not react. In case of Ca, P and Mg, 70 wt.%, 45 wt.% and 60 wt.%, respectively, was released from borate and borosilicate BGs. The remaining Ca, P and Mg content in the BGs seems to precipitate to CaP-rich species (containing Mg). In contrast to the already mentioned ions, copper was released slower and reached a total amount of 85 wt.% after 48h. Most interestingly, in an acidic environment, the release of Zn could be measured. In case of Zn- doped borate BG, a total amount of 7.8 ± 0.1 mg/L of Zn and in case of Zn-doped borosilicate BG, a total amount of 8.2 ± 0.3 mg/L of Zn were released after 48h. For Cu-Zn-doped borate and borosilicate BGs, a total release of 8.5 ± 0.1 mg/L was measured. In case of Zn containing BGs around 56 wt.% could be released, which indicates that Zn is also partially involved in the 52 3 Borate and borosilicate glasses doped with copper and/or zinc precipitation of Ca, P and Mg (since no remaining BG particles could be found in the LA solution).
B3 160 Si 13-93 280 B BS B3-Cu 140 BS-Cu B3-Zn 260 B3-Cu-Zn 120 BS-Zn BS-Cu-Zn BS 100 240 BS-Cu BS-Zn 80 BS-Cu-Zn 220
60 200 40
20 180 Cumulative concentration of B (mg/L) B of concentration Cumulative Cumulative concentration of Si (mg/L) Si of concentration Cumulative 0 160 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
200 Ca 13-93 140 K 180 B3 13-93 B3-Cu B3 160 120 B3-Zn B3-Cu 140 B3-Zn B3-Cu-Zn 100 BS B3-Cu-Zn 120 BS BS-Cu 80 BS-Cu 100 BS-Zn BS-Zn BS-Cu-Zn 80 60 BS-Cu-Zn 60 40 40
20 20
Cumulative concentration of Ca (mg/L) Ca of concentration Cumulative Cumulative concentration of K (mg/L) K of concentration Cumulative
0 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
45 70 Mg Na 13-93 13-93 B3 40 B3 60 B3-Cu B3-Cu 35 B3-Zn B3-Zn 50 B3-Cu-Zn 30 B3-Cu-Zn BS BS 40 BS-Cu 25 BS-Cu BS-Zn BS-Zn BS-Cu-Zn 20 30 BS-Cu-Zn 15 20 10
10 5
Cumulative concentration of Mg (mg/L) Mg of concentration Cumulative Cumulative concentration of Na (mg/L) Na of concentration Cumulative 0 0 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h) 3 Borate and borosilicate glasses doped with copper and/or zinc 53
30 13-93 P B3 40 13-93 B3-Cu Cu 39 B3 25 B3-Zn B3-Cu B3-Cu-Zn 38 B3-Zn 20 BS B3-Cu-Zn 37 BS-Cu BS BS-Zn 36 BS-Cu 15 BS-Cu-Zn BS-Zn 35 BS-Cu-Zn 10 34
33 5
32
Cumulative concentration of P (mg/L) P of concentration Cumulative Cumulative concentration of Cu (mg/L) Cu of concentration Cumulative 0 31 0 20 40 60 80 100 120 0 20 40 60 80 100 120 Time (h) Time (h)
Zn 13-93 9.0 B3 B3-Cu B3-Zn 8.5 B3-Cu-Zn BS BS-Cu 8.0 BS-Zn BS-Cu-Zn
7.5
7.0 Cumulative concentration of Zn (mg/L) Zn of concentration Cumulative
0 20 40 60 80 100 120 Time (h)
Figure 30: Release of ions during dissolution of the different investigated BGs in LA under static conditions (note the different scales of the y-axis)
3.3.4 Antibacterial efficiency of the BGs
In order to evaluate the antibacterial potential of the fabricated BGs, an agar diffusion tests was performed. According to the Standard SNV 195920-1992, materials which are leading to an inhibition zone greater than 1 mm can be considered as antibacterial [189]. According to Figure 31, all produced BGs are efficiently antibacterial against gram-positive and gram-negative bacteria. The only exceptions were found for 13-93 BG, which showed no antibacterial effect against S. aureus and for BS, which showed a minor effect against E.coli. For borate and borosilicate BGs, an improved antibacterial effect was observed by the addition of Cu and/or Zn against the gram-negative bacterium E. coli. However, for both tested bacteria, the improvement was greater for copper containing BGs compared to Zn-doped BGs. In case of the gram-positive bacterium S. aureus, all BGs showed a greater influence compared to the gram-negative bacterium E. coli. The different sensitivity of gram-negative and gram-positive bacteria has been observed in literature [93] and could be explained by the differences in cell 54 3 Borate and borosilicate glasses doped with copper and/or zinc wall composition, cell physiology and structure and in the metabolism of the two kind of bacteria. In general, it is believed that gram-negative bacteria are through their highly organized compact structure better protected against penetration and diffusion of antibacterial materials/drugs [215]. This is in accordance with the observed antibacterial effect of the tested BGs, showing a greater effect against the gram-positive bacterium S. aureus in comparison to the gram-negative bacterium E. coli.
13 S. aureus 9 * E.coli 12 * * 11 * 8 * * 10 7 * * 9 6 8 * * 5 7 6 4 5
3 4 Inhibition zone (mm) zone Inhibition
Inhibition zone (mm) zone Inhibition * 3 2 2 1 1
0 0
B3 BS B3 BS 13-93 B3-Cu B3-Zn BS-Cu BS-Zn 13-93 B3-Cu B3-Zn BS-Cu BS-Zn A) B3-Cu-Zn BS-Cu-Zn B) B3-Cu-Zn BS-Cu-Zn
C)
Figure 31: Antibacterial effects of BGs: Summary of mean inhibition zones of different BGs against A) gram-positive bacteria S. aureus and B) gram-negative bacteria E. coli bacteria; One-way ANOVA statistical analysis denotes non-significant differences (*p<0.05); C) exemplary images of the inhibition E. coli and S. aureus bacteria growth by the different BGs
3.3.5 Dose-depending effect of BGs on dendritic cells
In this part of the study, only borate BGs doped with copper and/or zinc were analysed based on their previously promising results (section 3.3.1 - 3.3.4). The chosen borate BGs were tested in contact with DCs in order to evaluate the possible immune-modulating properties of the fabricated BGs. To verify the influence of the increased pH during the dissolution of the BGs, conditioned CCM with and without HEPES were prepared. The pH measurements, summarized in Table 5, showed that the addition of HEPES has, as expected, a great influence on the pH. However at high concentrations an increase of pH up to 7.8 was also 3 Borate and borosilicate glasses doped with copper and/or zinc 55 observed for CCM containing HEPES. Additionally, conditioned CCM without HEPES showed a great increase of pH and reached maximum values of 8.7 for 10% concentrations. Therefore, conditioned CCM containing HEPES was used for cell culture experiments and to verify if a possible high pH could negatively influence the cells, pH-CCM with a pH of 8.5 was additionally used as control.
Table 5: Summary of measured pH of prepared conditioned CCM with and without HEPES.
CCM Without HEPES With HEPES CCM 7.62 ± 0.01 7.15 ± 0.02 0.1% B3 8.10 ± 0.05 7.16 ± 0.01 1% B3 8.55 ± 0.03 7.32 ± 0.01 10% B3 8.69 ± 0.08 7.79 ± 0.05 0.1% B3-Cu 8.17 ± 0.07 7.13 ± 0.01 1% B3-Cu 8.43 ± 0.06 7.20 ± 0.01 10% B3-Cu 8.65 ± 0.04 7.52 ± 0.02 0.1% B3-Zn 8.34 ± 0.03 7.11 ± 0.01 1% B3-Zn 8.44 ± 0.04 7.16 ± 0.01 10% B3-Zn 8.67 ± 0.05 7.62 ± 0.01 0.1% B3-Cu-Zn 8.21 ± 0.01 7.24 ± 0.01 1% B3-Cu-Zn 8.52 ± 0.04 7.32 ± 0.01 10% B3-Cu-Zn 8.71 ± 0.01 7.63 ± 0.02
900 CCM B3 800 B3-Cu B3-Zn 700 B3-Cu-Zn 600
500
400 * * 300 * *
Concentration (mg/L) Concentration * 200 * 100
0 P Ca B Cu
Figure 32: Amounts of ions released in the 10%-conditioned CCM after 24h of immersion of the different fabricated BGs, measured by ICP-OES. One-way ANOVA statistical analysis denotes significant differences (*p<0.05) 56 3 Borate and borosilicate glasses doped with copper and/or zinc
Additionally to pH measurements, the concentrations of ions in the conditioned CCM medium were evaluated for the 10% concentrations. As shown in Figure 32, high amounts of B could be released and no significant differences between the different borate BGs could be found (in accordance to the results obtained in section 3.3.3). Moreover, a decrease of P was observed, which proves (as already found in section 3.3.3) that CaP-rich species are formed on top of the BG particles. More importantly, copper was released from B3-Cu and B3-Cu-Zn BG, up to 61.8 ± 3.7 mg/L and 54.8 ± 1.0 mg/L, respectively. In accordance to the results obtained in section 3.3.3, no release of Zn was detected.
In order to evaluate if ions released from BGs could have toxic effects, DCs were stimulated for 48h in the presence of the different conditioned CCM. After 48h, the DCs were matured using LPS overnight and then evaluated using a standard cell viability assay. As shown in Figure 33, no significant differences between CCM and pH-CCM were found. This proves that the cells are not sensitive to a pH increase up to 8.5. Moreover, no differences between B3 and B3-Zn BGs as well as between B3-Cu and B3-Cu-Zn BGs were found. This is in accordance with the ICP measurements, where the only difference between the BGs was the release of copper. Both, B3 and B3-Zn BGs showed no significant influence on DCs in 0.1% and 1% concentration, however in 10% concentration they led to a reduction of living DCs of around 20-25%. Contrary, the released copper from B3-Cu and B3-Cu-Zn BGs seems to increase the toxic effect of the borate BGs by showing already a reduction of living cells at 1% concentration.
0.1% dilution 1% dilution 10% dilution 100 +,-
80 +,-
+,- 60 +,- +,-
+,-
40 Living cells (%) cells Living
20
0 CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn
Figure 33: Calculated percentages of living DCs after cultiviation in presence of different conditioned CCM for 48h. Significant changes between conditioned CCM and CCM(+)/pH-CCM(-) are marked (p<0.01).
3 Borate and borosilicate glasses doped with copper and/or zinc 57
0.1% dilution 0.1% dilution 4000 1% dilution 2500 1% dilution CD11c 10% dilution +,- +,- MHCII 10% dilution 3500 +,- +,- 2000 3000
2500 1500
2000
MFI MFI 1500 1000
1000 500 +,- 500 +,-
0 0 CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn
7000 0.1% dilution 0.1% dilution 1% dilution 5000 1% dilution CD83 10% dilution CD80 10% dilution 6000
+,- 4000 5000
4000
3000 MFI
3000 MFI 2000 2000
1000 +,- 1000 +,- +,-
0 0 CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn
0.1% dilution 0.1% dilution 1% dilution 250 1% dilution 7000 CD25 10% dilution 10% dilution CD86 +,- 6000 200 +,- 5000
150 4000 +,-
MFI +,- MFI +,- 3000 100
2000 +,- 50 1000 +,- +,- 0 0 CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM pH-CCM B3 B3-Cu B3-Zn B3-Cu-Zn
Figure 34: Summary of measured surface markers of DCs cultivated in the presence of conditioned CCM containing dissolution products of the fabricated BGs. Significant changes between conditioned CCM and CCM(+)/pH-CCM(-) are marked (p<0.05).
To better evaluate the effects of BGs on the activation and phenotypical maturation of DCs, the expression of typical surface markers was analysed by flow cytometry. In accordance to the observation made based on Figure 33, a dose-dependent effect of the conditioned CCM on DCs was found. As shown in Figure 34, for all analysed surface markers (besides CD11c), a reduction was observed with increasing copper content (B3-Cu and B3-Cu-Zn). This is also in accordance with the reduction of living cells cultivated in contact with CCM containing dissolution products of copper-doped BGs. Of high interest, B3 as well as B3-Zn BGs led to the 58 3 Borate and borosilicate glasses doped with copper and/or zinc increase of specific surface markers (CD11c, MHCII, CD80 and CD86), depending on the concentration. These data suggest that DCs cultivated in the presence of B3/B3-Zn conditioned CCM were activated. In contrast, the presence of copper led to a reduced activation of the DCs.
0.1% dilution 0.1% dilution 1% dilution 1% dilution 10% dilution 10% dilution 120000 **** **** 100000 ** **** **** 100000 **** 80000 *** *** 80000
60000 cpm
60000 cpm
*** 40000 40000 **** **** 20000 20000
0 0 CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM B3 B3-Cu B3-Zn B3-Cu-Zn DC:Spleen cell ratio 1:40 DC:Spleen cell ratio 1:133
0.1% dilution 0.1% dilution 1% dilution 70000 1% dilution 35000 10% dilution 10% dilution * 60000 30000
50000 25000
40000 20000
cpm cpm 30000 15000
10000 20000
5000 10000
0 0 CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM B3 B3-Cu B3-Zn B3-Cu-Zn DC:Spleen cell ratio 1:333 DC:Spleen cell ratio 1:400
Figure 35: Proliferation of T-cells co-cultivated with DCs in the presence of conditioned CCM. Significant changes between conditioned CCM and CCM are marked with asteriks (*p<0.05, **p<0.01, ***p<0.001, ****p<0.0001).
By performing in vitro mixed leucocyte reactions, the DC-mediated allogeneic T-cell proliferation capacity can be studied [191]. T-cells were co-cultivated with DCs for 72h, which were matured before in the presence of the conditioned CCM. Since no significant influence between CCM and pH-CCM was found, for further experiments just CCM was used. The proliferation of the T-cells was then analysed by a proliferation assay and the results are shown in Figure 35. The effect of dissolution products of BGs on the DCs maturation also showed consequences on the DCs capacity to stimulate T-cells. The activated DCs showed an increased stimulation effect on T-cell proliferation, especially in 1% and 10%-CCM. Moreover, the inactivated DCs cultured in CCM containing copper (B3-Cu and B3-Cu-Zn) in high concentration (10%-CCM) showed an inhibition effect on the proliferation of T-cells. 3 Borate and borosilicate glasses doped with copper and/or zinc 59
Interestingly, the 1%-CCM containing dissolution products from B3-Cu and B3-Cu-Zn BGs showed, similar to conditioned CCM without copper, a stimulation effect on T-cell proliferation.
0.1% dilution 0.1% dilution 6000 1% dilution 1% dilution IFN-y 10% dilution 140 TNF-alpha 10% dilution * 5000
120
* 4000 100
80 3000
pg/ml pg/ml 60 2000
40
1000 20
0 0 CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM B3 B3-Cu B3-Zn B3-Cu-Zn
0.1% dilution 0.1% dilution 500 1% dilution 1% dilution 180 GM-CSF * 10% dilution IL-1alpha * 10% dilution 160 * 400 140
120 300
100 pg/ml
pg/ml 80 200 60
40 100
20
0 0 CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM B3 B3-Cu B3-Zn B3-Cu-Zn * 0.1% dilution 0.1% dilution 1% dilution 1% dilution 300 * 10% dilution 10% dilution IL-6 3000 IL-17A * * 250 * * 2500 * 200 2000
150
1500 pg/ml * pg/ml
100 1000 * 500 50
0 0 CCM B3 B3-Cu B3-Zn B3-Cu-Zn CCM B3 B3-Cu B3-Zn B3-Cu-Zn
Figure 36: Summary of secreted proinflammatory cytokines during DCs - T-cell co-cultures in a ratio 1:40 for 72h. Significant changes between conditioned CCM and CCM are marked with asteriks (*p<0.05).
In addition, to analyse the effect of DCs (cultured in conditioned medium) on T-cells, the secretion of proinflammatory cytokines in the remaining CCM after co-cultivation of T-cells and DCs was evaluated (Figure 36). The dose-dependent effect of B3 and B3-Zn was 60 3 Borate and borosilicate glasses doped with copper and/or zinc also found. In accordance with the increased cell proliferation of T-cells, 10%-CCM containing dissolution products of B3 and B3-Zn BGs led to an increased secretion of TNF-alpha, GM-CSF, IL-17A and IL-6. More interestingly, similar to the results obtained by measuring the proliferation of T-cells, conditioned CCM containing copper led, in high concentrations, to a reduction of cytokines inside the CCM, whereas lower concentrations led to increased levels of TNF-alpha, IL-1aplpha, IL-17A and IL-6. These findings prove that there is a composition and dose-depending effect of BGs on the immune reaction shown here by the reaction, of DCs and T-cells.
3.4 Discussion
3.4.1 Influence of network former on BG characteristics
As described in section 2.2, the most important properties of BGs, such as the dissolution rate and the bioactivity, are defined by their chemical composition. The successful fabrication of the silicate, borosilicate and borate BGs (without addition of copper and zinc) was proven by XRD (indicating an amorphous structure) and FTIR (presence of Si and/or B within the glass network). In addition, the structure of the silicate 13-93 and the series of borate B3 BGs containing copper and/or zinc was studied by NMR measurements. NMR measurements showed that the 13-93 BG is composed of mainly Q2Si units and Q3Si units. In addition, the main amount of phosphorous is present in the form of orthophosphates [216]. Although no (comparable) results for the structure of 13-93 BG are available in the literature, similar results were obtained for the silicate 45S5 BG composition. In accordance to 13-93 BG, also 45S5 BG is composed of mostly orthophosphates, Q2Si and Q3Si units [217]–[219]. In contrast to silicate BGs, the structure of borate BGs is more complex. Due to the so-called borate anomaly (described in section 2.2.1), it is not clear whether a high amount of [BO3] units indicates a high polymerization or a low polymerization degree [70], [71], [73]. By NMR measurements, the borate B3 BG was found to contain boron in the form of [BO4] units and
[BO3] units. Moreover, also in the borate B3 BG, the main amount of phosphorus is present in the form of orthophosphates.
Since the amount of network modifiers was not changed, the difference within the thermal, dissolution and antibacterial properties of the silicate, borosilicate and borate BG can be attributed to the change of network former. Obvious differences, such as the decrease of the glass transition and crystallization temperatures of the different BGs with increasing 3 Borate and borosilicate glasses doped with copper and/or zinc 61 boron content are well known [62], [79], [82], [220]. In addition, also the lower chemical durability and the faster dissolution and apatite formation rate of boron-containing/borate BGs compared to silicate BGs is well documented [56], [79], [82], [221]. This could be due to the decreasing resistance against hydrolysis with increasing boron content, which can be explained by the initial formation of Si-O-B bonds followed by the formation of B-O-B bonds. Si-O-Si bonds showed the highest resistance against hydrolysis, followed by Si-O-B and the lowest resistance was found for B-O-B bonds [222]. The results of the dissolution tests obtained in this thesis indicate that the addition of boron to the silicate BG leads to the disruption of the glass network and therefore to a faster dissolution rate. By calculating the relative amount of Si released from the 13-93 and the BS BGs, an increased dissolution rate with increasing boron content can be observed. Whereas around 6% of the total amount of silicon was released from the 13-93 BG after being immersed in TRIS for up to 96h, around 14% of silicon was released from the BS BG. This effect was not that obvious during dissolution studies in SBF, where around 5% and 7% of the total incorporated amount Si were released from the 13-93 and BS BGs, respectively. However, during dynamic dissolution experiments using SBF, again a clear difference between the relative released amount of silicon from 13-93 (around 6%) and BS (around 18%) BGs in the first 3h of dissolution time could be observed. Therefore, the results of the dissolution studies support the claim that boron leads to the lower network stability of the borosilicate BG in comparison to the silicate BG and therefore increases the release rate of Si.
In order to understand the effect of adding boron to the silicate BG, further detailed analyses are necessary. Especially, the structure of borosilicate BGs needs to be evaluated, since the influence of associations between Si and B (Si-O-P linkages) might also have an great influence on the glass structure and therefore on the dissolution process [223]. In addition, the dissolution of BGs is a complicated process and is influenced by several mechanisms. This could be shown by the choice of dissolution medium. For instance, obvious differences between the dissolution of borate and borosilicate BGs in TRIS and SBF could be measured, whereas the dissolution behaviour of silicate BGs seems not to be influenced whether TRIS or SBF was used. This influence can be mainly explained by the in section 2.2.2 described differences in the apatite formation ability of the three kinds of glasses. In case of borate and borosilicate BGs, the rate-limiting step of dissolution was found to be the ion diffusion to the interphase [79]. Due to the high amount of Ca and P in the SBF solution, apatite formation might be faster and therefore the formed apatite layer might retard the further dissolution of the borate and borosilicate BGs, as already reported in literature [224]. 62 3 Borate and borosilicate glasses doped with copper and/or zinc
In addition to the differences found due to excess/limit of ions in the dissolution media, also the pH seems to play an important role. In literature [225], it could be already shown that the dissolution behaviour strongly changes due to changing of the pH from a neutral to an acidic environment. In this study, it was found that Si was almost not dissolved at all in an acidic environment. In contrast, the remaining elements (e.g. Na, Ca, P) were released in a fast rate [225]. Similar observations were made here. Whereas the B3 and BS BG showed a fast overall dissolution rate (besides the release of Si), Si from the 13-93 BG was not dissolved, but at the same time the network modifiers Mg, Na, K and Ca were fast released. This might be explained by the fact that lactic acid is a weak acid and can act as a buffer [225]. During the dissolution of the different glasses, more H+ might be consumed leading to faster release of the modifiers. The retarded release of Si in borosilicate BGs was also inferred by measuring the pH during the dissolution. Only a minor increase of pH was measured for 13-93 BG indicating the slow dissolution of this BG. In contrast, B3 BG showed an initial fast increase of pH in the first 24h, indicating the fast dissolution in this time frame. In case of BS BG, the pH change is characterized by a fast increase (probably due to release of the “borate part”), followed by a minor decrease (probably to the “silicate part”). These differences during dissolution in an acidic environment might also explain the differences found in the antibacterial effect of the glasses. The efficiency of BGs against E.coli and more obviously against S. aureus was found to be higher with increasing boron content. One reason might be that during the antibacterial test, an acidic environment occurred and therefore the boron containing BGs were able to release ions and to increase the pH more efficiently than the silicate BGs. Overall, these results showed the great importance of using different dissolution media in order to analyse the different dissolution behaviours of different BGs under a variety of conditions.
The differences found between the fabricated BGs by structural and thermal analysis as well as by dissolution and antibacterial tests are summarized in Table 6. Although the results indicate that by adding boron in different amounts, several key properties of the resulting BGs can be tuned, the results also show the importance of clearly investigating the effects that the addition boron exert on the properties on the BGs.
3 Borate and borosilicate glasses doped with copper and/or zinc 63
Table 6: Summary of properties of silicate 13-93, borosilicate BS and borate B3 BG obtained by structural and thermal analysis, dissolution experiments as well as antibacterial tests.
13-93 BG BS BG B3 BG Mainly composed of Mainly composed of Structural - Q2Si and Q3Si units [BO3] and [BO4] units Intermediate glass Highest glass transition Lowest glass transition transition and Thermal and crystallization and crystallization crystallization temperature temperature temperature Intermediate Lowest dissolution Highest dissolution dissolution rate, partly Dissolution rate, not dissolvable in rate, fast dissolvable in dissolvable in an acidic an acidic environment an acidic environment environment Moderate efficiency Low efficiency against Moderate efficient against E. coli and good Antibacterial E. coli, not efficient against E. coli and S. efficiency against S. against S. aureus aureus aureus
3.4.2 Influence of incorporation of copper/zinc on BG characteristics
In general, the addition of network modifiers leads to the turning of bridging oxygens into non-bridging oxygens, which in turn changes the BG structure [54], [226]. Although in this work, the total amount of network modifier was not changed, the charge-to-size ratio of the copper, zinc and calcium ions plays a pivotal role in determining the glass structure [227], [228] and therefore it can directly affect the resulting properties of the fabricated ion-doped BGs. Here, calcium was replaced either by copper or zinc or by both of them. Although the charge-to-size ratio of copper and zinc is relatively similar, the charge-to-size ratio of calcium is different. Whereas the atomic ratio is larger for calcium, the atomic weight is higher for copper and zinc [229], [230]. Therefore, it is expected that the replacement of calcium by copper and/or zinc has a pronounced effect on the structure of the resulting BGs, resulting in different dissolution rates and bioactivities. By measuring the thermal properties, such effects can be assessed. For example, ion-doped borate and borosilicate BGs showed lower glass transition temperatures than neat borate and borosilicate BGs. Interestingly, the addition of copper and/or zinc to borate BGs led to an increased crystallization peak, whereas the 64 3 Borate and borosilicate glasses doped with copper and/or zinc addition of copper and/or zinc to borosilicate BGs led to a decreased crystallization peak in comparison to pure borosilicate BG. Based on the characterization of the ion-doped borate and borosilicate BGs and the results obtained by dissolution tests in SBF, TRIS and LA, antibacterial and cell culture experiments, some observations could be made.
i) Influence of copper and/or zinc on the structure and dissolution rate of ion- doped borate and borosilicate BGs
The NMR measurements on BGs in the series of borate BGs indicated an increased amount of phosphorus due to the addition of copper and/or zinc in the form of Si-O-P bonds or as pyrophosphate (P2O7)4-), which might lead to an increased network stability. In accordance with this observation, zinc inhibited the release of Si during dissolution of the zinc-containing borosilicate BGs in TRIS (from around 40 mg/L to 32 mg/L). Surprisingly, the addition of copper seems to favour the release of Si (from 40 mg/L to 55 mg/L). However, in case of the network modifiers K and Mg, a clear decrease of dissolution rates by the addition of copper and zinc was observed for borosilicate BGs, with a slightly greater decrease for zinc- compared to copper-doped BS BGs. This more pronounced effect of zinc could not be observed for ion-doped borate BGs, which might be explained by the effect, that copper and/or zinc have on the coordination of boron ions within the borate BGs. Since the binding energy of
[BO4] groups is lower than that of [BO3] groups and an increase of [BO3] groups was indicated by NMR measurements, it is expected that the ion-doped borate (and borosilicate) BGs are more stable and therefore the dissolution rates decrease. During the dissolution of the series of borate and borosilicate BGs in TRIS, no significant influence on the boron release rate was found. However, also here by analysing the release rates of the network modifiers K, Mg and Na, a decreased release rate of these ions was found by the addition of copper and/or zinc to the borate BGs. This effect was more obvious for copper.
Besides having a significant influence on the structure and resulting dissolution behaviour, the addition of copper and/or zinc also has an impressive effect on the antibacterial efficiency as well as on the cell behaviour:
ii) Antibacterial efficiency of borate and borosilicate BGs by adding copper and/or zinc
It is well known that the antibacterial efficiency of different BGs can be either attributed to the change of pH and/or to the release of ions [86]. Dissolution experiments using different media showed that addition of copper and/or zinc to borate and borosilicate BGs did not have a significant influence on the pH. The changed antibacterial efficiency of the 3 Borate and borosilicate glasses doped with copper and/or zinc 65 present ion-doped BGs can be therefore mainly attributed to the release of copper and/or zinc. In both cases, ion-doped borate and borosilicate BGs, the antibacterial efficiency against E. coli was found to be greater for copper containing BGs than for zinc containing BGs. This result might be explained by the well-known efficiency of copper against bacteria [75], [94] and additionally by the greater amount of copper incorporated in comparison to zinc. However, this more pronounced effect of copper could not be observed in case of S. aureus, which might be due to the overall greater effect that the ion-doped BGs exert on S. aureus. This might be explained by the highly organized compact structure of the membrane of E. coli, which protects the bacteria cell better against antibacterial ions [215]. In any case, the addition of copper and/or zinc in borate and borosilicate BGs was shown to efficiently improve the antibacterial effect of the BGs. Based on dissolution experiments in TRIS and SBF, no significant release of zinc was expected and therefore the efficiency of zinc as antibacterial ion was surprising. However, during the agar diffusion test, an acidic environment might occur, which (as shown in dissolution experiments using LA) favours the release of zinc ions. Moreover, also the presence of zinc on the surface of the BG pellets already affect the bacterial growth, as shown in the literature [231]–[233].
iii) Effect of addition of copper and/or zinc to borate BGs on immune cells
Based on the significant effects that BGs (undoped and doped with Cu and/or Zn) have been shown to have on the regeneration of soft and hard tissues, several cell studies have been carried out in the past using relevant cells, e.g. bone cells, skin cells and bloods cells [29], [31], [190], [233]–[239]. However, here we aimed to investigate the effect that ion-doped borate BGs have on the immune system (in the form of DCs and T-cells), since the immune system is strongly involved in any regeneration process and DCs are known to interact and infiltrate biomaterials [240]–[243]. Whereas pure B3 BG as well zinc-doped B3-BG (B3-Zn) showed only in high concentration toxic effects on DCs, the addition of copper seems to enhance the toxic effect of the B3-Cu and B3-Cu-Zn BGs. Based on the obtained results of the ICP measurements, we could calculate a critical biological level of copper of below 5 mg/L. This level is lower compared to the toxicity level for fibroblasts reported by Wang et al. (10 mg/L) [244]. Therefore, it seems that DCs are more sensitive to copper ions than fibroblasts. Besides toxicity, by measuring the amount of surface markers, the activation or inactivation of the DCs by the ion-doped B3 BGs was also analysed. Interestingly, B3 and B3-Zn BG led to the activation of the DCs, whereas B3-Cu and B3-Cu-Zn led to the inactivation in 1% concentration in cell culture medium. Since only activated (differentiated) DCs induce an immune reaction and inactivated DCs promote rather an immune tolerance, our results clearly show that the immune system can be affected by the addition of metallic ions [245]. This effect was also 66 3 Borate and borosilicate glasses doped with copper and/or zinc further proven by the ability of the DCs, in contact with the different BGs, to stimulate T-cell proliferation, indicating that more studies are necessary to clearly understand the effect that ion-doped BGs exhibit on the immune system, this should become an area of intensive research in future.
In summary, significant effects of adding copper and/or zinc on the structure and resulting properties of borosilicate and borate BGs were identified. Besides the important characteristics of BGs, the dissolution rate and bioactivity, especially the biological properties were affected. A summary of the effects discovered in this part of the study is given in Table 7.
Table 7: Summary of the effects, that the addition of copper or zinc exhibit on the properties of borate and borosilicate BGs obtained by structural and thermal analysis, dissolution experiments, antibacterial tests as well as cell tests using immune cells.
Addition of copper Addition of zinc Structural Increases BG stability Increases BG stability Lowers glass transition Lowers glass transition temperature temperature Thermal Affects the crystallization Affects the crystallization temperature depending on BG temperature depending on BG composition composition Leads to a slower dissolution rate, Leads to a slower dissolution rate, especially in the case of especially in the case of borate borosilicate BGs Dissolution BGs Zinc was incorporated in apatite Release of copper could be proven formation, could be just released in several dissolution media in lactic acid Increases antibacterial efficiency Shows also great efficiency Antibacterial Increases antibacterial efficiency against the “more-protected” E. coli Leads to dose-dependent effect on Leads to dose-dependent effect on immune cells immune cells Immune cells Critical biological level of 5 mg/L No obvious difference compared Is able to inactivate DCs, could to undoped B3 BG could be promote immune tolerance measured 4 Fabrication of bioactive glass containing wound dressings by freeze drying 67
4 Fabrication of bioactive glass containing wound dressings by freeze drying8
4.1 Introduction
As described in section 2.3, freeze dried foams containing therapeutic agents (e.g. BGs) are suitable candidates for wound dressings exploiting their ability to constantly release incorporated therapeutic agents, which should accelerate wound healing and vascularization. By combining the fabricated BGs described in Chapter 3 with MC and MH (sections 2.4 and 2.5), in this part of the study multifunctional foams were developed and characterized for wound dressing applications. This Chapter presents the processing and full characterization of composite MC-MH foams containing BG particles (B3 or B3-Cu).
4.2 Experimental procedure
4.2.1 Production of foams
Prior to freeze drying, 2.5 wt.% MC (viscosity 4,000 cP, Sigma Aldrich) was dissolved in distilled water (DW) at 40°C. For foams containing MH (MGO >800 mg/kg, Watson & Son, New Zealand) and BG particles, the solution was cooled down to room temperature, then 2.5 wt.% MH was added and stirred for 15 min. For foams containing additionally BG, 0.75 wt.% of B3 or B3-Cu was added (15 wt.% of BG compared to MC-MH weight) and stirred for another 15 mins. The BGs were fabricated as described in section 3.2.1, crushed using a jaw crusher (Retsch) to a particle size below 0.1 µm and then grounded using a planetary mill (zirconia, Retsch). To obtain BG powders with an average BG particle size between 5 and 20 µm (proved by SEM images), milling balls with a diameter of 5 mm for 12 runs á 5 min, followed by milling balls with a diameter of 1 mm for 10 runs á 2 min (both made out of zirconia) were used. The four different (MC, MC-MH, MC-MH-B3 and MC-MH-B3-Cu) solutions were filled in a 48-well-plate, frozen at -20 °C overnight and then freeze dried (-80 °C, 0.01mbar, Alpha 2-4
8 Some of the results presented in this chapter were previously published by the authors (as first author) and are hereby adapted from reference [246] 68 4 Fabrication of bioactive glass containing wound dressings by freeze drying
LSC plus, Christ). After drying, the samples were removed from the well-plate and cylindrical samples of 9 mm diameter and 12 mm height were obtained [246].
4.2.2 Characterization
4.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX)
SEM (Leo 1530, Oberkochen, Germany) combined with energy-dispersive X-ray analysis (EDX) (Thermo Scientific UltraDry, Madison, WI) was used to examine the morphology of different fabricated foams and to evaluate the distribution of the BG particles inside the foams. The SEM was operated at 1 kV and working distance of 2 mm. Foams were fixed on a sample holder by carbon tape.
4.2.2.2 Porosity
The porosity of the produced freeze-dried foams was calculated based on the following formula:
휌 푃 = 1 − 푓표푎푚 (2) 휌푚푎푡푒푟푖푎푙
where 휌푓표푎푚 is the density of the fabricated foams and 휌푚푎푡푒푟푖푎푙 is the density of the respective used material. The mean density of the foam (휌푓표푎푚) was calculated by dividing the mass (measured by an analytical balance BM-252, A&D) through the volume of the 10 different foams of each kind. The density of the material (휌푚푎푡푒푟푖푎푙) was measured based on the Archimedes’ method (using an analytical balance BM-252, A&D) using films made by drying the different solutions made for freeze drying in a petri dish at room temperature (10 pieces á 1 cm2 were used).
4.2.2.3 Fourier transform infrared spectroscopy FTIR
FTIR (IR Affinity-1S, Shimadzu) was performed on composite foams before and after immersion in simulated body fluid (SBF), produced as described in section 3.2.2.7 [186]. The conditions for this analysis were as follow: spectral range 4000 to 400 cm-1; window material: CsI; 40 scans and resolution: 4 cm-1. 4 Fabrication of bioactive glass containing wound dressings by freeze drying 69
4.2.2.4 Mechanical testing
By compression tests using a universal testing machine (Instron 5960, Germany), the compressive strength of the cylindrical foams (9 mm diameter and 12 mm height) were assessed. At least 10 replicates were tested using a speed of 5 mm/min and a load cell of 100N. Since during the compression test, the foams get densified and no compression strength at break could be obtained, the mean compression strength was calculated based on the measured compression stress from 40-60% of compression strain (as shown exemplary in Figure 37) [247].
300
250
200
150
100
Compression stress (kPa) stress Compression 50
0
0 20 40 60 80 100 Compression strain (%)
Figure 37: Exemplary stress-strain-curve of a MC foam showing the densification and the resulting increasing compression stress with ongoing compression strain. Based on this, the mean compressive strength was calculated by calculating the average of the measured compression stress from 40-60% of compression strain (marked in grey)
4.2.2.5 Contact angle measurement
In order to evaluate the wettability of the foams, the contact angle can be measured by depositing a water droplet on a surface. The contact angle is then calculated by establishing a tangent of the foam surface and the liquid drop [248]. Here, the contact angle (DSA 30, Krüss) on at least 5 replicates was assessed using a 3 µl water drop.
4.2.2.6 X-ray diffraction (XRD) analysis
After SBF dissolution tests, the phase compositions of the freeze-dried foams were characterized by powder X-ray diffraction (XRD) measurements (Miniflex 600 HR, Rigaku, 70 4 Fabrication of bioactive glass containing wound dressings by freeze drying
Japan). Data were collected with a step size of 0.02ᵒover a 2θ range from 20ᵒ to 60ᵒ. Prior analysis, foams were dried at room temperature and measurements were carried out in triplicates.
4.2.3 In vitro dissolution in SBF
By using the same SBF and protocol as described in section 3.2.2.7, the acellular bioactivity of the foams was assessed. The samples were immersed in a ratio of 1.5 g of sample to 1L of SBF [187]. After different immersion durations (3, 7 and 14 days) in SBF, the samples were collected for analysis by FTIR, XRD and SEM. Additionally, the pH of the remaining SBF was recorded using a pH meter (FiveEasy Plus pH meter FP20) and then the released amount of MH was measured using a UV-Vis spectrophotometer (Specord 40, Analytik Jena, Germany) at 300 nm based on a calibration curve (R2=0.9989). For all measurements, 1 ml of solution was used and the calibration curve was measured based on 7 standards in the range of 100mg/ml to 1mg/ml. All samples and measurements were performed in triplicates.
4.2.4 Antibacterial efficiency of the foams
The antibacterial efficiency of the fabricated foams was evaluated using S. aureus and E. coli following a previously established protocol [249]. After 24h of incubating the bacteria strains in lysogeny broth medium (#968.1, Roth) at 37°C in an orbital shaker, the optical density (600nm, Genesys 30 TM, Thermo Scientific) of the bacteria solution was adjusted to 0.015. Then, 10 µl of the arranged bacteria solution was added to 1ml of the CCM containing a freeze-dried sample (same ratio as described for cell experiments). Then, the solutions containing the samples were incubated for 24h at 37 °C in an orbital shaker and subsequently the optical density was recorded. CCM without sample and CCM containing different amounts of MH were taken as control. All measurements were performed in triplicates. Based on the optical density measurements, the relative bacterial viability (%) was calculated using the following equation:
푂퐷 푅푒푙푎푡𝑖푣푒 푏푎푐푡푒푟𝑖푎푙 푣𝑖푎푏𝑖푙𝑖푡푦 (%) = 푠푎푚푝푙푒 × 100 (3) 푂퐷퐶퐶푀 4 Fabrication of bioactive glass containing wound dressings by freeze drying 71
4.2.5 In vitro cell tests
For cell tests, the foams were cut using a razor blade (Herkenrath Solingen) in pieces of 9 mm diameter and 2 mm height. Disinfection was performed for 1h using UV light. Conditioned media was prepared as described in section 3.2.4. Briefly, freeze dried foams were immersed in CCM (Dulbecco’s modified eagle medium (DMEM) supplemented with 10% fetal bovine serum and 1% penicillin and streptomycin) in a ratio of 0.01g/1ml CCM. Foams were, similar to all cell tests, incubated in a humidified atmosphere of 95% air and 5% CO2 at 37 °C. After incubation for 24h, the conditioned media was collected and analysed using ICP-OES (ICP-OES, Vista MPX, Varian) in order to measure the amount of ions in the conditioned media. To analyse the conditioned media using ICP-OES, the conditioned medium was stabilized using 1M nitric acid to reach a pH of 2. Samples and measurements were performed in triplicates.
4.2.5.1 Indirect cytotoxic test using mouse embryotic fibroblast (MEF) cells
As a preliminary test, an indirect cell test using MEF cells was conducted by seeding 100,000 MEF cells in 1ml CCM in 24-well-plates. Then the samples, which were fixed in TC- inserts (Sarstedt), were placed on top without touching the cells on the well plate bottom. The cells were then incubated as described above for 24h. As positive reference, cells without any samples and as negative reference, cells incubated in CCM containing 6% of Dimethyl sulfoxide (DMSO) were taken. For each kind of sample and control, four replicates were used. After 24h incubation, the viability of the MEF cells was evaluated by a mitochondrial activity calorimetric assay, which is based on tetrazolium salt (WST-8 assay, Sigma-Aldrich). Additionally, the morphology of the MEF cells was evaluated using staining with rhodamine phalloidin and 4′,6-diamidino-2-phenylindole (DAPI) [147], [250]. Whereas rhodamine phalloidin is staining actin cytoskeleton in cells, DAPI is binding to the Adenine and Thymine bases within the DNA of the cell nuclei and therefore staining the cell nuclei. Images were taken with a fluorescence microscope (Axio Observer D1).
4.2.5.2 Direct cell test using human dermal fibroblast (hDF)
To further prove the cytocompatibility of the samples, foams were tested in direct contact with hDF cells. Therefore, 50,000 hDFs were directly seeded on top of each sample using a 50 µl CCM drop. Then, the samples (located in a 24-well-plate) were incubated for 15 min and subsequently filled with 1ml CCM. Foams with cells seeded on the surface were then incubated for 24h and 7d. Following this, the viability of hDFs was assessed as described above. Additionally, the morphology of the hDF cells grown on the samples was evaluated by 72 4 Fabrication of bioactive glass containing wound dressings by freeze drying
SEM analysis after fixing the samples as described in literature [250], [251]. For SEM analysis, foams were washed with PBS and then fixed with fixative I (0.2 M sodium cacodylate trihydrate, 0.1 wt.% glutaraldehyde, 2 wt.% paraformaldehyde and 5 wt.% sucrose) and fixative II (0.2 M sodium cacodylate trihydrate, 0.3 wt.% glutaraldehyde and 2 wt.%) for 15 min, respectively. After fixation, scaffolds were dehydrated in a graded ethanol series (30, 50, 80, 90 and 95.8 vol.%) and finally dried in a critical point dryer (EM CPD300, Leica, DE).
4.2.5.3 In vitro scratch test using mouse embryotic fibroblast (MEF) and human keratinocytes-like (Hacat) cells
Based on a protocol established by Liang et al. [252], an in vitro scratch tests was done using MEF and Hacat cells in order to analyse the migration of the different cells. In addition to MEF cells, Hacat cells were used since keratinocytes are involved in wound closure as described in section 2.1.2 [253]. Conditioned CCM used for this test was produced as described in section 4.2.5. First, 100,000 MEF cells or 500,000 Hacat cells per 1ml CCM were seeded in a 24-well plate and incubated for 24h as described above. Then, by using a 200 µl pipette tip, a scratch was manually created on the grown cell (MEF or Hacat) monolayer. Subsequently, the cells were washed using PBS and then incubated with conditioned CCM for 24h. Again, a positive control (pure CCM) and a negative control (CCM supplemented with 6% DMSO) were used. The cell migration (closure of wound) was observed during the incubation after different time points (2.5h, 5h, 8h and 24h) by using a light microscope (Primo Vert, Carl Zeiss). All measurements were done in triplicates and the width of the scratch was calculated using ImageJ software [254].
4.2.5.4 Co-culture of human dermal fibroblasts (hDFs) and human umbilical vein endothelial cells (HUVECs)
In order to co-culture hDFs and HUVECs, first 70,000 hDFs per ml of CCM (as described above) were seeded per well in a 24-well-plate and incubated for 24h as described above. Then, 35,000 HUVECs per ml of H-CCM (CCM as described above plus endothelial cell growth supplements/Heparin (Promocell)) were seeded on top of the cultured 70,000 hDF cells, after removing the CCM from the hDFs. After another incubation time of 24h, the H-CCM was replaced by conditioned CCM. As positive control, fresh H-CCM and as negative control, H- CCM containing 6% DMSO was used. Additionally, for comparison just 70,000 of hDF cells or just 35,000 of HUVECs were cultured using the same protocol. After 24h of cultivation, the cells (hDFs, HUVECs and hDFs-HUVECs) were analysed using the cell viability test, SEM analysis and fluorescence staining, as described above. All measurements were done in triplicates. 4 Fabrication of bioactive glass containing wound dressings by freeze drying 73
4.3 Results
4.3.1 Crosslinking efficiency of MC with MH
As shown in Figure 38, the FTIR spectra of MC shows the typical bands reported [255]–[257]: at 3440 cm-1 the stretching vibrations of hydroxyl groups, at 1150 cm-1 the stretching vibrations of the C-O oxygen bridge, at 1050 cm-1 the C-O-C stretching mode of the glycosidic unit and at 940 cm-1 the vibration of the OCH3 group. Moreover, in the range 2925- 2830 cm-1 peaks representing the stretching vibrations of C-H can be found [255]–[257]. In addition, the FTIR of MH shows also the reported typical bands, e.g. related to sugar (the main component) in the range of 1500-750 cm-1 and the OH stretching vibrations at 3250 cm-1 [258], [259].
C-O-C C-H C-O
O-H C-H MC-MH
MC
MH
4000 3500 3000 2500 2000 1500 1000 500 0 Wavenumber (cm-1)
Figure 38: FTIR spectra of MC, MH and MC-MH showing the typical bands reported for MC and MH [255]–[257]
Besides the fact that both typical bands known for MC and MH can be found in MC-MH foams, the crosslinking reaction can be observed in the presence of hydroxyl groups. Higher wavenumbers in the OH region (around 3575 cm-1, Figure 39 A) are attributed to the vibrations of “free” OH groups, which can be detected for MC. However, these bands are not detectable anymore for MC-MH foams, suggesting that this free OH groups are getting connected to the MH. Moreover, it is possible to distinguish between the three components of the OH bands found in MC-MH foams (Figure 39 B). The first two bands at around 3430 cm-1 and 3340 cm-1 can be assigned to intramolecular hydrogen bonded OH groups, whereas the 74 4 Fabrication of bioactive glass containing wound dressings by freeze drying third band centered at 3228 cm-1 can be assigned to the intermolecular hydrogen bonded OH groups. Especially the intermolecular hydrogens bonded hydroxyl groups seem to be increased in the MC-MH foam compared to the MC foam, which proves that a crosslinking reaction occurs between MC and MH.
MH MC-MH MC
3800 3600 3400 3200 3000 2800 A) wavenumber (cm-1) B)
Figure 39: A) FTIR spectra of MC, MH and MC-MH in the range of 3700-3000 cm-1 and B) zoom on components of the OH vibrations band for the crosslinked MC-MH foam
4.3.2 Characterization of the foams
After verifying the crosslinking reaction, the produced foams were additionally loaded with BG particles, namely B3 or B3-Cu. Since the addition of BG particles could have a significant influence on the pore morphology and overall foam porosity, the produced foams were evaluated using SEM. As shown in Figure 40, all freeze dried foams exhibit a porous structure. However, a difference could be found between MC/MC-MH and MC-MH foams containing BG. Whereas foams without BG seem to have a homogenous porosity and smooth surfaces, the foams containing BG seem to be less homogenous in the pore distribution and denser. Moreover, the surfaces of the foams are rougher. Additionally, based on the SEM images shown in Figure 40, a pore size range of around 50-200 µm could be found for all fabricated foams (measured by the ImageJ software). Based on density measurements, the porosity of the foams could be calculated. As summarized in Table 8, only a minor impact on the porosity by the addition of BG particles could be found. In addition to SEM images, EDX was performed in order to confirm the presence of BG in the foams. In both cases, MC-MH-B3 and MC-MH-B3-Cu, the BG related elements could be found. 4 Fabrication of bioactive glass containing wound dressings by freeze drying 75
Table 8: Density and porosity of the produced foams
Foam MC MC-MH MC-MH-B3 MC-MH-B3-Cu Density in g/cm3 0.12 ± 0.02 0.27 ± 0.03 0.34 ± 0.02 0.45 ± 0.01 Porosity in % 93.4 ± 0.5 93.6 ± 0.4 95.2 ± 0.4 96.8 ± 0.1
Figure 40: SEM images showing the morphology of the fabricated freeze-dried foams observed by SEM 76 4 Fabrication of bioactive glass containing wound dressings by freeze drying
By performing EDX analysis, the presence as well as the distribution of the BG particles inside the MC-MH foams could be evaluated. It is assumed that the total amount of B3/B3-Cu is not changing during the freeze-drying process, therefore MC-MH-B3 and MC-MH-B3-Cu both contain 15 wt.% of the respective BG. As shown in Figure 41, the presence of elements (B, Na, Mg, P, K and Ca) related to B3/B3-Cu could be found. Moreover, in case of B3-Cu, also copper could be detected on the surface of the composite foam. Based on several EDX measurements at different spots, the homogenous distribution of the BG particles inside the MC-MH foam could be also proved.
80 21 C MC-MH-B3 C MC-MH-B3-Cu 70 18 O O 60 15 50 12
40
cps/eV cps/eV 9 30
6 20
3 K Ca 10 B B Na Mg P Na K Ca Mg P Cu 0 0 0 2 4 6 8 10 0 2 4 6 8 10 EDX spectra EDX spectra
Figure 41: EDX spectra of the composite MC-MH foams containing B3 or B3-Cu with corresponding SEM images showing the distribution of the measured elements.
In addition to the morphology and porosity, the contact angles as well as the mechanical properties were investigated. As shown in Table 9, the introduction of MH into the MC foams led to a reduction of the contact angle. The contact angle of foams containing additionally BG was not measurable, since the water drop was directly absorbed. Therefore, the BG containing foams can be described as superhydrophilic (Figure 42) [260].
4 Fabrication of bioactive glass containing wound dressings by freeze drying 77
Figure 42: Water drop on top of different of foams during contact angle measurement. MC-MH-B3 and MC-MH-B3-Cu directly absorbed the water drop, indicating a superhydrophilic behavior [260]
Compression tests, similar to contact angle measurements, showed that, by adding BG particles to the foams, the resulting properties of the MC-MH foams can be improved. According to Table 9, the compressive strength was not significantly changed by the addition of MH, however, by the addition of B3/B3-Cu, the foams were significantly strengthened. Additionally, after compression tests, the foams were still intact. In case of MC-MH and MC-MH foams containing BG, the foams were even able to recover their shape to a certain point (Figure 43) after a long time (a few hours to days).
Table 9: Summarized contact angle and mean compression strength results (at 40-60% of compression strain) of tests on the fabricated foams
Foam MC MC-MH MC-MH-B3 MC-MH-B3-Cu Contact angle [°] 101 ± 7 77 ± 9 n. d. n. d. Mean compression 17 ± 3 13 ± 4 43 ± 6 40 ± 6 strength [kPa] n. d. = not detectable 78 4 Fabrication of bioactive glass containing wound dressings by freeze drying
Figure 43: Images of foams directly after compression tests (left) and after 72h (right)
4.3.3 Acellular bioactivity using SBF
By performing dissolution studies using SBF, the bioactivity of materials can be evaluated [186]. Therefore, neat and composite MC foams were immersed in SBF for up to 14 days and evaluated using SEM, FTIR and XRD. Moreover, the pH of the remaining SBF was recorded. As shown in Figure 44, in case of neat MC and MC-MH foams, the pH did not significantly changed during immersion time. In contrast, MC-MH foams containing BG showed an increase of pH, which is in accordance to the pH increase found in section 3.3.3. The increase of pH is a clear sign of dissolution products from the BG inclusions and indicates also the formation of an apatite layer on the sample surfaces, as described in section 2.2.2.
MC MC-MH 8.0 MC-MH-B3 MC-MH-B3-Cu
7.9
7.8
7.7 pH
7.6
7.5
7.4
2 4 6 8 10 12 14 immersion time (days)
Figure 44: Change of pH as a function of immersion time of neat and composite foams in SBF 4 Fabrication of bioactive glass containing wound dressings by freeze drying 79
By SEM, the surfaces of the remaining foams were further investigated. As shown in Figure 45, MC and MC-MH foams did not show any formation of a CaP-rich phase, as expected.
Figure 45: SEM images of MC and MC-MH foams after immersion in SBF for 14 days
In order to further analyse the MC and MC-MH foams after immersion in SBF, FTIR measurements were done. As shown in Figure 46 A), no obvious difference between MC foams before and after immersion in SBF could be found. However, MC-MH foams seem to release MH during immersion in SBF, since the measured FTIR spectrum of MC-MH foams after immersion in SBF is similar to pure MC-foams. This observation could be further proven by UV-Vis measurements. Figure 46 B) shows clearly the release of MH from the MC-MH foams (as well as from BG containing MC-MH foams). However, due to the complexity of MH, it is not possible to clearly analyse the released products via UV-Vis. MH is composed of several compounds, having different UV-Vis maxima and different degradation rates [261]. Therefore, the amount of MH released could be actually higher or lower than the measured one.
MC-MH 4.0 MC 3.5 after 14d MC-MH-B3C
3.0 g/ml)
2.5 MC-MH-B3 2.0 before 1.5 after 14d MC-MH
Honey release ( release Honey 1.0
before 0.5 MC 0.0 500 1000 1500 2000 2500 3000 3500 4000 2 4 6 8 10 12 14 A) Wavenumber (cm-1) B) Immersion time (days)
Figure 46: A) FTIR of MC and MC-MH foams before and after immersion in SBF for 14 days. B) Release of MH as a function of immersion time in SBF 80 4 Fabrication of bioactive glass containing wound dressings by freeze drying
In contrast to MC and MC-MH foams, as shown in Figure 47, BG containing foams show first signs of formation of a CaP-rich phase after 3 days. With increasing immersion time, the formation of calcium-phosphate species on the surface of the foams appears more clearly. Especially in the case of MC-MH-B3-Cu foams, the surface seems to be fully covered after 14 days of immersion in SBF. Therefore, it seems that the addition of Cu-BG leads to faster formation of CaP-rich layer and thus to a higher bioactivity of the composite foams.
Figure 47: SEM images of MC-MH-BG and MC-MH-BG-Cu foams immersed in SBF for 3, 7 and 14 days, indicating the formation of a calcium phosphate rich phase on the foam surfaces
To evaluate the found species on the surface of the BG-containing MC-MH foams, FTIR as well as XRD were conducted on the samples after immersion in SBF for 14 days. According to Figure 48 B), peaks attributed to the formation of CaP-rich phases on the surface of the freeze dried foams can be found. For instance, the peaks at 1000-1100 cm-1 can be ascribed to 4 Fabrication of bioactive glass containing wound dressings by freeze drying 81
P-O stretching, whereas the peaks at 600 cm-1 and 560 cm-1 can be ascribed to P-O vibrations [114], [213], [214]. To further examine if the formed CaP rich deposit crystallize into hydroxyapatite, XRD measurements were conducted. During these measurements, no visible peaks could be found (Figure 48 B). Therefore it seems that the formed CaP deposit does not crystallize. This is in accordance to the observations made in section 3.3.3.2, where no crystalline phase after immersion the different BG into SBF could be found. However, also here it is important to mention that this result could be due to the detection limit of XRD. In summary, the dissolution study using SBF showed that, by the addition of BG into the MC-MH foams, the bioactivity of the composite foams could be enhanced. Moreover, it could be also shown that MH was released (to a certain extent) during the immersion in SBF.
MC-MH-B3-Cu MC-MH-B3 MC-MH-B3-Cu P-O P-O MC-MH-B3 After 14d after 14d
Before before
After 14d Intensity (a.u.) Intensity
after 14d Absorbance (a.u.) Absorbance Before before
1300 1200 1100 1000 900 800 700 600 500 20 25 30 35 40 45 50 55 60 A) Wavenumber (cm-1) B) 2 (°)
Figure 48: A) FTIR and B) XRD of BG containing foams before and after immersion in SBF for 14 days
4.3.4 Antibacterial efficiency
Besides providing bioactivity, wound dressings should also provide ideally antibacterial properties. An antibacterial test using bacteria involved in wound infections was performed [2], [38], [162]. As expected, foams based just on MC did not provide any antibacterial effect (Figure 49). More interestingly, the addition of MH improved the antibacterial effect of the freeze-dried foams, especially against the gram-negative bacteria E. coli. The effect of MC-MH foams was similar to MH alone, no significant difference could be found. Surprisingly, the addition of B3 seemed to disturb the antibacterial effect of MH. In contrast, the addition of B3-Cu further improved the antibacterial effect of the composite foam. Overall, the dual effect of copper and MH seems to kill efficiently both kinds of bacteria.
82 4 Fabrication of bioactive glass containing wound dressings by freeze drying
E. coli S. aureus 140
* 120 * 100 # 80
60 # 40
Relative bacteria viability (%) viability bacteria Relative # 20 # # # 0 MC MC-MH MC-MH-B3 MC-MH-B3-Cu MH
Figure 49: Relative viability of gram-negative E. coli and gram-positive S. aureus bacteria, cultivated in contact with the different MC foams. One-way ANOVA statistical analysis denotes significant differences (*p<0.05), # symbolizes significant differences (p<0.05) compared to MC for both kind of bacteria.
4.3.5 Compatibility with different skin cells
4.3.5.1 Conditioned CCM containing dissolution products used for cell experiments
In order to understand and discuss the effects that MC-based foams containing additionally BG and MH could have on cells, conditioned CCM was prepared and examined using ICP and pH measurements. As shown in Figure 50 A, no significant difference between the pH of CCM and CCM containing dissolution products of MC and MC-MH could be found. As expected, the addition of B3 as well as B3-Cu led to an increase of pH up to 8.6 ± 0.1. In accordance with the results obtained by pH measurements, MC and MC-MH foams had no influence on the ion concentration of the CCM. In contrast, B3 and B3-Cu led to a mayor increase of B and Ca and to a minor increase of Mg and P. Moreover, in case of MC-MH-B3-Cu, 3.6 ± 0.2 mg/L copper could be detected in the conditioned CCM. Besides the copper, no significant difference between B3 and B3-Cu could be found. 4 Fabrication of bioactive glass containing wound dressings by freeze drying 83
* 9.00 *
8.75
8.50
8.25 pH
8.00
7.75
7.50
MC MC-MH Control MC-MH-B3 A) MC-MH-B3-Cu
B 250 Ca Cu 225 K 200 Mg P 175
150
125
100 * * 75
Ion concentration (mg/L) concentration Ion 50
25 * * * * * * 0 *
CCM MC MC-MH MC-MH-B3 B) MC-MH-B3-Cu
Figure 50: Results of A) pH and B) ICP measurements of conditioned CCM containing dissolution products of different fabricated foams. One-way ANOVA statistical analysis denotes significant differences compared to with the pure CCM (*p<0.05). 84 4 Fabrication of bioactive glass containing wound dressings by freeze drying
4.3.5.2 Indirect cell tests using MEF cells
Figure 51: Fluorescence images of actin filaments (red) and cell nuclei (blue) of MEF cells cultured for 24h in CCM containing dissolution products of freeze dried MC based foams.
In order to evaluate if the produced freeze-dried foams based on MC exhibit cytotoxic effects, preliminary cell tests using MEF cells were performed. As shown in Figure 51, the morphology of MEF cells cultured in CCM containing dissolution products of MC and MC-MH foams shows no significant differences to the cells cultured in CCM (positive control). Moreover, the addition of B3 BG appeared not to negatively influence the growth of MEF cells. However, the number of MEF cells in contact with B3-Cu BG seems to be lower compared to 4 Fabrication of bioactive glass containing wound dressings by freeze drying 85 the positive control, but still it is higher compared to MEF cells cultured in 6% DMSO. To quantify the effect of the different conditioned CCM on the growth of MEF cells, the cell viability was additionally measured (Figure 52). MC and MC-MH foams showed a significant improvement of cell viability, whereas MC-MH-B3 foams seem not to have any significant impact compared to the positive control. As already observed in the fluorescence images (Figure 51), the addition of B3-Cu had a negative impact on the cell viability. However, this negative impact was not as high as for 6% DMSO (negative control).
140 # # 120
100
80
60
40 Relative cell viability (%) viability cell Relative
20
0
MC CCM MC-MH MC-MH-B3 6% DMSO MC-MH-B3-Cu
Figure 52: Relative viability of MEF cells cultured in conditioned CCM containing dissolution products of the different fabricated foams. By one-way ANOVA statistical analysis significant differences (*p<0.05) could be found for all samples, except when denoted by #. 86 4 Fabrication of bioactive glass containing wound dressings by freeze drying
4.3.5.3 Direct cell test using human dermal fibroblast (hDF)
Figure 53: Fluorescence (left) and SEM images of hDFs grown on MC, MC-MH, MC-MH-B3 and MC-MH- B3-Cu foams for 7 days
4 Fabrication of bioactive glass containing wound dressings by freeze drying 87
By performing a direct cell test using hDFs, the biocompatibility of the produced foams was further investigated. In Figure 53, images show hDFs grown on the surface of the different foams. No significant difference between the different foams can be observed. In all cases, it seems that the surface of the foams is completely covered with hDFs. The morphology of hDFs shown in the SEM images is typically for hDFs, which grow in a dense layer on top of the porous foams [262]–[264]. Moreover, the cell viability of hDFs cultured on the different foams was measured after 1 and 7 days. A clear increase of cell viability from day 1 to day 7 can be observed for all types of foams. This result indicates that besides offering a suitable surrounding for attachment, the different MC-based foams also promote cell proliferation. Additionally, although (surprisingly) foams containing B3 show better results after 1 day of cultivation, after 7 days no significant difference between the foams could be measured. Therefore, in a direct contact and for longer cultivation periods, B3-Cu containing foams were confirmed to be non-cytotoxic.
* Day 1 Day 7 *
MC MC-MH MC-MH-B3 MC-MH-B3-Cu
Figure 54: Cell viability of hDFs grown directly on different MC-based foams for 1 and 7 days. One-way ANOVA statistical analysis denotes significant differences (*p<0.05). 88 4 Fabrication of bioactive glass containing wound dressings by freeze drying
4.3.5.4 In vitro scratch test using mouse embryotic fibroblast (MEF) cells and human keratinocytes-like (Hacat) cells
Figure 55: Light microscope images showing exemplary the migration of MEF cells under different conditions during the scratch test 4 Fabrication of bioactive glass containing wound dressings by freeze drying 89
To analyse the capability of the releasing products of the fabricated foams to support the migration of relevant cells, an indirect in vitro scratch test was performed. Figure 55 shows exemplary the migration of MEF cells cultured in CCM and in CCM containing 6% DMSO versus time. It is obvious that MEF cells cultured in CCM are able to close “the wound” after 24 hours, whereas in the presence of 6% DMSO the MEF cells seem even to die [252]. By further analysing the microscope pictures obtained during the test, the migration ratio for the different samples could be calculated. As shown in Figure 56 A, no significant difference between the positive control (just CCM) and CCM containing dissolution products of the different freeze dried MC-based foams could be found. Since, as described in section 2.1.2, during the proliferation phase keratinocytes migrate into the wound and start the re- epithelialization, the scratch test was also performed using keratinocytes (Figure 56 B). Here, the presence of dissolution products of MC-MH and MC-MH-B3 appeared to improve the migration ability of the keratinocytes and therefore improve their capability to close the “wound”. On the contrary, MC and MC-MH-B3-Cu foams did not have any significant impact on the migration ratio comparted to the negative control.
CCM CCM 120 120 6% DMSO 6% DMSO MC MC MC-MH 100 MC-MH 100 MC-MH-B3 MC-MH-B3 MC-MH-B3-Cu MC-MH-B3-Cu 80 80
60 60
40 Migration ratio (%) ratio Migration Migration ratio (%) ratio Migration 40
20 20
0 0 5 10 15 20 25 30 35 0 5 10 15 20 25 30 35 A) Time (h) B) Time (h)
Figure 56: Migration and therefore closure of scratch (indicated by the migration ratio) of A) MEF and B) Hacat cells as a function of time
4.3.5.5 Co-culture of human dermal fibroblasts (hDFs) and human umbilical vein endothelial cells (HUVECs)
By performing cell studies using one type of cells (e.g. fibroblasts, keratinocytes) cellular interactions, which may occur during wound healing in vivo cannot be studied. In this investigation therefore, hDFs were co-cultured with HUVECs. These types of cells were chosen since cytokines released by (among others) endothelial cells trigger the migration of fibroblasts during wound healing [265]. Moreover, cytokines released by fibroblasts (such as VEGF) trigger proliferation and migration of endothelial cells [266]. 90 4 Fabrication of bioactive glass containing wound dressings by freeze drying
Figure 57: Fluorescence images of calcein (green) and cell nuclei (blue) of co-cultured hDF and HUVECs for 24h in CCM containing dissolution products of freeze dried MC based foams.
As shown in Figure 57, co-cultured hDFs and HUVECs were seem to grow in a dense layer when in contact with MC and MC-MH foams compared to the positive control. However, by the addition of B3 and B3-Cu, the cell layer was not as dense as found for the positive control, but more cells could be found compared to the negative control (6% DMSO). Since it is not possible to distinguish between the two types of cells in fluorescence images, SEM images were additionally inspected (Figure 58). Whereas hDFs are flat and wide spread, HUVECs exhibit a more round shape [267], [268]. In accordance with the fluorescence images, the number of cells grown in 6% DMSO seems to be lower in comparison to the positive control. 4 Fabrication of bioactive glass containing wound dressings by freeze drying 91
Moreover, no significant differences between the cells cultured in CCM and conditioned CCM could be found.
Figure 58: SEM images of co-cultured hDF and HUVECs for 24h in CCM containing dissolution products of freeze-dried MC based foams. HDFs are marked with arrows having a solid line and HUVECs are marked with arrows having a dashed line.
To better evaluate the impact of the dissolution products from the MC-based foams, cell viability was additionally measured. As shown in Figure 59, the viability of hDFs and HUVECs in co-culture decreased in the presence of B3 and B3-Cu. However, the viability was still higher than that of the negative control. To understand the effects of co-culturing, the same experiments were conducted with hDFs and HUVECs separately. In case of HUVECs, also a reduction of cell viability by B3 and B3-Cu could be found. However, the reduction was not apparent in case of hDFs cultured in contact with B3 BG. Therefore, it can be concluded that 92 4 Fabrication of bioactive glass containing wound dressings by freeze drying
B3 seems to have a negative impact on HUVECs, but not on hDFs, whereas B3-Cu exhibits cytotoxicity on both types of cells.
MC MC-MH 180 MC-MH-B3 160 MC-MH-B3-Cu CCM 140 # 6% DMSO # # 120
100 # # 80
60
Relative cell viability (%) viability cell Relative 40
20
0 hDF HUVEC Co-Culture
Figure 59: Relative viability of hDFs and HUVECs co-cultured in conditioned CCM containing dissolution products of the different fabricated foams. By one-way ANOVA statistical analysis significant differences (*p<0.05) could be found for all samples, except denoted by #.
4.4 Discussion
As described in section 2.1.3, research focusing on the development of multifunctional wound dressings is facing two main challenges, namely the generation of suitable biomaterials and the fabrication, out of these biomaterials, of suitable wound dressings. As described, wound dressings need to fulfil several requirements. These requirements (summarized in section 2.1.3) can be divided in three groups: practical, physical and biological requirements. Based on these requirements, the freeze-dried foams can be evaluated in terms of their potential for wound dressing material.
4.4.1 Practical properties
A suitable wound dressing should be produced in high quantities and it must be readily available and cost effective [34], [35]. Due to the fact that MC based foams can be produced by dissolving MC in water, no toxic solvents are needed which can lead to possible residuals inside the obtained samples. Moreover, the use of toxic solvents has several disadvantages in respect to environmental and safety reasons [269]. However, besides 4 Fabrication of bioactive glass containing wound dressings by freeze drying 93 offering economical and safety advantages, the use of water as solvent leads to a restriction of the choice of material. Here, it could be shown that by dissolving MC (in combination with MH and BG) in water, no phase separation occurred and therefore MC (and MC-MH) scaffolds with smooth surface could be produced by freeze drying. Additionally, also the BG particles could be homogenously dispersed in the MC-MH sol and composite scaffolds with well distributed BG particles could be obtained. By using the freeze-drying method, it could be shown that high quantities could be produced by using raw materials which are not extremely expensive (e.g. costs of water compared to organic solvents). The freeze-dried foams are easy to cut in desired shapes; they can be readily handled and applied to the wound. Moreover, it is possible to disinfect the produced foams by UV-light. However, disinfection might be most probably not enough for clinical use; therefore the sterilizability of the foams needs to be examined.
4.4.2 Physical properties
Foams must provide some specific physical properties suitable for wound dressings. Here, one key factor is the porosity [270]. An open structure with interconnected pores of a suitable size is needed to guarantee waste and gas exchange. All produced foams exhibit a pore size of more than 90%, which could be even slightly increased by the addition of BG particles. This is in contrast to results obtained in literature, where similar porosities could be found for other materials (e.g. alginate [271], gelatin-MH [272], silk fibroin-MH [273], [274]), but the addition of BG led to a reduction of porosity [275]–[277]. Besides porosity, a crucial aspect is the mechanical performance of such porous structure. Here, similar results to those of other foam materials could be obtained [272]. Similar to the literature, the addition of BG particles leads to a minor decrease of the compression strength. However, it could be shown that the addition of BG is a promising possibility to improve the mechanical properties. Therefore, the mentioned disadvantage of freeze dried foams (section 2.3), the lack of suitable mechanical strength, could be tackled by making composites. Most interestingly, the introduction of MH leads to a self-healing foam, which makes it ideal for applying the foams into a wound. During the introduction of the foams into the wound site, it can be pressed in and, with time, the foams will expand and will perfectly fill the wound. Therefore, due to the suitable porosity and mechanical performance, freeze dried MC-based foams are able to fulfil the following requirements: covering and protecting the wound, proper oxygenation, thermal insulation and removal of waste. Moreover, since the foams turn into a gel during application (as shown in the acellular bioactivity study in Section 4.2.3), they are able to keep a moist environment and therefore reduce the water and blood loss. 94 4 Fabrication of bioactive glass containing wound dressings by freeze drying
These requirements must be fulfilled over a certain period of time. Therefore, besides mechanical properties, also the degradation rate of the foams is crucial. Here, we successfully showed that it is possible to crosslink MC with MH leading to an improved mechanical performance (self-healing ability), degradation behaviour and release rate (proved by the MH and BG release). MH offers not only the advantage of being a non-toxic crosslinking agent in comparison to other chemical crosslinking agents (e.g. glutaraldehyde [148], [278]), but also it introduces favourable biological effects.
4.4.3 Biological properties
The release of MH additionally provides advantages in terms of biological properties of the produced foams. One necessary requirement that multifunctional wound dressings suitable for chronic wounds have to fulfil is to provide antibacterial effects [9], [36]. In recent years, it was not possible to develop any new antibiotic and, additionally, new multi-resistant bacteria appeared [3], [279]. MH is well-known for its antibacterial effect [164] and it could be shown that in combination with MC, MH had especially an antibacterial effect against the gram-negative bacterium E. coli. Surprisingly, this effect was not evident in case of MC-MH-B3 foams. This result could be due to the effect of the basic pH due to the release of dissolution products of B3, which compensates the acidic pH due to the release of MH [3], [38], [85]. However, in case of B3-Cu, an optimal antibacterial effect against both, gram-negative and gram-positive bacteria could be observed. BG as well as herbal extracts such as MH showed in the past that they can preserve their effectiveness [22]. For instance, in recent research, bacteria were repeatedly exposed to a silicate BG (S53P4 composition containing in wt.%: 53
SiO2, 23 Na2O, 20 CaO and 4 P2O5) and no change of the minimal inhibitory concentration could be observed [32].
Besides the antibacterial effect, ideal wound dressings should also provide an optimal environment for the migration and proliferation of cells. This could be warranted by a suitable porosity and pore structure. Here, the pore size is also a crucial aspect. All produced foams exhibit suitable pore size for cell infiltration, distribution of nutrients and angiogenesis [270]. Moreover, it is a common approach to test the bioactivity of samples by immersion in SBF. The bioactivity is then determined by the ability to form hydroxyapatite on the surface of the sample [54], [186], [187]. Here, the MC-MH foams doped additionally with BG showed the formation of CaP-rich species during immersion time, however it seems that this calcium- phosphate layer does not crystallize into hydroxyapatite. Although the ability to form hydroxyapatite is crucial for applications in hard tissue engineering, it is still not clear 4 Fabrication of bioactive glass containing wound dressings by freeze drying 95 whether it is necessary in wound healing applications [83]. On one hand, Wilson et al. [280] showed that due to the formation of hydroxyapatite on top of BG a strong bonding occurred between the BG and soft tissue, but, on the other hand, studies [84], [281] have shown that hydroxyapatite can lead to soft tissue calcifications, which should be avoided. Therefore, it is still an open question if the formation of hydroxyapatite is needed for wound healing applications and consequently it is not a clear disadvantage (or advantage) if the freeze dried foams here are not showing any hydroxyapatite formation.
The performed cell tests, in contrast, show clearly the favourable biological effect of the composite MC-MH foams. A previous study [175] focusing on the effect of MH showed that high concentrations of MH can have a cytotoxic effect on macrophages, hDFs and human pulmonary microvascular endothelial cells. On the other hand, the same study showed that in low concentrations MH seems not to have any effect on the tested cells [175]. Here, the addition of MH leads in all performed cell tests to increased cell viability and additionally in case of Hacat cells, to an improved cell migration. Therefore, by crosslinking MC with MH, a favourable release and therefore a positive biological effect of MH could be achieved. The performed cell tests showed also that B3 BG has no negative impact on the cell viability of the produced foams. Therefore, in contrast to results from literature [82], [190], borate B3 BG seems, if combined with MC-MH foams, not to lead to cytotoxic effects. However, by using B3- Cu BG, the cell viability of all kind of cells (HUVECs and hDFs) was reduced in the indirect cell test. This effect was not as strong as compared to the negative control, and direct cell studies for longer periods using hDFs showed that cells in presence of copper can still proliferate. Moreover, scratch tests using hDFs and Hacat cells showed no influence of copper on the migration of the used cells. The results are in accordance to the results obtained in section 3.3.5, where a critical level of 5 mg/L of copper could be found. This level is even lower than the one reported in literature [244], which was at 10 mg/L of copper in contact with 3T3 fibroblasts. In accordance to another study [282], where the authors found that even at a level of 10 mg/L copper seems not to have a high impact on the viability of endothelial cells, also here the impact on hDFs was higher than that on HUVECs. These results clearly show that more tests are necessary to exactly determine whether the MC-MH foams containing B3-Cu are suitable as wound dressings with intrinsic antibacterial functionality.
96 4 Fabrication of bioactive glass containing wound dressings by freeze drying
5 Fabrication of bioactive glass containing wound dressings by 97 electrospinning
5 Fabrication of bioactive glass containing wound dressings by electrospinning9
5.1 Introduction
As described in section 2.3, electrospun fibre mats containing therapeutic agents (e.g. BG particles) are suitable candidates for wound dressings, based (among others) on their ability to mimic the structure of the ECM [115], [116]. By combining the BGs described in Chapter 3 in particulate form with MC and MH (section 2.4 and 2.5) in electrospun fibres, wound healing patches were generated. In this Chapter the processing as well as the full characterization of electrospun fibre mats based on MC and MH containing B3 BG particles will be presented.
5.2 Experimental procedure
5.2.1 Production of electrospun fibre mats
MC and MH (section 4.2.1) and B3 BG in particulate form (section 2.2) were used to produce the electrospinning solution. The used MC, MH and B3 particles (preparation described in section 4.2.1) were dissolved in formic and acidic acid (VWR, Darmstadt, Germany) in a ratio 1:1 (AA/FA). Based on optimization of the electrospinning solution, PCL (80 kDa, Sigma Aldrich, Germany) was additionally used. The details of the optimized solutions are summarized in Table 10. Electrospinning was performed by using a commercially available setup (Starter Kit 40KV WEB, Linari Engineering srl, Valpiana (GR), Italy). Solutions based on PCL were stirred for 4 hours, whereas the solutions based on blends were stirred overnight. After adding MH, all solutions were stirred one hour more. After adding B3 particles, solutions PCL-MH-B3 and Blend-MH-B3 were stirred for another 5 min. During optimized electrospinning the following parameters were used: an applied voltage of
9 Some of the results presented in this chapter were previously published by the author (as first author) and are hereby adapted from reference [355] 98 5 Fabrication of bioactive glass containing wound dressings by electrospinning
20 kV, a flow rate of the polymeric solution of 1cc, a needle diameter of 21G and the distance needle-collector of 11 cm.
Table 10: Overview of the concentrations of the different electrospinning solutions used to prepare fibrous mats
B3 [wt.% respect to Fibre mat PCL [%w/v] MC [%w/v] MH [%w/v] polymer] PCL 15 PCL-MH 15 4 PCL-MH-B3 15 4 30 Blend 15 4 Blend-MH 15 4 4 Blend-MH-B3 15 4 4 30
5.2.2 Characterization
5.2.2.1 Scanning electron microscopy (SEM) and energy-dispersive X-ray analysis (EDX)
SEM combined with EDX (some instrument described in section 4.2.2.1) was used to examine the morphology of different fibre mats and to evaluate the distribution of the BG particles inside the mats. Based on SEM images, the fibre diameter of 50 fibres from each sample was calculated using ImageJ (NIH, Bethesda, MD) [254]. The SEM was operated at 1 kV and working distance of 2 mm. Fibre mats were fixed on a sample holder by carbon tape.
5.2.2.2 Fourier transform infrared spectroscopy FTIR
FTIR (some instrument described in section 4.2.2.3) was performed on composite fibre mats before and after immersion in SBF, produced as described in section 3.2.2.7 [186].
5.2.2.3 Mechanical testing
By using a universal testing machine (Instron 5960, Germany), tensile strength of the produced fibre mats was assessed Samples were cut in rectangular shape (length 20 mm, width 3mm) and fixed in a frame/holder (paper, 30 mm2) to ensure no pretension [114]. At least 10 replicates were tested using a load cell of 100N and a speed of 10 mm/min. 5 Fabrication of bioactive glass containing wound dressings by 99 electrospinning
5.2.2.4 Contact angle measurement
In order to evaluate the wettability of the fibre mats, the contact angle was measured as described in section 4.2.2.5.
5.2.2.5 X-ray diffraction (XRD) measurement
After SBF dissolution tests, the phase compositions of the electrospun fibre mats were characterized using a powder X-ray diffraction (XRD) measurements as described in section 4.2.2.6.
5.2.3 In vitro dissolution in SBF
By using the same SBF and protocol as described in section 3.2.2.7 (SBF was produced according to the recipe of Kokubo et al. [186]), the bioactivity of the fibre mats was evaluated. The samples were immersed in a ratio of 1.5 g of sample to 1L of SBF [187]. After different immersion durations in SBF (3, 7 and 14 days), the samples were collected for analysis by FTIR, XRD and SEM. Additionally, the pH of the remaining SBF was recorded (section 4.2.3) and by using a UV-Vis spectrophotometer (section 4.2.3). All samples and measurements were performed in triplicate.
5.2.4 Antibacterial efficiency of the fibre mats
The antibacterial efficiency of the fabricated fibre mats was evaluated using both, gram positive and gram negative bacteria, S. aureus and E. coli, as described in section 4.2.4. All measurements were performed in triplicates.
5.2.5 In vitro cell tests
Prior to cell tests, the fibre mats were cut manually using a razor blade (Herkenrath Solingen) and fixed in scaffold supports (scaffdex, Sigma Aldrich). Disinfection was performed for 1h using UV light. Conditioned media was prepared as described in section 3.2.4. Briefly, electrospun fibre mats were immersed in CCM (DMEM supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin) in a ratio of 1mg/1ml CCM at 37 °C. Electrospun fibre mats were incubated in a humidified atmosphere (95% air and 5% CO2). After incubation for 24h, the conditioned media was collected and analysed using ICP-OES (ICP-OES, Vista MPX, 100 5 Fabrication of bioactive glass containing wound dressings by electrospinning
Varian) in order to measure the amount of ions in the conditioned media. To analyse the conditioned media using ICP-OES, the conditioned medium was stabilized using 1ml of 1M nitric acid to reach a pH of 2. Samples and measurements were performed in triplicate.
5.2.5.1 Direct cell tests using human dermal fibroblasts (hDFs)
To further prove the cytocompatibility of the samples, fibre mats were tested in direct contact with hDF cells. Therefore, 50,000 hDFs were directly seeded on top of each sample using a 50 µl CCM drop. Then, the electrospun fibre mats were incubated for 15 min (fixed in scaffold supports in a 24-well-plate). Subsequently, they were filled with 1ml CCM. Electrospun fibre mats with cells seeded on the surface were then incubated for 1 and 7d. Following this, by performing a mitochondrial activity calorimetric assay, which is based on tetrazolium salt (WST-8 as-say, Sigma-Aldrich), the viability of hDFs was assessed. Additionally, the cell morphology was evaluated by SEM analysis. For SEM analysis, the samples were fixed as described in section 4.2.5.2.
5.2.5.2 In vitro scratch test using human dermal fibroblast (MEF) and human keratinocytes-like (Hacat) cells
Based on a protocol established by Liang et al. [252], an in vitro scratch tests was carried out using Hacat cells, spontaneously immortalized human keratinocyte line [283], in order to analyse the migration of the different cells. Hacat cells were used since keratinocytes are involved in wound closure as described in section 2.1.2 [174]. Conditioned CCM used for this test was produced as described in section 4.2.5. First, 500,000 Hacat cells per 1ml CCM were seeded in a 24-well plate and incubated for 24h as described above. Then, by using a 200 µl pipette tip, a scratch was created on the grown dense Hacat cell monolayer. Subsequently, the cells were washed using PBS and then incubated with conditioned CCM (section 5.2.5) for 24h. Again, a positive control (pure CCM) and a negative control (CCM supplemented with 6% DMSO) were used. The migration of Hacat cells was observed during the incubation after different time points (2.5h, 5h, 8h and 24h) by using a light microscope (Primo Vert, Carl Zeiss). All measurements were done in triplicates and by using ImageJ software [254], the width of the scratch was analysed. 5 Fabrication of bioactive glass containing wound dressings by 101 electrospinning
5.3 Results
5.3.1 Production and characterization of MC-MH-B3 fibre mats
In order to verify the suitability of using MC in combination with benign solvents to produce fibre mats by electrospinning, different parameters were investigated. As shown in Figure 60, solutions containing 6%w/v and 4%w/v of MC in AA/FA were tested. It could be clearly observed that for the higher concentrations no continuous fibres could be produced most probably due to the fast evaporation of the solvent. However, for 4%w/v MC solutions, it was possible to obtain fibres in the nanometre range. Additionally, the influence of the applied voltage was studied. In case of 15kV, electrospun mats full of beads were obtained. Similar to results from the literature [284], better results were obtained for higher voltages, however due to technical limitations it was not possible to test at voltages higher than 20 kV. Therefore, for further experiments an applied voltage of 20kV was used in combination with 4 %w/v MC solution, since this combination led to the best results in terms of fibre formation.
Figure 60: SEM images of MC-based fibres produced by electrospinning using different parameters (the insets are showing the respective fibre mat in higher magnification). 102 5 Fabrication of bioactive glass containing wound dressings by electrospinning
Further, MH was introduced into the optimized MC-based fibres in a ratio 1:1 based on the promising results presented in Chapter 4. Then, additionally, B3 BG was introduced into the MC-MH fibre mats. According to Figure 61, MC-MH fibre mats are characterized by nanofibers and numerous beads. However, the addition of B3 BG seems to improve the electrospinning potential of the MC-MH solution. Fibre mats showing continuous fibres without beads could be obtained with increased fibre diameter (still in the nanometre range).
Although it was possible to produce electrospun fibre mats based on MC in the nanometre range with and without the addition of MH and B3, the obtained fibre mats were too thin to be able to remove them from the aluminium foil. Even after several hours of electrospinning, it was not possible to obtain a thickness suitable for further use. Therefore, MC-based fibres were blended with PCL, a well-known polymer for electrospinning [114], in order to achieve suitable fibre mats of sufficient structural integrity for wound healing applications.
Figure 61: SEM images of MC-MH fibres containing different amounts of BG particle addition (compared to the polymer content) (the insets are showing the respective fibre mat in higher magnification).
5.3.2 Characterization of optimized fibre mats
Based on the results presented in section 5.3.1, MC fibres were blended with PCL to achieve higher quantities of fibre mats. As control, PCL, PCL-MH and PCL-MH-B3 fibre mats 5 Fabrication of bioactive glass containing wound dressings by 103 electrospinning were additionally fabricated. Based on Figure 62, for all different kind of electrospinning solutions, fibre mats without beads could be obtained. Based on the SEM images, the average fibre diameter was calculated (Table 11). In case of Blend fibres, the introduction of MH as well as B3 seems not to have a significant influence. In contrast, the addition of MH and B3 (although not significantly) increases the average fibre diameter of PCL-based fibres. This result, as well as the increase of standard deviation (showing a greater fibre diameter distribution) has been also reported in literature [113].
A) B)
C) D)
E) F)
Figure 62: SEM images of PCL- and Blend-based fiber mats: A) Blend, B) PCL, C) Blend-MH, D) PCL-MH, E) Blend-MH-B3 and F) PCL-MH-B3. The insets are showing the respective fibre mat in higher magnification.
In order to verify the presence as well as the distribution of B3 BG particles inside the electrospun fibre mats, EDX analysis was performed. Based on the EDX measurements made 104 5 Fabrication of bioactive glass containing wound dressings by electrospinning on different spots of different fibre mats, the well distributed presence of B3 particles in Blend-MH-B3 and PCL-MH-B3 fibre mats could be proven as shown in Figure 63.
Al 80 60 Blend-MH-B3 PCL-MH-B3 Al 70 50
60 40 50
40 30
cps/eV cps/eV
30 20 C 20 C 10 O 10 O B B O Mg P K Ca Na P K Ca 0 0 0 2 4 6 8 10 0 2 4 6 8 10 EDX spectra EDX spectra
Figure 63: EDX spectra of the Blend-MH-B3 and PCL-MH-B3 fibre mats, indicating the presence of B3 particles inside the fibre mats (Al was measured since the fibre mats were attached to alumina foil).
The wettability of the electrospun fibre mats was also investigated. As summarized in Table 11, the Blend fibre mats exhibited a lower contact angle than PCL fibre mats, showing their hydrophobic nature [251]. Moreover, it could be shown that for both, Blend and PCL- based fibre mats, the contact angle could be further reduced by the addition of MH and B3 particles (Table 11).
Table 11: Average fibre diameter and contact angle measured for different PCL- and Blend-based fibre mats
Blend- PCL-MH- Fibre mat Blend Blend-MH PCL PCL-MH MH-B3 B3 Average fibre 165 ± 74 171 ± 64 195 ± 76 173 ± 61 238 ± 106 375 ± 323 diameter [nm] Contact angle 72 ± 1 62 ± 2 45 ± 2 140 ± 3 90 ± 2 76 ± 2 [°] Ultimate tensile 43 ± 11 5 ± 2 4 ± 1 5 ± 1 2 ± 1 3 ± 1 strength [MPa] Tensile Strain 51 ± 11 60 ± 15 75 ± 11 52 ± 5 31 ± 10 52 ± 14 [%]
5 Fabrication of bioactive glass containing wound dressings by 105 electrospinning
Moreover, the mechanical performance of the fibre mats was evaluated by tensile tests (Table 11). Here, Blend fibres showed a higher ultimate tensile strength (UTS, 43 ± 11 MPa) compared to PCL fibres (5 ± 1 MPa). The addition of MH was seen to lead to a dramatically decrease of UTS of Blend fibres, whereas the influence on PCL fibres was less. In contrast, the addition of B3 BG particles seems not to have a significant influence on the mechanical properties.
5.3.3 Acellular bioactivity using SBF
In accordance to the evaluations made in sections 2.23 and 4, the produced electrospun fibre mats were also evaluated using SBF. In contrast to freeze dried foams (section 4.3.3), the immersion of electrospun fibre mats does not lead to a significant increase of pH (Figure 64). This could be explained by the minor amounts of B3 BG particles incorporated in the fibre mats.
Blend 7.85 Blend-MH Blend-MH-B3 7.80 PCL 7.75 PCL-MH PCL-MH-B3 7.70
7.65 pH
7.60
7.55
7.50
7.45
2 4 6 8 10 12 14 Immersion time (days)
Figure 64: Change of pH as a function of immersion time of different electrospun fibre mats in SBF
As expected, PCL, PCL-MH, Blend and Blend-MH fibre mats did not show any sign of apatite formation after 14 days of immersion in SBF (Figure 65). 106 5 Fabrication of bioactive glass containing wound dressings by electrospinning
Figure 65: SEM images of electrospun fibre mats immersed in SBF for 14 days showing the degradation of the electrospun fibre mats. Images in higher magnification (inserts) do not indicate apatite formation.
In contrast, according to Figure 66, fibre mats including B3 BG particles (PCL-MH-B3 and Blend-MH-B3) show first signs of apatite formation after 3days immersion in SBF. This becomes more obvious after 7 days. However, it seems that after 14 days less apatite formation is visible, probably due to the ongoing degradation and the loss of the formed apatite during sample washing. In comparison, Blend-MH-B3 samples show faster and more obvious formation of an apatite layer compared to PCL-MH-B3 samples. Moreover, the images indicate that apatite forms mainly on the B3 BG particles and not on the polymer surface.
5 Fabrication of bioactive glass containing wound dressings by 107 electrospinning
Figure 66: SEM images of electrospun fibre mats containing B3 BG particles after 3, 7 and 14 days of immersion in SBF showing the degradation and apatite formation of the fibre mats. Inserts show the formed apatite in higher magnification.
To further evaluate the characteristics of the formed apatite layer, samples were characterized by FTIR measurements. According to Figure 67 A, FTIR spectra of electrospun samples prior immersion in SBF show the main bands related to PCL. The peaks at 2944 and
2866 cm-1 can be attributed to stretching vibrations of CH2, the peak at 1722 cm-1 can be assigned to vibrations of C-O and the peaks at 1294 cm-1, 1240 cm-1 and 1169 cm-1 can be attributed to stretching vibration of C-O and C-C [114], [285]. Moreover, the already described peaks (section4.3.1) related to MH and MC overlap with the peaks related to PCL, therefore no clear differences between the spectra of the different fibres can be observed. The addition of B3 particles inside the Blend-MH and PCL-MH fibres led to a shoulder in the spectra, centred at 860 cm-1, thus confirming the addition of BG to the fibre [196]. Figure 67 B) shows the FTIR 108 5 Fabrication of bioactive glass containing wound dressings by electrospinning spectra of fibre mats immersed in SBF for 14 days. New peaks appear at 1027 cm-1 and at
560-600 cm-1, ascribable to PO43- asymmetric stretching vibration and to P-O bending vibration, respectively [114].
PCL-MH-B3 PCL-MH-B3
PCL-MH PCL-MH
PCL
PCL Blend-MH-B3
Blend-MH-B3 Blend-MH Blend-MH
Blend Blend
1400 1200 1000 800 600 400 1400 1200 1000 800 600 400 - A) Wavenumber (cm-1) B) Wavenumber (cm 1)
Figure 67: FTIR spectra of produced electrospun fibre mats before and after immersion in SBF for 14 days. The peaks are discussed in the text.
To evaluate whether amorphous CaP-rich species or crystalline apatite is forming, the electrospun fibre mats were additionally characterized by XRD. According to Figure 68 A), the typical broad peaks for PCL could be found in fibres [286]. Besides these peaks, no additionally peaks could be detected. After 14days of immersion in SBF, new peaks are clearly visible (Figure 68B). These peaks can be mainly attributed to sodium chloride, which was not completely washed out from the samples after immersion in SBF [287]. However, hydroxyapatite shows similar peaks [288] and therefore the formed apatite layer seems (partially) crystallized.
PCL-MH-B3
PCL-MH PCL-MH-B3
PCL-MH PCL PCL
Blend-MH-B3
Blend-MH Blend-MH-B3 Blend-MH Blend Blend
20 30 40 50 60 70 80 20 30 40 50 60 70 80 A) 2Q (°) B) 2Q (°)
Figure 68: XRD pattern of produced electrospun fibre mats before and after immersion in SBF for 14 days 5 Fabrication of bioactive glass containing wound dressings by 109 electrospinning
Further, the release of MH was also evaluated. Like the results obtained in section 4.3.3, due to the fact that MH is composed of several components, it is challenging to measure the exact amount of MH released by the electrospun fibre mats. In Figure 69, it is obvious that no MH release was measured from pure PCL and Blend fibres (as expected). Moreover, MH incorporated in PCL fibres shows a fast release in the first 7 days and an overall higher released amount after 14 days. In contrast, MH incorporated in Blend fibres seems to be more steadily released, showing a more favourable release rate. This result can be explained by the cross-linking of MC with MH. In case of PCL fibres, the addition of B3 BG particles did not show a significant influence on the release rate of MH. In contrast, Blend-MH-B3 fibre mats show a higher MH release than Blend-MH fibre mats, most probably due to the increased degradation of MC-MH by the increased surface area (obtained by the degradation of B3 BG particles). This result supports the finding that in case of PCL fibres, MH is just washed out, whereas in case of MC fibres, the MH is released in a more controlled manner during the degradation of the MC- MH fibres.
10 PCL PCL-MH PCL-MH-B3 8 Blend Blend-MH Blend-MH-B3 6
4
2
Cumulative honey release [µg/mL] release honey Cumulative 0
2 4 6 8 10 12 14 Time (d)
Figure 69: Cumulative release of MH as a function of immersion time in SBF for up to 14 days
5.3.4 Antibacterial efficiency
Since contaminations are a huge risk for chronic wounds and an ideal wound dressing should provide antibacterial properties, the antibacterial potential of the produced electrospun fibre mats was evaluated. All tested samples did not show any antibacterial 110 5 Fabrication of bioactive glass containing wound dressings by electrospinning potential (Figure 70). In contrast to 4.3.4, the addition of MH does not lead to an antibacterial effect. This result can be explained by the very low amount of MH incorporated inside the ES samples compared to the freeze-dried samples. As shown in Figure 70 B), in case of pure MH (the release properties of composites needs to be taken into account), at least 25 mg/ml needs to be incorporated to achieve an antibacterial effect. In case of ES samples, only around 1mg/ml of MH was incorporated. Moreover, the addition of B3 BG particles was not effective to provide antibacterial effect, which is in accordance to the results obtained in section 4.3.4.
E. coli S. aureus 1.6 S. aureus E. coli 120 1.4
1.2 100
1.0 80
0.8 60 0.6 40
0.4 (%) viability Relative OD measurement (600 nm) (600 measurement OD 0.2 20
0.0 0
PCL Blend Control PCL-MH Control Blend-MH 100 mg/ml 50 mg/ml 25 mg/ml A) PCL-MH-B3 Blend-MH-B3 B) 12.5 mg/ml6.25 mg/ml3.125 mg/ml
Figure 70: Measured viability of E. coli and S. aureus cultivated A) in CCM containing dissolution products of different electrospun fibre mats and B) in different concentrations of MH. All samples were measured in triplicate.
5.3.5 Compatibility with different skin cells
5.3.5.1 Conditioned CCM containing ionic dissolution products
To examine the results of in vitro cell biology tests presented in the following sections, the used conditioned CCM containing ionic dissolution products of the electrospun fibre mats was measured by ICP-OES measurements. According to Figure 71, besides boron, no significant difference between the pure CCM and CCM containing dissolution products of PCL- MH-B3 and Blend-MH-B3 could be found. Since CCM contains already high amounts of Ca, K, Mg and P, the release of these elements from the BG seems not to have a significant influence on their concentrations. 5 Fabrication of bioactive glass containing wound dressings by 111 electrospinning
B 250 Ca K 225 Mg 200 P
175
150
125
100
75
Ion concentration (mg/L) concentration Ion 50
25 * 0 *
PCL Control Blend PCL-MH Blend-MH PCL-MH-B3 Blend-MH-B3
Figure 71: Results of ICP measurements on conditioned CCM containing ionic dissolution products of different fabricated fibers. One-way ANOVA statistical analysis denotes significant differences compared to pure CCM (*p<0.05).
5.3.5.2 Direct cell tests using human dermal fibroblasts (hDFs)
The effect of the produced electrospun fibre mats on cell attachment and proliferation was evaluated in contact with hDFs. Figure 72 shows fluorescence and SEM images of hDFs attached to the surface of the different fibre mats. In case of Blend-based fibre mats, hDF cells were well attached after 24 hours of cultivation and the amount of hDF cells seems to have increased after 7 days of cultivation. According to the SEM images, the surface of the electrospun Blend-based fibre mats is completely covered with a dense layer of hDF cells. Similar results could be obtained for PCL-MH and PCL-MH-B3 samples. However, a lower amount of hDF cells attached in the first 24 hours of cultivation on neat PCL fibre mats. Even after 7 days of cultivation, the surface seems not to be covered by cells. 112 5 Fabrication of bioactive glass containing wound dressings by electrospinning
Figure 72: Fluorescence and SEM images of hDFs grown on electrospun fiber mats for 1 and 7 days. Fluorescence images after 1d show low amount of hDFs attached on the PCL fiber mats, whereas the number of cells seems to increase by the addition of MH and B3 and by using Blend-based fiber mats. In accordance, fluorescence and SEM images after 7 days show dense cell layer grown on all kind of fiber mats, exceptional neat PCL fiber mats. 5 Fabrication of bioactive glass containing wound dressings by 113 electrospinning
To further quantify the results obtained in Figure 72, WST measurements were carried out. According to Figure 73, a clear difference between neat PCL-fibre mats and Blend-based fibre mats can be observed. The addition of MH to PCL fibres did not show a positive influence on the cell viability, but by additionally adding B3 particles to the PCL-MH fibre mats, a significant negative influence on the cell viability for both cultivation periods could be measured. In case of Blend based fibre mats, no significant influence of MH and B3 could be observed. However, this result could be due to the already excellent bioactivity of neat Blend- fibre mats, indicated by the dense layer of hDFs grown on the surface of the Blend-fibre mats (Figure 72).
Day 1 Day 7 * * *
* Cell viability (a.u.) viability Cell
* * * * PCL PCL-MH PCL-MH-B3 Blend Blend-MH Blend-MH-B3
Figure 73: Cell viability of hDFs grown directly on different electrospun fiber mats for 1 and 7 days. One-way ANOVA statistical analysis denotes significant differences compared to neat PCL fiber mats (*p<0.05).
5.3.5.3 In vitro scratch test (migration test) using human keratinocytes-like (Hacat) cells
In addition to cell attachment and proliferation, the impact of the different electrospun fibre mats on cell migration of Hacat cells were examined by performing an in vitro scratch test. The aim of this test is to observe the closing of the wound by the migration of Hacat cells along the edge of the created scratch to enable new cell-cell contact [252]. According to Figure 74, both types of electrospun fibre mats showed no influence on cell migration in comparison to the positive control (pure CCM). In case of PCL fibre mats, by the addition of MH, a positive 114 5 Fabrication of bioactive glass containing wound dressings by electrospinning influence on cell migration could be measured. In case of Blend fibre mats, already neat Blend fibre mats showed a positive effect on the migration of Hacat cells, which was shown to be significantly improved by the addition of MH. The addition of B3 BG particles does not seem to have a significant effect on both Blend-MH-B3 and PCL-MH-B3 fibre mats, compared to fibre mats just containing MH (Blend-MH and PCL-MH).
Control+ Control+ 140 Control- 140 Control- PCL Blend PCL-MH Blend-MH 120 120 PCL-MH-B3 Blend-MH-B3
100 100
80 80
60 60
Migration ratio (%) ratio Migration Migration ratio (%) ratio Migration 40 40
20 20
0 5 10 15 20 25 30 35 0 5 10 15 20 25 30 35 A) Time (h) B) Time (h)
Figure 74: Migration (presented by the migration ratio) of Hacat cells cultured in the presence of A) PCL based fiber mats and B) Blend based fiber mats as a function of time
5.4 Discussion
The potential of electrospun fibre mats as wound dressing materials was evaluated according to the required properties. Requirements for wound dressings (summarized in section 2.1.3) are divided in three main groups: practical, physical and biological requirements.
5.4.1 Practical properties
As described in section 2.3, by changing the electrospinning parameters (e.g. polymer concentration, applied voltage, humidity), the resulting properties of electrospun fibre mats can be manipulated. However, a disadvantage of the electrospinning technique could be the use of toxic solvents. Therefore, an increasing amount of research studies is focusing on the so-called “green electrospinning”, where benign solvents are used [250], [251], [269]. Due to its solubility in non-toxic solvents (e.g. water), MC is an ideal candidate for green electrospinning. By using the benign solvents acetic acid and formic acid, it was possible to 5 Fabrication of bioactive glass containing wound dressings by 115 electrospinning produce MC fibre mats with smooth fibre surfaces in the nanometre range. However, in this research it was not possible to obtain high quantities of electrospun MC fibre mats due to technical limitations (higher voltages would be necessary). Therefore, MC was blended with PCL to obtain higher quantities of material for realistic applications. PCL is well-known for its suitability for green electrospinning [114] and it was therefore chosen as second material in addition to MC. By combining MC with PCL, beadless fibre mats in high quantities with fibres in the higher nanometre range were successfully obtained. The blend fibre mats were found to be reproducible. However, the need of using high voltages (20kV) could be a limitation.
Moreover, it could be shown that by combining PCL-MC with MH, fibre mats with smooth fibres and without beads could be produced. In addition, BG particles could be homogenously added to the Blend-MH fibre mats (as well as to PCL-MH fibre mats). The obtained multicomponent fibre mats could be easily cut in different shapes, they were easily handled and therefore show high potential for application on a wound. Additionally, the fibre mats were successfully disinfected using UV-light (contaminations were never observed). Nevertheless, disinfection only might be not suitable for clinical use being necessary to prove the sterilizability of the fibre mats (e.g. by gamma radiation).
5.4.2 Physical properties
Besides meeting the practical requirements of a wound dressing, electrospun fibre mats need also to exhibit adequate physical properties. In general, an ideal dressing for the treatment of wounds should have a similar structure, scale and mechanical stability to the human skin. That means, fibre mats need to be porous with an average fibre diameter of 50-500 nm, which is comparable to native collagen bundles [289]–[292]. All produced fibres in this project are in the required range, whereas Blend-based fibres exhibited smaller fibre diameters compared to PCL-based blend fibres. Fibres based on silk fibroin [289], PCL [293], gelatine-PCL [294], PET [295] and PVA [296] showed average fibre diameter in the micromere scale. However, by combining PCL with chitosan, smaller fibre diameters could be obtained [297]. Moreover, in both cases, Blend- and PCL-based fibres, the addition of MH led to a (not significant) increase of the mean fibre diameter. This effect could be also already observed in literature [274], [293], [295], [296], [298]. Similar to the addition of MH, also the addition of BG particles led to a (not significant) increase of the average fibre diameter, in accordance to the literature [294], [299]. 116 5 Fabrication of bioactive glass containing wound dressings by electrospinning
In terms of protein adsorption and cellular adhesion/spreading, the surface properties of a possible wound dressing are crucial. Studies have proven that hydrophilic surfaces in contrast to hydrophobic surfaces improve cell adhesion, spreading and migration [300]– [302]. Our studies showed, by combining PCL with MC, that a more hydrophilic electrospun fibre mat surface can be created. Moreover, the wettability could be further improved by the addition of MH and BG (for both cases, Blend- and PCL-based fibres). In contrast, a previous study by Kadakia et al. [289] showed that by addition of MH the contact angle of pure silk fibroin fibre mats was increased from 61° ± 1° to 81° ± 1°. Another study performed by Liverani et al. [251] showed that the addition of BG had no effect on the wettability of the produced composite PCL fibre mats. However, in this research a different type and size of BG particles was used. Nevertheless, in contrast to results obtained in literature, it was possible to improve the wettability and therefore the surface properties of the Blend-based fibres by the addition of MH and microsized BG particles.
As mentioned above, an ideal wound dressing should have similar mechanical properties as human skin. By various in vitro and in vivo tests it could be shown that skin has an tensile strength of 5-30 MPa and an elongation at break in the range of 35 to 115% [303]. The pure Blend-fibre mats achieved a tensile strength of more than 40 MPa with a tensile strain of around 50%. These mechanical properties could be further improved by adding MH and BG, whereas the resulting tensile strain could be enhanced and the UTS reduced. Similar results could be achieved for PCL-based fibre mats, whereas the influence of doping with MH and BG was not as high as for Blend-based fibre mats (most probably due to the already lower UTS of neat PCL fibre mats). Moreover, similar effects of adding BG to polymer fibres have been reported in literature for silk-fibroin-carboxymethyl cellulose fibre mats (reducing initial UTS from 10 MPa to 5 MPa by adding 20% BG) [116] and PCL fibre mats (reducing initial UTS from 6 MPa to 3 MPa by adding 30% nanosized BG) [251]. In case of PCL-simvastatin fibre mats no significant influence of adding 20% BG particles could be found, most probably due to the already initial low UTS of neat PCL-simvastatin fibre mats of 2 MPa [304]. The reduction of the UTS (as well as other mechanical properties) could be explained by the weak adhesion of the BG particles to the used polymer [114], [251], [305].
Biodegradability of a wound dressing could be advantageous in order to avoid the regular change of the wound dressing, whereas the degradation rate should match the regeneration rate of the skin [108]. As described in section 2.4.2, the degradation of methylcellulose can be tailored by cross-linking. In this research, we could show that MC was successfully cross-linked with MH, whereas MH offers the advantage of being a non-toxic, 5 Fabrication of bioactive glass containing wound dressings by 117 electrospinning naturally crosslinker. Moreover, MH released from Blend-based fibres showed a favourable release rate, whereas MH released from PCL-based fibres showed a burst release. Similar results have been found in literature, when honey was incorporated in PCL fibres [293]. Besides the controlled release of MH, it was also shown that ions were released from the BG doped fibres, proving that the electrospun fibre mats are suitable platforms for dual drug delivery.
5.4.3 Biological properties
The addition of bioactive substances into scaffolds, such as MH and BG, as discussed above, can change the physical properties of the electrospun fibre mat and the cellular response to the wound dressing. Due to the addition of MH, it was expected to obtain an antibacterial effect. However, no antibacterial efficiency against E. coli and S. aureus was measured. It has been shown that at least 25 mg/ml of MH is necessary to obtain an antibacterial effect [259]. Therefore, it is expected that by increasing the content of MH in the fibres or by using a higher mass of fibre mat, an antibacterial effect can be achieved [259]. Similar results were obtained in a study by Yang et al. [298] using silk fibroin fibre mats. By using similar amounts of MH as used in this study, only a low antibacterial efficiency (efficiency of 6-12% against E. coli and 10-24% against S. aureus) was measured, whereas for higher concentrations of MH a greater antibacterial effect was obtained. This result was also confirmed by other studies using silk fibroin fibres [289].
Besides antibacterial properties, the addition of MH as well as BG should improve the bioactivity of the resulting electrospun fibre mat. As already discussed, the addition of MH and BG leads to a greater hydrophilicity of the surface of the fibres, which should improve cell adhesion and proliferation [300]–[302]. In accordance with the wettability of the different produced fibre mats, direct cell biology tests using hDFs showed a reduced cell viability of neat PCL fibres compared to neat Blend fibres. Moreover, the attachment and proliferation of hDFs was shown to be improved by the addition of MH and B3 to the PCL fibre mats, which is in accordance with the contact angle measurements. In case of Blend fibre mats, no significant influence of MH and B3 could be found. This might be due to the already good cytocompatibility of neat Blend fibre mats. Moreover, the ICP results showed that addition of B3 led to a significant higher amount of boron in the CCM. In contrast, indirect in vitro scratch tests using Hacat cells showed no negative impact of the neat PCL fibre mats for up to 24 h, most probably due to the fact that the test was done indirectly and no toxic substance was 118 5 Fabrication of bioactive glass containing wound dressings by electrospinning released from the PCL fibre mats. In accordance, none of the different fabricated PCL-based and Blend fibre mats showed any negative impact on cell migration. Moreover, the addition of MH and BG led in both cases, Blend and PCL-based fibre mats, to improved cell migration, which is favourable for wound healing applications.
Previous studies focusing on the effect of adding MH to electrospun fibre mats have shown different results that must be discussed in the context of the present findings. Although Arslan et al. [295] found in low concentrations of MH no cytotoxic effects of MH, high concentrations of MH added to polyethylene terephthalate fibre mats led to an altered morphology and confluency of mouse fibroblasts (L929 cells). In contrast, MH added to silk fibroin fibrous matrices was shown to improve cell viability of mouse fibroblasts (L929 cells) and in vivo tests showed even a positive effect on wound healing, which was comparable to the commercial available wound dressing AquacelAg dressing [298]. In another study, it was shown that adding honey to PCL fibres has an (initial) positive impact on the proliferation of macrophages and additionally a (not significant) improved wound closure in vivo was reported [293]. Based on these results, it is expected that a positive impact of PCL and Blend- based fibres containing MH can be achieved by using higher MH concentrations. In accordance, based on results from literature on PCL [305] and PCL-gelatine [116] fibres containing BG, an improved wound closure is expected by introducing either higher amounts of BG or by testing a larger mass of composite fibre mats, for example considering thicker mats. These results show that an optimization of the developed electrospun fibre mats is necessary. However, the results of this study indicate that by combining MC with PCL a new technology to develop electrospun wound dressings can be established and further improved by adding tailored concentrations of MH and BG particles. 6 Fabrication of bioactive glass containing wound dressings by 3D printing 119
6 Fabrication of bioactive glass containing wound dressings by 3D printing
6.1 Introduction
As described in section 2.3, 3D printing offers the possibility to fabricate complex 3D structures, which can be made patient specific [84, 93]. By combining the fabricated BGs described in Chapter 3 with MC and MH (section 2.4 and 2.5) and using 3D printing technology, a new group of wound healing scaffolds can be developed. In this Chapter the development as well as the characterization of 3D printed scaffolds based on MC and MH containing different amounts of B3 BG particles will be presented.
6.2 Experimental procedure
6.2.1 Scaffold production
MC and MH (the same compounds as described in section 4.2.1) and B3 BG (section 2.2) were used to produce the suspension (ink) for 3D printing. In order to produce the 3D printing ink, DW or Dulbecco’s phosphate-buffered saline (PBS, Gibco) were used as solvent. PBS or DW was heated up to 50°C to ensure a homogenous dispersion of MC and after reaching the temperature, different amounts of MC (5, 10 and 15 wt.%) were added to the solvent under constant stirring. In case of MC-MH solutions and MC-MH solutions containing B3 BG particles (preparation described in section 4.2.1), the desired amount of MH was added after stirring the MC solution for 15 min. After another 15 min of constant stirring, the desired amount of B3 BG was added to the MC-MH solutions. After stirring the suspension for 30 min at 50°C, it was stored in the fridge for 1h, 4h or overnight to allow the complete hydration of the MC based hydrogels, as suggested in previous work [306].
For the printing process, a three-axes movable bioplotter (BioScaffolder 3.0, GeSiM GmbH, Germany) equipped with a pressurized air system was used. Using the “ScaffoldGenerator” software, the shape and dimensions of the resulting scaffolds were defined. To perform printing, a syringe (which was connected to the plotter head) was filled 120 6 Fabrication of bioactive glass containing wound dressings by 3D printing with the different printing inks and a needle of 20G (diameter of 610 µm) was used. Printing inks were stored at room temperature at least for 1 hour and printing was performed at room temperature. During optimization, different pressures and plotting speeds were tried in order to obtain the optimal parameters of each solution. The optimized parameters including the composition and concentration of the different printing inks are summarized in Table 12.
Table 12: Summary of the concentrations and printing parameters for the different printing inks. All inks were based on PBS.
Sample MC MH B3 [wt.% respect Printing Printing [wt.%] [wt.%] to polymer pressure speed content] MC 10 550 kPa 2.0 mm/s MC-MH 10 5 550 kPa 1.6 mm/s MC-MH-10B3 10 5 10 550 kPa 1.6 mm/s MC-MH-20B3 10 5 20 550 kPa 2.5 mm/s MC-MH-40B3 10 5 40 550 kPa 4.5 mm/s
Using the optimized printing parameters, simple prismatic porous scaffolds based on square structures with an edge length of 10 mm, a height of 1.2 mm and a pore size of 0.9 mm were printed. In order to achieve a height of 1.2 mm, 4 layers with a height of 0.3 mm per layer were printed. The printing process and the resulting shape of the scaffolds can be seen in Figure 75. For compression tests, scaffolds with an edge length of 5 mm and a height of 3 mm were printed. After the printing process, all scaffolds were stabilized for 1h at 37°C in air (no significant shrinkage was observed).
Figure 75: Images showing the printing process of a MC-MH scaffold on top of a well plate (left) or inside the well plate (right) 6 Fabrication of bioactive glass containing wound dressings by 3D printing 121
6.2.2 Characterization
6.2.2.1 Optical microscope images
A light microscope (Primo Vert, Carl Zeiss) was used to examine the morphology of different printed scaffolds and to evaluate the printing ink. Based on light microscope images, the diameter of 10 struts from each scaffold and the pore side length of each scaffold were calculated using ImageJ (NIH, Bethesda, MD).
6.2.2.2 Fourier transform infrared spectroscopy (FTIR)
FTIR was performed on composite 3D printed scaffolds in order to examine the chemical structure of the composites and to evaluate the presence of the B3 BG particles inside the printed structure (according to section 4.2.2.3).
6.2.2.3 Mechanical testing
Using a universal testing machine (some instrument described in section 4.2.2.4), the compression strength of the produced 3D printed structures was assessed. Samples with an edge length of 5 mm and a height of 3 mm were printed as described in section 6.2.1. At least 10 replicates were tested using a speed of 10 mm/min and a load cell of 100N.
6.2.2.4 Contact angle measurement
In order to evaluate the wettability of the 3D printed scaffolds, the contact angle can be measured by depositing a water droplet on a surface as described in section 4.2.2.5. At least 5 replicates were measured using a 3 µl water drop. For contact angle measurements, samples without porosity were printed using the same printing parameters as described in Table 12. Dense samples with a total height of 0.6 mm (2 layers á 0.3 mm each) and 10 mm side length were created.
6.2.3 Swelling behaviour and change of weight of the 3D printed scaffolds
In order to test the stability of the printed scaffolds, swelling and degradation studies using DW and DMEM were performed. Both tests were carried out in triplicates. 122 6 Fabrication of bioactive glass containing wound dressings by 3D printing
6.2.3.1 Swelling and change of weight under “dry” conditions in DW
To test the samples under “dry” conditions (to stimulate dry wounds), scaffolds were immersed in 1ml of DW for up to 7 days. Prior immersion, all samples were weighed using an analytical balance (BM-252, A&D). After different time points (t=1h, 3h, 6h, 24h and 72h), excess of DW was removed and the weights of the scaffolds were recorded. Afterwards the desired excess of DW was added again. Weight change was calculated based on the following formula:
푤 −푤 푤푒𝑖𝑔ℎ푡 푐ℎ푎푛𝑔푒 = 푡 푡0 × 100 (4) 푤0
where 푤0 is the specimen weight at time point 0 and 푤푡 at time t.
6.2.3.2 Swelling and change of weight under “wet” conditions
To test the samples under “wet” conditions (to simulate wet wounds), scaffolds were immersed in 5ml of DMEM for up to 168h. The test was performed in accordance to the test described in section 6.2.3.1. The weight of the scaffolds was recorded after 1h, 3h, 6h, 24h, 72h and 168h.
6.2.4 In vitro direct cell tests using hDFs
Samples having the same geometry as described in section 6.2.2.4 were disinfected for 1h under UV light. To prove the cytocompatibility, the 3D printed sheets were tested in direct contact with hDFs (section 4.2.5 and 5.2.5). Therefore, 100,000 hDFs were directly seeded on top of each sample using a 50 µl CCM drop (procedure as described in section 5.2.5.1). Then, the samples (located in a 24-well-plate) were incubated for 15 min and subsequently filled with 1ml CCM. Samples with cells seeded on top were then incubated for 24h and 7d. Following this, the cell viability was assessed by a mitochondrial activity calorimetric assay, which is based on tetrazolium salt (WST-8 assay, Sigma-Aldrich). Additionally, the morphology of the hDF cells was evaluated using staining with calcein and 4′,6-diamidino-2- phenylindole (DAPI) [121], [171] as described in section . Viable cells with intact cell membranes are able to use nonspecific cytosolic esterases to convert nonfluorescent calcein- AM into bright green-fluorescent calcein. Therefore, calcein stains the cell membrane, whereas DAPI binds to the Adenine and Thymine bases within the DNA of the cell nuclei and therefore it stains the cell nuclei. Images were taken with a fluorescence microscope (Axio Observer D1). 6 Fabrication of bioactive glass containing wound dressings by 3D printing 123
6.3 Results
6.3.1 Optimization of printing ink and printing parameters
In order to fabricate MC-based scaffolds by 3D printing, inks containing DW as solvent combined with different MC concentrations were prepared and stored in a fridge at 4 °C for different time periods to allow complete hydration. Based on previous studies [307], concentrations of 5 wt.%, 10 wt.% and 15 wt.% were investigated. Solutions (of each concentration) stored for 1h and 4h in the fridge at 4 °C were shown to be still not completely hydrated (at 4°C, hydration takes place by hydrophilic interactions between the solvent (DW) and the –OH groups of the MC [308]), which could be observed by the formation of two different phases (a liquid, clear phase and a wax-like, opaque phase). Therefore, solutions stored in the fridge overnight were only used for further experiments. By printing simple structures (4 layers, length of 10 mm, 900 µm distance between struts), the 10 wt.% concentration was shown to give the best printing results. In case of lower concentrations, the printed strands were not stable and no 3D structure could be achieved. On the contrary, the viscosity of the 15 wt.% MC in DW was too high and no continuous struts could be obtained (Figure 76).
Figure 76: Printing optimization: whereas high amounts of MC lead to the formation of not continuous struts, low amounts lead to the merging of strands. 10 wt.% MC scaffolds showed continuous struts formation and a stable 3D structure.
Next, we investigated the effect of changing the solvent. Since the 3D printed scaffolds should be ideally (for a certain time) stable in the human body, a gel formation temperature 124 6 Fabrication of bioactive glass containing wound dressings by 3D printing below 37°C needs to be obtained. Since it is well known that the addition of salt leads to a reduction of the gel formation temperature [309]–[315], PBS was chosen as solvent. The printability of the MC-PBS solutions was not significantly changed compared to MC-DW solutions. In order to evaluate the effect of changing the solvent on the gel formation temperature, MC-DW and MC-PBS structures (dimension as described above) were immersed in DW at 37°C for up to 24h. Whereas MC-PBS scaffolds showed swelling after immersion in DW for 24h, MC-DW structures showed degradation after the first contact with DW. After 24h, the MC-DW structures were completely destroyed (Figure 77). To further improve the stability of the printed structures, MC-PBS samples were stored at 37 °C for 1 hour to promote the formation of a physical hydrogel based structure [306].
Figure 77: MC-DW and MC-PBS samples after immersion in DW for 24h at 37 °C. MC-DW structures start to degrade/dissolve, whereas the structure of MC-PBS structures was intact.
By using the 10 wt.% MC ink (dissolved in PBS), fabricated and stored using the optimized parameters, the printing process was optimized. Several parameters influence the printability of hydrogels, e.g. temperature, printing speed, needle diameter, viscosity, etc. Based on a previous study [137], [306], room temperature (around 21°C) was chosen as printing temperature. According to Negrini et al. [306], lower temperatures lead to a too low viscosity of the ink resulting in an accumulation during printing and higher temperatures lead to a too high viscosity resulting in an in-homogenous printing with voids and irregularities [306]. Also here, good printability was observed at room temperature; therefore no further optimization of the temperature was required. By printing more complex structures, it could be shown that a relative high printing pressure of 500 kPa and relative low printing speed of 1mm/min are necessary to obtain stable structures. However, 3D scaffolds suitable for mechanical testing (providing a height of at least 3 mm) could not be obtained. Therefore, for further investigations neat MC scaffolds were not used. 6 Fabrication of bioactive glass containing wound dressings by 3D printing 125
Based on the optimized parameters identified to fabricate and print MC-based inks, MC solutions containing additionally MH were developed. In order to determine the optimal concentration of MH, 10 wt.% MC solutions containing MH in a ratio of MC:MH of 1:1, 2:1 and 4:1 were prepared. Solutions containing 10 wt.% of MC and 10 wt.% of MH were not printable, since the addition of MH had a softening effect and no continuous, stable strands could be obtained. Both 10 wt.% MC inks, containing additionally 5 or 2.5 wt.% of MH showed good printability. Therefore, the solution containing MC and MH in a ratio 2:1 was chosen to have the highest possible amount of MH in the printed structures. To additionally investigate the effect of adding B3 BG particles, MC-MH solutions containing 10 wt.%, 20 wt.% and 40 wt.% of B3 BG particles in respect to the polymer content were fabricated using the same parameters as described for the MC-solution fabrication. To obtain scaffolds with desired dimensions, the printing pressure and speed of the four MC-MH solutions containing 0, 10, 20 and 40 wt.% of B3 BG particles inclusions were separately optimized, as summarized in Table 12. Scaffolds fabricated using these parameters (Figure 78) were used for further experiments.
Figure 78: Images showing 3D printed scaffolds made from MC-MH-20B3 ink
6.3.2 Characterization of optimized 3D printed scaffolds
In order to characterize the 3D printed scaffolds, light microscope images of the different printed samples were obtained. As shown in
Figure 79, the struts of the MC-MH scaffolds are transparent and no clear inclusions can be detected, which proves that a homogenous hydrogel composed of MC and MH could be obtained. Moreover, with increasing content of B3 BG particles inside the MC-MH hydrogel, the struts get more dense and less transparent. In case of 10 wt.% of B3 BG, the thickness of the struts seems not to be uniform and some accumulation of the hydrogel occurs at the crossing points. With increasing amount of B3 BG particles inside the MC-MH scaffolds, the struts get ticker and more straight. Therefore, it seems that the printing accuracy was improved by adding B3 BG particles into the MC-MH hydrogel. 126 6 Fabrication of bioactive glass containing wound dressings by 3D printing
Figure 79: Light microscope images of 3D printed scaffolds showing indicating printability changes as function of BG content In order to prove these observations more precisely, the light microscope images were analysed using the ImageJ software. As shown in Figure 80, the initial introduction of B3 particles seems to reduce the mean thickness of the printed struts, whereas with increasing amount of particles, the mean thickness increased as well. In accordance, the mean side length of the pores was found to be the highest for the MC-MH-10B3 scaffold and decreased with increasing B3 BG content. Moreover, the assumption that the printing accuracy increases with increasing B3 BG content could be proven, as indicated by the decreasing error bar with increasing B3 BG content.
1200 600 #
1000 500
800 m)
400 m)
600
300 Thickness ( Thickness 200 ( Thickness 400
100 200
0 0 MC-MH MC-MH-10B3 MC-MH-20B3 MC-MH-40B3 MC-MH MC-MH-10B3 MC-MH-20B3 MC-MH-40B3 Mean thickness of printed struts Mean side length of pores
Figure 80: Mean thickness of the printed struts and mean side length of the pores of the printed MC-MH based scaffolds based on light microscope images. One-way ANOVA statistical analysis denotes significant differences between all samples (*p<0.05), only exception is marked with #. 6 Fabrication of bioactive glass containing wound dressings by 3D printing 127
The fabricated 3D printed scaffolds were further characterized using FTIR. As shown in Figure 81, the typical bands of MC can be found in the pure MC scaffolds. Briefly, at around 1150 cm-1, the stretching vibrations of the C-O oxygen bridge and at 1050 cm-1, the C-O-C stretching mode of the glycosidic unit can be found. Moreover, the peak at 940 cm-1 can be attributed to the vibration of OCH3 groups [255]–[257]. The spectra of MC-MH scaffolds additionally indicate the presence of MH by several peaks in the range of 750 cm-1 to 1500 cm-1, which can be related to the main component of MH: sugar [258], [259]. Although, the presence of B3 BG particles could not be clearly detected, it can be observed that the peaks related to MH and MC are not as clearly visible as in scaffolds not containing any B3 BG. Especially in the range of 900 cm-1 to 1100 cm-1, the intensity of the reported peaks decreased and the peaks appear more broad. This might be due to the presence of absorption bands in this range related to B-O stretching of tetrahedral BO4 units in the BG [78], [196].
OCH3 C-O-C C-O MC-MH-30B3
MC-MH-20B3
MC-MH-10B3
Absorbance (a.u.) Absorbance MC-MH
MC
400 600 800 1000 1200 1400 1600 Wavenumber (cm-1)
Figure 81: FTIR spectra of 3D printed scaffolds based on different printing solution containing MC, MH and different amounts of B3 BG particles. The peaks are discussed in the text.
The mechanical properties of the 3D printed scaffolds were assessed by compression strength tests. Figure 82 shows exemplary typical stress-strain curves of the different 3D printed scaffolds. Interestingly, a clear increase of the overall compressive strength of MC-MH foams containing 10 wt.% B3 BG can be observed, whereas the addition of 20 wt.% B3 BG seems not to have any significant influence on the strength. Moreover, MC-MH-40B3 scaffolds offer the highest compressive strength during the first 80% of compressive strain. 128 6 Fabrication of bioactive glass containing wound dressings by 3D printing
Additionally, in contrast to MC-MH, MC-MH-10B3 and MC-MH-20B3 scaffolds (showing an exponential increase of compressive strength), MC-MH-40B3 samples exhibit a more linear increase of compressive strength.
35 MC-MH MC-MH-10B3 30 MC-MH-20B3 MC-MH-40B3 25
20
15
10 Compressive stress (kPa) stress Compressive 5
0
0 20 40 60 80 100 Compressive strain (%)
Figure 82: Typical stress-strain curves of 3D printed MC-MH scaffolds containing different amounts of B3 BG particles. Mean compressive strength was calculated based on the measured compressive stress in the marked area (70-80% of compressive strain).
Based on the stress-strain curves, the mean compressive strength in the range of 70 to 80 % (grey area in Figure 82) of the MC-MH based scaffolds containing different amounts of B3 BG was calculated (Figure 83 A). In accordance to the stress-strain curves, no significant difference between the mean compressive strength of MC-MH (5±1 kPa) and MC-MH-20B3 (7±1 kPa) could be found (only a minor enhancement of the compressive strength). Moreover, the MC-MH-10B3 scaffold showed a significantly increased mean compressive strength (11±1 kPa), which is more than the double compared to the MC-MH scaffold. In accordance, the MC- MH-40B3 scaffolds showed a slightly higher enhancement of the mean compressive strength than the MC-MH scaffold containing just 10 wt.% of B3 BG, a value of 15±4 kPa could be measured. However, here it is important to mention that the error bar increased significantly for the MC-MH-40B3 scaffold. 6 Fabrication of bioactive glass containing wound dressings by 3D printing 129
* 100 25 *
* 80 20 * * 60 15
10 40 Contact angle (°) angle Contact
5 20 Mean compressive strength (kPa) strength compressive Mean
0 0 A) MC-MH MC-MH-10B3 MC-MH-20B3 MC-MH-40B3 B) MC-MH MC-MH-10B3 MC-MH-20B3 MC-MH-40B3
Figure 83: Mean compressive strength (based on the measured compressive stress in the range of 70- 80% of compressive strain) and B) contact angle of the different 3D printed scaffolds
In addition, the contact angle of printed dense samples was measured to evaluate the effect of the addition of B3 BG. As shown in Figure 83 B, no significant difference between the contact angle of MC-MH samples and of MC-MH samples containing 10 and 20 wt.% of B3 BG could be measured. The results of all three samples lie in the range of 60-80°. The increased error bar of the results for the MC-MH-10B3 sample is in accordance with the found inaccuracy during the printing process, leading also to a higher roughness (which in turn influences the contact angle). Moreover, the MC-MH sample containing 40 wt.% B3 BG showed a significant reduction of the contact angle to a value of 35 ± 8°. It seems that B3 BG particles in low concentration are completely embedded in the polymer matrix and therefore they do not affect the contact angle, whereas for higher concentrations of B3 BG, a favourable reduction of contact angle was achieved. Figure 84 shows typical water drops on top of the different samples during contact angle measurements, whereas a clear difference between the MC-MH-40B3 sample and the remaining samples can be observed. 130 6 Fabrication of bioactive glass containing wound dressings by 3D printing
Figure 84: Typical water drops on top of different 3D printed MC-MH samples containing different amounts of B3 BG particles during contact angle measurement.
6.3.3 Swelling behaviour and change of weight of 3D printed scaffolds under different conditions
Based on the fact that moist/wet wound treatment is known to achieve better wound healing results (e.g. less inflammatory reaction and reduced scar formation) than wound treatments, which are not able to keep a moist environment [316], it is important to measure the swelling behaviour of the 3D printed scaffolds and therefore the ability of keeping a moist environment. The amount of aqueous solution, which can be retained by the hydrogel-based 3D printed scaffold, depends on the structure of the hydrogel network, but also on the environmental conditions (temperature, pH, ionic strength of water solution) [317]. Therefore, two different dissolution media (DW and DMEM) were used to test the swelling behaviour of the 3D printed scaffolds under different conditions (excess or limitation of dissolution medium).
6 Fabrication of bioactive glass containing wound dressings by 3D printing 131
6.3.3.1 Swelling and change of weight under wet conditions
2750 MC-MH 2500 MC-MH-10B3 MC-MH-20B3 2250 MC-MH-40B3 2000 using wet conditions
1750
1500
1250
1000
750
500 Relative change of weight (%) weight of change Relative 250
0 0 20 40 60 80 100 120 140 160 180 Time (h)
Figure 85: Relative change of weight as a function of immersion time in DMEM under wet conditions. All samples were immersed in 5ml of DMEM and the weight was recorded without excess of DMEM.
In order to analyse the swelling behaviour, the 3D printed scaffolds were immersed in 5 ml of DMEM. According to Figure 85, all MC-MH scaffolds containing different amounts of BG particles showed an initial fast increase of weight. In case of MC-MH and MC-MH scaffolds containing 10 and 20 wt.% of B3 BG, the maximum of weight was reached after 3 hours, achieving around 2000 % of the original weight for the three kinds of samples. Afterwards, the weight decreased slowly with immersion time reaching an end weight of around 200% for MC-MH, MC-MH-10B3 and MC-MH-20B3. Although after 7 days almost the same relative change of weight was reached, the swelling (increase and decrease of weight) was different for the three kinds of samples. No obvious difference between MC-MH and MC-MH-20B3 samples could be found, whereas it seems that MC-MH-10B3 scaffolds can keep higher amounts of DMEM for longer time. In contrast to MC-MH, MC-MH-10B3 and MC-MH-20B3 scaffolds, the swelling profile of MC-MH-40B3 scaffold is different. The increase of weight (and resulting swelling) of the MC-MH-40B3 scaffold was lower compared to the other samples, reaching its maximum of around 300% after 24 hours. Afterwards, the weight decreased slightly until the third day of immersion, followed by a defined decrease, reaching a final weight of around 40% of the original weight. Therefore, the MC-MH-40B3 scaffold seems to degrade. However, the reduction of weight could be also explained by the dissolution of BG particles within the MC-MH-40B3 scaffold. 132 6 Fabrication of bioactive glass containing wound dressings by 3D printing
6.3.3.2 Swelling and change of weight under dry conditions
2000 MC-MH 1750 MC-MH-10B3 MC-MH-20B3 1500 MC-MH-40B3 using dry conditions 1250
1000
750
500 Relative change of weight (%) weight of change Relative 250
0 0 10 20 30 40 50 60 70 80 Time (h)
Figure 86: Relative change of weight as a function of immersion time in DW under dry conditions. All samples were immersed in 1ml of DW and weights were recorded without excess of DW.
As mentioned above, wound dressings need to be able to keep a moist environment in case of dry wounds. Therefore, in addition to the immersion of the 3D printed scaffolds in higher amounts of liquids, the scaffolds were also brought in contact with a minor amount of DW. Under “dry” conditions, the relative changes of weight are different to those in “wet” conditions. As shown in Figure 86, MC-MH scaffolds showed the highest uptake of water (max. weight of around 1900%), followed by MC-MH-20B3 scaffolds (max. weight of around 1600%). Although under “wet” conditions MC-MH-10B3 scaffolds exhibited the highest maximal weight change, under “dry” conditions a maximal weight increase of around 1300% was reached. In contrast, the maximum weight increase of MC-MH-40B3 scaffolds under “dry” conditions was found to be similar to the weight change measured under “wet” conditions (around 260%). Interestingly, the weight change profile of all printed samples is similar, characterized by an initial fast increase in the first 3-6 hours, followed by fast decrease until 24 hours. After 24h, all samples slowly decreased in weight, resulting in degradation with relative weights of around 90% for MC-MH, MC-MH-10B3 and MC-MH-20B3 as well as a relative weight of around 50% for MC-MH-40B3 scaffolds. 6 Fabrication of bioactive glass containing wound dressings by 3D printing 133
6.3.4 Preliminary cell compatibility of 3D printed scaffolds
Figure 87: Fluorescence images of hDFs grown on 3D printed MC-MH based scaffolds containing different amounts of B3 BG after 24 hours of culture
In order to evaluate the cytocompatibility of the 3D printed MC-MH based scaffolds, a direct cell test using hDFs was performed. Figure 87 shows fluorescence images of hDFs attached to the surface of the 3D printed scaffolds. No obvious qualitative differences between the fluorescence images could be found, all 3D printed scaffolds showed a good cytocompatibility. 134 6 Fabrication of bioactive glass containing wound dressings by 3D printing
* * 120 *
100
80
60
40 Relative cell viability (%) viability cell Relative
20
0
MC MC-MH Control MC-MH-10B3 MC-MH-20B3 MC-MH-40B3
Figure 88: Cell viability of hDFs grown directly on different 3D printed MC-MH based scaffolds containing different amounts of B3 BG after 24h of culture. One-way ANOVA statistical analysis denotes significant differences (*p<0.05).
In order to quantify the cytocompatibility of the different 3D printed MC-MH scaffolds, WST measurements were carried out. In accordance to results obtained by fluorescence images, none of the tested samples showed great impact on cell viability. However, a significant reduction of the viability of hDFs seeded on MC scaffolds compared to the control could be measured. Moreover, MC-MH samples containing additionally MH showed a significant improvement of cell viability compared to pure MC samples. Therefore, it could be concluded that the addition of MH improves the cytocompatibility of the samples. Similar effects could be observed by the addition of 10 wt. % of B3 BG. However, with increasing B3 BG content, the cell viability appears to be slightly (not significantly) reduced, compared to the control and the MC-MH scaffold.
6.4 Discussion
6.4.1 MC as printing material
MC is well known for 3D printing of support materials (used to produce more complex scaffolds made out of non-water soluble materials), which can be easily removed by using water [318], [319]. Moreover, MC is already widely used in the area of tissue engineering to 6 Fabrication of bioactive glass containing wound dressings by 3D printing 135 produce cell-sheets [137], [318], [320]. However, cell-sheets are only able to replicate the shape of the form, in which they were fabricated and can therefore not be fabricated in complex shapes. To overcome this drawback, knowledge on how to produce stable cell-sheets at physiological temperatures can be combined with 3D printing to obtain stable hydrogel- scaffolds [306]. So far, MC alone has not been often used for 3D printing, more research is focusing on MC-Blends (e.g. with alginate [321]–[324], agarose [325], hydroxypropyl methylcellulose [318], hyaluronic acid [326] and self-assembled peptide [327]). There are only a few studies reporting the 3D printing of pure MC [137], [306].
In order to use MC-based hydrogels for 3D printing, the printability of the neat MC hydrogel needs to be evaluated. As described in section 6.3.1, MC solutions based on DW showed a good printability (after optimization). However, they were not suitable for the fabrication of 3D scaffolds for wound healing applications, since they showed no stability in aqueous environment at physiological conditions. Therefore, in accordance to previously published studies [137], [306], a salt-containing solution was used for further experiments.
As mentioned in section 6.3.1, the addition of salt leads to the reduction of the gel formation temperature [309]–[315]. For instance, Cl- ions and SO4- ions (deriving for instance from salts of NaCl, KCl or Na2SO4) induce water molecule spillage from the polymeric structure [137], [328]–[330]. The water molecules place themselves around the used salts and reduce thereby the intermolecular hydrogen-bond formations between the hydroxyl groups of the MC and the water molecules. This in turn leads to an increase of the hydrophobic interactions between MC molecules, therefore inducing a decreased gelation temperature [307], [317], [331]–[333]. An ideal scenario would be that, by the addition of salt at temperatures below
37°C, hydrophobic groups of MC (-OCH3) are predominant and lead to the formation of a physical cross-linked structure (in contrast to the predominance of hydrophilic groups of MC (-OH) at 20°C) [307], [317], [331], [334]. Therefore, by increasing/decreasing the temperature, the hydrophobicity of the hydrogel and therefore the gel formation ability can be tuned. Moreover, the gel formation is time reversible since no covalent bonds are involved [139], [335]. This is a great advantage and gives the opportunity to play with salt concentrations in order to achieve the optimal gel formation temperature.
Although higher concentrations of salt will lead to lower gelation temperatures (which would be preferred option here), a too high concentration of salt can be risky for cells. The osmolarity (concentrations of salts) should not be higher than 300 mOsm, otherwise salt content could have a negative influence on cell viability [307]. Therefore, in contrast to literature [137], [306], here one-fold PBS was used. The used PBS has an osmolarity of 270 to 136 6 Fabrication of bioactive glass containing wound dressings by 3D printing
300 mOsm per L, which does not exceed the suggested level. Although it was possible to print 3D scaffolds by using the optimized MC-PBS solution, it was not possible to print higher scaffolds (3 mm), which do not collapse under their own weight. Therefore, the salt content should be further increased to obtain stable and high 3D scaffolds. Since the goal here was not to print neat MC scaffolds and it was hypothesized that by adding MH (and achieving a cross- linking effect), the printability could be further improved, the effect of increasing the salt concentration on the printability (and cell viability) was not further investigated. However, in the future this approach should be considered in order to obtain pure MC 3D scaffolds with optimum properties.
6.4.2 Effect of addition of MH and B3 BG
As expected, by adding MH to the MC-based scaffolds (in a ratio 1:2), it was possible to improve the printability of 3D scaffolds. Based on the optimized 3D printed MC-MH scaffolds, the impact of adding B3 BG particles was examined. The initial addition of B3 BG was shown to lead to the impairment of the printability of the scaffolds, but with increasing amount of B3 BG, the printability could be improved compared to the MC-MH scaffolds. However, here it is important to mention that the final scaffolds need to be further investigated (e.g. by µCT) in order to evaluate in detail the porosity structure.
Hydrogels in general are 3D networks of crosslinked polymeric chains and can be separated into chemically and physically crosslinked materials [336]. As shown in section 4.3.1, the addition of MH into the MC hydrogel leads to the formation of covalent bonds: MH acts therefore as chemical crosslinker. In contrast, the addition of salt was shown to be a physical crosslinker, characterized by its reversible nature. Additionally, at the interface between the B3 BG particles and the MC-MH hydrogel, physical crosslinks, for instance in the form of ion-bonding, can occur [336]. Since the amount of MH was not changed and therefore the degree of chemical crosslinking was kept constant, only the change of the extent of physical cross-linking due to the addition of B3 BG particles needs to be further considered. In general, there are three ways how BG particles can interact with the polymeric matrix:
1) no specific interaction
2) specific physical interactions take place leading to an enhancement of crosslinking density (physical cross-linking) 6 Fabrication of bioactive glass containing wound dressings by 3D printing 137
3) specific physical interactions lead to the adsorption of polymer molecules on the surface, which could result in the disturbance of the polymeric network structure (decrease of cross-linking).
The kind of interaction between the B3 BG particles and the polymeric MC-MH matrix is strongly related to the amount of BG incorporated [336]–[338]. The pure 3D printed MC-MH scaffold exhibited compressive strengths of around 5 kPa, which is lower but in a similar range to the reported values of e.g. hyaluronic acid/MC composite [326] and alginate/MC composites scaffolds [324] (both in the kPa range). Moreover, the initial addition of B3 BG particles leads to a significant increase of compression strength (about 11 kPa), followed by a decrease with increasing B3 BG content (about 7 kPa). The highest compressive strength could be found for MC-MH-40B3 scaffolds (about 15 kPa). These results correlate well with the fact that the amount of filler is crucial regarding the interface reactions. Similar results were found by Luo et al. [339], where BG was added to sodium alginate. In this study, the authors found that initially an increase of compression strength could be achieved by adding 13-93 BG particles, followed (after reaching a peak at a filler content of 1:2) by a decrease of compressive strength [339]. In contrast, Killion et al. [340] found a steady increase of compressive strength with increasing silicate sol-gel BG content in PEGDMA matrices [340]. The conflicting results indicate that more in depth studies are necessary to understand the interface-reactions in dependence on the filler content.
In theory, the incorporation of bioactive inorganic fillers as B3 BG particles can have a similar effect on the hydrophilicity of the composite and therefore also on the water adsorption capability of the polymer matrix. For instance, the generation of interfaces between the particles and the matrix can allow the aqueous media to penetrate easily into the composite [341]. As expected, the addition of B3 BG particles to the MC-MH scaffolds leads to a significant reduction of the contact angle (from around 70° to 35°). In accordance, the swelling behaviour of the 3D printed scaffolds also changed with increasing B3 BG content. The greatest effects could be observed for MC-MH-40B3 (showing a lower swelling degree). However, although the swelling degree was reduced for 40 wt.% of B3 BG, similar results (around 200 % swelling degree) could be found for pure 8 wt.% MC scaffolds [137], [290]. Here it is also important to mention that the results in literature cannot be directly compared to the present results, since different test conditions (amount of dissolution media, amount of sample, measuring time points, etc.) were used [137], [320]. Interestingly, Contessi et al. [320] found, by testing the swelling behaviour of 8 wt.% MC scaffolds in DW and DMEM, that MC scaffolds are more stable in water than in DMEM. However, in higher salt concentrations were used compared to this study and therefore the different osmolality of the 3D printed scaffolds 138 6 Fabrication of bioactive glass containing wound dressings by 3D printing seems to play an important role [320]. Since the osmolality of the 3D printed MC-MH scaffolds (produced by using PBS) is similar to the osmolality of DMEM, it is most likely that no salt is leached out of the scaffolds. This behaviour might explain the greater stability of the printed scaffolds in DMEM compared to DW (where it is more likely that the salt is leached out).
6.4.3 Suitability of MC-MH based hydrogels containing B3 BG as bioinks for biofabrication
The ECM is the natural environment of cells consisting of a 3D fibrous network embedded in a gel-like matrix [336]. Therefore, instead of fabricated 3D scaffolds (in form of foams, fibre mats or 3D printed scaffolds) where cells are “just” seeded on top, increasing research efforts are focusing on embedding cells into hydrogels for fabrication of tissue-like constructs [342], [343]. One possible approach is here biofabrication, where cells are incorporated in a bioink, which is used to 3D print scaffolds of desired shape [344]. Prior to using the developed MC-MH based hydrogels as bioink, the material cytocompatibility must be evaluated. Here, only a slight reduction of cell viability was determined for neat MC scaffolds. In order to evaluate the possible toxic effects of the salt, Altomare et al. [330] tested 4 wt.% MC hydrogels containing different amounts of salts and showed that even higher amounts of salt than the amounts used here were not toxic to fibroblast (L929) cells [330]. In addition, Contessi et al. [320] conducted a preliminarily bioprinting study by 3D printing muscle myoblast cells (C2C12) within a 8 wt.% MC hydrogel. The results showed that cells were well distributed within the printed structure and showed good viability [320]. Therefore, based on these studies and the cytocompatibility results obtained here, the MC-MH hydrogels containing different amounts of B3 BG particles are proposed as attractive candidates for developing bioinks for biofabrication. 7 Summary and Outlook 139
7 Summary and Outlook
7.1 Concluding remarks
The research carried out in the framework of this doctoral thesis has focused on providing a contribution to tackling the current need in developing alternative multifunctional wound dressings, especially for the treatment of chronic and infected wounds [4], [7], [10]. Within this thesis, composite wound dressings were fabricated by combining newly developed antibacterial and angiogenic BGs with MC and MH by using three different processing techniques, namely freeze drying, electrospinning and 3D printing.
In the first part of this thesis, based on the well-known silicate 13-93 BG composition [54], [75], a series of borate and borosilicate BGs doped additionally with copper and/or zinc were fabricated and characterized. Based on the structural, thermal and chemical analysis of the fabricated new BG compositions, several observations in relation to the used network formers silica and boron trioxide could be made. As expected, the (partial) replacement of silica with boron trioxide leads to an increase in dissolution rate and the dissolution behaviour of the three investigated of BGs was confirmed to strongly depend on the used dissolution media. Although the results obtained by NMR, FTIR spectroscopy and DTA measurements correlate well with the results obtained by several dissolution tests using different dissolution media, interpretation of results should be done carefully. In the past, several studies were performed in order to understand the complex dissolution behaviour of BGs, however these studies were mainly focused on silicate BGs [47], [225], [345]–[347]. The dissolution process of borate and borosilicate BGs is different to that of silicate BGs and several mechanisms influence the complex dissolution process of such multi-component glasses [81], [288]. For instance, as shown in Chapter 3, the used of dissolution medium (e.g. SBF, TRIS, LA) has different effects on the dissolution behaviour of the tested silicate, borate and borosilicate BGs.
In addition to the observations made by changing the network former, it was confirmed that the replacement of calcium by copper and/or zinc had a great influence on the BG properties. Based on the variance between the charge-to-size ratio of calcium and copper/zinc [227], [228], the structure and resulting dissolution behaviour of ion-doped BGs were found to be different compared to the undoped borate and borosilicate BGs. However, as 140 7 Summary and Outlook mentioned above, although the results obtained by the different analyses performed correlated well with the results obtained by dissolution experiments, such observations are considered to be preliminary indications and need further evaluation (e.g. by NMR and Raman analysis of a series of borosilicate BGs).
Besides the basic characterization of the developed BGs and the evaluation of their dissolution behaviour, the relevant biological properties of the BGs were in focus of this research. As shown by testing the BGs in contact with two bacteria types, the gram-positive bacterium S. aureus and the gram-negative bacterium E. coli, the addition of zinc and especially copper led to an improved antibacterial efficiency of the borate and borosilicate BGs. Moreover, due to their faster dissolution rate and therefore high suitability for wound healing applications, the series of borate BGs containing copper and/or zinc were tested in contact with immune cells. Surprisingly, this was the first time that BGs have been tested in contact with immune cells, although the immune system is strongly involved in every regeneration process in the human body [240], [242], [243], [348]. DCs for instance are known to interact with biomaterials [241] and thus, in future, every biomaterial considered for applications in the human body should be tested in terms of its influence on the immune system. The results obtained here indicate a dose-dependent effect of the different ions released by the borate BGs on DC behaviour. Depending on the BG composition and concentration, the released ions were shown to induce an immune reaction (e.g. in case of B3- Zn BG) or to lead to an immune tolerance by inactivation of the DCs (e.g. in case of B3-Cu BG). This effect in turn was shown to influence the ability of DCs to stimulate T-cell proliferation, indicating the significant influence that BGs could have on the immune system, which remains unexplored.
Therefore, the first research question asked in the introduction “Are the developed ion-doped BGs suitable for applications in wound healing?” can be answered: the results obtained in Chapter 3 indicate the suitability of borate (and borosilicate) BGs doped with copper and/or zinc. Although the ion-doped BGs were found to be antibacterial and the immune reaction can be tailored by using different concentrations and kinds of BGs, the effect that BGs could have on the regeneration of wounds needs to be further evaluated. Since BGs alone are not practicable as wound dressings, the favourable biological effects that BGs will have on wound regeneration must be exploited in practice by developing composites, which combine BGs with biopolymers.
In this thesis, MC and MH were chosen to be combined with the developed ion-doped BG particles using three well-known polymer processing techniques: freeze drying, 7 Summary and Outlook 141 electrospinning and 3D printing. Prior to discussing the differences between the wound dressings fabricated using the three different methods, the suitability of MC and MH as materials for wound regeneration in general was investigated and it can be discussed based on the results obtained in this project. As mentioned in section 2.5, MH was chosen based on two main reasons: i) due to its great biological effects and ii) due to its potential as natural cross-linker [142], [179]. The successful (mild) crosslinking effect of MH on MC could be proven by FTIR measurements and was further indicated by the sustained release of MH, characterized using UV-VIS measurements. In addition, the amount of MH incorporated was shown to be non-toxic in all performed cell culture tests (in Chapters 4-6) and the results further indicated a positive influence of MH on fibroblasts, keratinocytes and endothelial cells. Moreover, depending on the concentration, MH was found to be effective against the bacteria E. coli and S. aureus. MC was also shown to be a suitable biomaterial for the fabrication of wound dressings. Especially the possibility to dissolve MC in water and to change the gel formation temperature by adding salts makes it an attractive choice for applications in the field of tissue engineering [137]. The results of this part of the study clearly answered the second research questions; “Can suitable wound dressings be fabricated by combining BGs with methyl cellulose (MC) and Manuka honey (MH)?”. Indeed MC as well MH offers the required properties for suitable wound dressings, being thus promising candidates for the fabrication of such devices.
By using freeze drying (Chapter 4), electrospinning (Chapter 5) and 3D printing (Chapter 6), wound dressings were fabricated based on MC, MH and additionally BGs in particulate form. All three methods were found to be suitable to fabricate different forms of wound dressings: foams (freeze drying), fibre mats (electrospinning) and scaffolds (3D printing). As described in Chapter 4, MC-MH foams doped with BG particles exhibit a suitable morphology and provide adequate mechanical and chemical properties. Foams produced by freeze drying can be easily reproduced and fabricated in huge amounts. In contrast to electrospinning and 3D printing, the composition of the freeze-dried foams can be easily changed and fine-tuned (electrospinning and 3D printing require an adjustment of several parameters, e.g. printing speed, voltage). However, by using the electrospinning method (Chapter 5), suitable fibre mats able to mimic the natural ECM can be fabricated. Also electrospun fibre mats usually show suitable mechanical, physical and chemical properties for applications in (soft) tissue engineering. However, the use of benign solvents is necessary and huge quantities of electrospun fibre mats could be just achieved for Blend-based fibre mats. Freeze dried foams and electrospun fibre mats both showed promising results in cell biology tests. In both cases, the addition of MH and BG improved the attachment and proliferation of 142 7 Summary and Outlook hDFs. Moreover, in vitro scratch tests showed the potential of these two kinds of wound dressing to improve wound closure. The main difference found in this thesis between freeze dried foams and electrospun fibre mats is their antibacterial efficiency. Whereas the dual effect of copper-ions and MH showed great results against bacteria, the electrospun fibre mats lacked antibacterial properties. In general, freeze dried foams are especially interesting for the use in deep wounds, whereas electrospun fibre mats are interesting for superficial wounds. However, the combination of both might be also an interesting alternative. The main findings and advantages/disadvantages of freeze drying and electrospinning as processing techniques for wound dressings are summarized in Table 13.
As third fabrication method, 3D printing was investigated in this thesis (Chapter 6). In accordance to freeze drying and electrospinning, a wound dressing fabricated by 3D printing composed of MC-MH-BG was successfully fabricated. However, in order to obtain a stable 3D printed scaffold, the viscosity of the MC-MH ink had to be increased. The addition of salts in form of PBS was shown to be a suitable method to obtain a stable and printable MC-MH ink. By adding additional BG particles, the mechanical and degradation properties of the resulting 3D printed MC-MH scaffolds could be tuned. Moreover, preliminary cell tests proved the cytocompatibility of the printed BG containing MC-MH scaffolds. These results not only demonstrate the suitability of the printed scaffolds as wound dressings, they also indicate the possibility to use MC-MH based inks containing additionally BGs for biofabrication. In the last years, an increasing amount of research efforts is focusing on the possibility of embedding cells into hydrogels and the use of such hydrogels containing cells for 3D printing [349]–[351].
Based on the obtained results (Table 13), the last research question; “Are freeze drying, electrospinning and 3D printing suitable fabrication methods to produce modern wound dressings incorporating BG in particle form?” can be answered: the three methods investigated are suitable to produce wound dressings for different kinds of wounds. In contrast to 3D printing, by freeze drying and electrospinning, wound dressings close to commercially available wound dressings in the form of foams and fibre mats, respectively, were fabricated. These two kinds of dressings were not only mechanically, physically and chemically tested, also their potential in contact with cells and bacteria was investigated. In case of 3D printing, the potential of MC-MH hydrogels containing additionally BG as composite ink was shown. However, only an initial evaluation of the printability and cytocompatibility of MC-MH-BG inks was carried out. More research is necessary to use this ink in combination with cells for biofabrication, e.g. for the development of skin tissue analogues.
7 Summary and Outlook 143
Table 13: Advantages and disadvantages of freeze drying, electrospinning and 3D printing for the fabrication of MC-MH based wound dressing containing additionally BG particles, according to the results of the present project
Method Practical properties Physical properties Biological properties Freeze drying + Reproducible in high + High porosity + Antibacterial by MH and quantities (interconnectivity must be copper
+ No need of toxic further evaluated) + Cell studies using
solvents + Improved mechanical fibroblasts indicate good
+ Cost-effective properties by addition of biocompatibility
BG + Easy to handle and cut + Improved cell migration
in desired shapes + Keep moist environment (Hacat cells and by turning into a gel fibroblasts) (scratch test) - Sterilization needs to be
evaluated + Suitable degradation - Cytotoxicity level of and release rate of ions copper needs to be further and MH evaluated Electrospinning +/- Need of benign + Nano-fibre mats in the - Lack of antibacterial solvents required range could be properties
- Need of PCL to produce produced + Direct cell studies using
high quantities + Hydrophilicity fibroblasts indicate good
+ Reproducible + Suitable mechanical biocompatibility
+ Easy to handle and cut properties and + Improved cell migration in desired shapes degradation/release of Hacat cells (scratch test) properties - Sterilization needs to be evaluated 3D printing + Good printing accuracy + Designed porosity and + Direct cell studies using by adding BGs interconnectivity fibroblasts indicate good
+ Complex shapes + Tuneable mechanical biocompatibility printable properties by crosslinking
+ Reproducible in high degree and addition of BGs
quantities + Suitable degradation
+ No toxic solvents properties needed
- Sterilization needs to be evaluated 144 7 Summary and Outlook
7.2 Future directions
Novel BGs doped with copper and/or zinc and their application in combination with MC and MH as wound dressing were investigated in this thesis. The results of the experimental study are promising and thus new challenges emerged, which need to be investigated in further studies. Based on the results obtained and presented in the different chapters, the following suggestions for further investigations were identified.
Borate and borosilicate glasses doped with copper and/or zinc
In this project, novel borate and borosilicate BGs doped with copper and/or zinc were successfully fabricated and were characterized by several methods. Based on the structural analysis of the fabricated BGs, several observations could be made, which open the possibility to extend the research as follows:
- A series of borosilicate BGs with different silica to boron trioxide ratios should be fabricated to investigate the effect that the addition of boron trioxide has on the resulting BG network by using e.g. Raman and NMR measurements. Besides evaluating the structural effects, also the effect of adding boron trioxide on the degradation and ion release properties should be examined using the series of borosilicate BGs.
- A series of Cu-doped BGs as well as Zn-doped BGs with varying Cu/Zn content should be fabricated to investigate the effect that the addition of copper or zinc has on the network structure of the borate BGs. Based on the structural analysis and dissolution evaluation, the results of this new study should allow to predict the concentrations of biologically active ions released from the BGs during implantation. In addition, in case of zinc, a borate BG should be developed able to release zinc in a relevant amount, which was not achieved in the present investigation.
Besides gaining more knowledge about the design of BGs and the influence that the change of network formers and modifiers could have, the fabricated series of borosilicate BGs and of Cu/Zn containing BGs should be used to study the dose-depending effect of ions released by the BGs. Here a special focus should be given to boron, which is known to be cytotoxic above a certain concentration level. More interestingly, a therapeutic window of copper and zinc should be found, ideally providing a favourable biologically effect on cells, being antibacterial and not toxic. Based on the therapeutic window, the designed BGs can be used for several applications, in addition to wound dressings, for example skin regeneration and vascularization approaches in tissue engineering. 7 Summary and Outlook 145
Fabrication of bioactive glass containing wound dressings
In general, the potential of combining BGs (doped with copper and/or zinc) with MC and MH for the fabrication of wound dressings (using electrospinning, freeze drying and 3D printing) should be further explored. The release profiles of the ionic dissolution products of the BGs and the release of MH from the composite wound dressings are of key importance. Hereby, it should be mentioned that another method of examining the release of MH should be explored (e.g. by measuring the glucose content in the dissolution media). More studies are necessary to assess whether the release profiles are reproducible and the released amounts are within the therapeutic window of the respective ions and MH. In this regard, possible future directions could be:
- Changing the concentration of the BGs within the fabricated wound dressings in order to improve the mechanical properties as well as the amount of ions released by the BGs
- Optimizing the cross-linking process by changing the concentration of MH in terms of mechanical stability, degradation properties as well as release of MH
These two possible future tasks should be done for all three types of fabrication methods used within this project. However, whenever the concentration and therefore the release amount of MH and/or BG are changed, it is important to test if the released concentrations of ions/MH are still not toxic to cells. In addition, the results of this study indicate synergistic effect of MH and ions (as shown in the antibacterial study). These synergistic effects should be further evaluated in order to find out the optimum amount of BGs in combination with MH. Moreover, for each fabrication method future topics could be identified:
i) Freeze drying
As shown in the chapter focusing on 3D printing (chapter 6), the addition of salt in form of PBS has a great influence on the viscosity of the produced hydrogel and thus on the mechanical and degradation properties of the structure made out of this hydrogel. Since the viscosity is also a crucial aspect in producing freeze dried foams, the influence of adding salt should be also investigated when processing freeze dried MC-MH-BG foams. Based on the promising results obtained in chapter 6, it is can be anticipated that the porosity, mechanical stability and especially degradability of the foams can be tuned by adding PBS. Further research should also include the evaluation of the porosity, especially focusing on the interconnectivity of the pores. Although results obtained by using SEM are promising, it is necessary to examine the 3D porosity for instance by µCT. 146 7 Summary and Outlook
ii) Electrospinning
In this project, in order to fabricate high quantities of MC-based electrospun fibre mats, MC was blended with PCL. A future direction should be to improve the use of MC for electrospinning by adding salt to the pure MC electrospinning solution. Based on the resulting increased viscosity of the MC-PBS solution, it might be possible to fabricate higher quantities of electrospun MC-based fibre mats without the need of adding PCL. In addition, the use of salt to increase the viscosity might also be suitable to reduce the needed electrospinning voltage to produce pure MC as well as Blend fibre mats. Although the cell biology tests were promising, the fabricated MC-MH-BG fibre mats in this project did not show antibacterial properties. In order to diminish this drawback, thicker electrospun fibre mats can be produced or more BG/MH can be added to the electrospun fibre mat. Another possibility would be to impregnate the electrospun fibre mat with MH after the electrospinning process.
iii) 3D printing
As already mentioned in the discussion of Chapter 6 (section 6.4.3), the results presented in Chapter 6 indicate the great possibility of using the developed MC-MH-BG inks for biofabrication. This technique offers the advantage to fabricate patient-specific wound dressings in a one-step process. In order to be suitable for biofabrication, the flow and shear rheological properties of the inks must be evaluated and tuned according to the needs of the cells, which are incorporated into the ink prior to printing.
Finally, critical industrial needs, such as in vivo validation, scaling up and regulatory issues must be considered prior to developing a final biomedical product. One important task, which requires further investigations, is the sterilization process of the fabricated freeze dried, electrospun or 3D printed structures. After solving this issue, as well as the other discussed tasks, the wound dressing should be tested in contact with other bacteria as well as in contact with fungi or virus. Moreover, in depth in vitro tests using 3D wound healing models (including infected wound models) [352]–[354] should be done prior to conducting in vivo experiments. In addition, the capability of the MC-MH wound dressings containing BGs to improve wound regeneration should be tested in comparison to commercial available wound dressings, such as the borate BG-based wound dressing MirragenTM [100], [101]. Nevertheless, the results of this thesis indicate the great potential that ion-doped BGs in general and more specifically, in combination with MC and MH, have in wound regeneration applications. References 147
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