Francesca Cecchinato on Magnesium-Modified Titanium Coatings and Magnesium Alloys for Oral and Orthopaedic Applications: in Vitro Investigation

IN VITRO

INVESTIGATION

ON MAGNESIUM-MODIFIED ON FRANCESCA CECCHINATO AND COATINGS TITANIUM FOR ALLOYS MAGNESIUM ORAL ORTHOPAEDIC AND APPLICATIONS:

DOCTORAL DISSERTATION IN ODONTOLOGY

ON MAGNESIUM-MODIFIED TITANIUM COATINGS AND MAGNESIUM ALLOYS FRANCESCA CECCHINATO MALMÖ UNIVERSITY 2015 FOR ORAL AND ORTHOPAEDIC APPLICATIONS: IN VITRO INVESTIGATION

ON MAGNESIUM-MODIFIED TITANIUM COATINGS AND MAGNESIUM ALLOYS FOR ORAL AND ORTHOPAEDIC APPLICATIONS: IN VITRO INVESTIGATION Malmö University Faculty of Odontology Doctoral Dissertations 2015

© Francesca Cecchinato, 2015 Photographs and illustrations: Francesca Cecchinato ISBN 978-91-7104-641-3 (print) ISSN 978-91-7104-642-0 (pdf) Holmbergs, Malmö 2015 FRANCESCA CECCHINATO ON MAGNESIUM-MODIFIED TITANIUM COATINGS AND MAGNESIUM ALLOYS FOR ORAL AND ORTHOPAEDIC APPLICATIONS: IN VITRO INVESTIGATION

Malmö University, 2015 Department of Prosthodontics Faculty of Odontology Malmö, Sweden This publication is also available in electronic format at: www.mah.se/muep Dedicated to my father and mother. This thesis is number 47 in a series of investigations on implants, hard tissue, and the Thesis defended 6.2.1995. External examiner: Docent L. Linder. locomotor apparatus originating from the Department of Biomaterials, University of Gothenburg and the Department of Prosthodontics/Material Sciences, Malmö 13. Patricia Campbell BA, 1995. On Aseptic Loosening in Total Hip Replacement: the University, Sweden. Role of UHMWPE Wear Particles. Thesis defended 7.2.1995. External examiner: Professor D. Howie. 1. Anders R Eriksson DDS, 1984. Heat-induced Bone Tissue Injury. An in vivo investigation of heat tolerance of bone tissue and temperature rise in the drilling of 14. Ann Wennerberg, DDS, 1996. On Surface Roughness and Implant Incorporation. cortical bone. Thesis defended 19.4.1996. External examiner: Professor P-O. Glantz. Thesis defended 21.2.1984. External examiner: Docent K-G. Thorngren.

2. Magnus Jacobsson MD, 1985. On Bone Behaviour after Irradiation. 15. Neil Meredith BDS MSc FDS RCSm, 1997. On the Clinical Measurement of Thesis defended 29.4.1985. External examiner: Docent A. Nathanson. Implant Stability Osseointegration. Thesis defended 3.6.1997. External examiner: Professor J. Brunski.

3. Fredric Buch MD, 1985. On Electrical Stimulation of Bone Tissue. Thesis defended 28.5.1985. External examiner: Docent T. Ejsing-Jörgensen. 16. Lars Rasmusson DDS, 1998. On Implant Integration in Membrane-Induced and Grafter Bone. Thesis defended 4.12.1998. External examiner: Professor R. Haanaes. 4. Peter Kälebo MD, 1987. On Experimental Bone Regeneration in Titanium Implants. A quantitative microradiographic and histologic investigation using the Bone Harvest Chamber. 17. Thay Q Lee MSc, 1999. On the Biomechanics of the Patellfemoral Joint and Patellar Thesis defended 1.10.1987. External examiner: Docent N. Egund. Resurfacing in Total Knee Arthroplasty. Thesis defended 19.4.1999. External examiner: Docent G. Nemeth.

5. Lars Carlsson MD, 1989. On the Development of a new Concept for Orthopaedic 18. Anna Karin Lundgren DDS, 1999. On Factors Influencing Guided Regeneration and Implant Fixation. Augmentation of Intramembraneous Bone. Thesis defended 2.12.1989. External examiner: Docent L-Å Broström. Thesis defended 7.5.1999. External examiner: Professor B. Klinge.

6. Tord Röstlund MD, 1990. On the Development of a New Arthroplasty. Thesis defended 19.1.1990. External examiner: Docent Å. Carlsson. 19. Carl-Johan Ivanoff DDS, 1999. On Surgical and Implant Related Factors Influencing Integration and Function of Titanium Implants. Experimental and Clinical Aspects. 7. Carina Johansson Res Tech, 1991. On Tissue Reaction to Metal Implants. Thesis defended 12.5.1999. External examiner: Professor B. Rosenquist. Thesis defended 12.4.1991. External examiner: Professor K. Nilner. 20. Bertil Friberg DDS MDS, 1999. On Bone Quality and Implant Stability 8. Lars Sennerby DDS, 1991. On the Bone Tissue Response to Titanium Implants. Measurements. Thesis defended 24.9.1991. External examiner: Dr J.E. Davies. Thesis defended 12.11.1999. External examiner: Docent P. Åstrand.

9. Per Morberg MD, 1991. On Bone Tissue Reactions to Acrylic Cement. 21. Åse Allansdotter Johansson MD, 1999. On Implant Integration in Irradiated Bone. Thesis defended 19.12.1991. External examiner: Docent K. Obrant. An Experimental Study of the Effects of Hyperbaric Oxygeneration and Delayed Implant Placement. 10. Ulla Myhr PT, 1994. On factors of Importance for Sitting in Children with Cerebral Thesis defended 8.12.1999. External examiner: Docent K. Arvidsson-Fyrberg. Palsy. Thesis defended 15.4.1994. External examiner: Docent K. Harms-Ringdahl. 22. Börje Svensson FFS, 2000. On Costochondral Grafts Replacing Mandibular Condyles in Juvenile Chronic Arthritis. A Clinical, Histologic and Experimental Study. 11. Magnus Gottlander MD, 1994. On Hard Tissue Reactions to Hydroxyapatite- Thesis defended 22.5.2000. External examiner: Professor Ch. Lindqvist. Coated Titanium Implants. Thesis defended 25.11.1994. External examiner: Docent P. Aspenberg. 23. Warren Macdonald BEng, MPhil, 2000. On Component Integration on Total Hip Arthroplasties: Pre-Clinical Evaluations. 12. Edward Ebramzadeh MScEng, 1995. On Factors Affecting Long-Term Outcome of Thesis defended 1.9.2000. External examiner: Dr A.J.C. Lee. Total Hip Replacements.

6 7 This thesis is number 47 in a series of investigations on implants, hard tissue, and the Thesis defended 6.2.1995. External examiner: Docent L. Linder. locomotor apparatus originating from the Department of Biomaterials, University of Gothenburg and the Department of Prosthodontics/Material Sciences, Malmö 13. Patricia Campbell BA, 1995. On Aseptic Loosening in Total Hip Replacement: the University, Sweden. Role of UHMWPE Wear Particles. Thesis defended 7.2.1995. External examiner: Professor D. Howie. 1. Anders R Eriksson DDS, 1984. Heat-induced Bone Tissue Injury. An in vivo investigation of heat tolerance of bone tissue and temperature rise in the drilling of 14. Ann Wennerberg, DDS, 1996. On Surface Roughness and Implant Incorporation. cortical bone. Thesis defended 19.4.1996. External examiner: Professor P-O. Glantz. Thesis defended 21.2.1984. External examiner: Docent K-G. Thorngren.

6 7 24. Magne Røkkum MD, 2001. On Late Complications with HA Coated Hip 35. Andreas Thor DDS, 2006. On platelet-rich plasma in reconstructive dental implant Arthroplasties. surgery. Thesis defended 12.10.2001. External examiner: Professor P. Benum. Thesis defended 8.12.2006. External examiner: Professor E.M. Pinholt.

25. Carin Hallgren Höstner DDS, 2001. On the Bone Response to Different Implant 36. Luiz Meirelles DDS MSc, 2007. On Nano Size Structures for Enhanced Early Bone Textures. A 3D analysis of roughness, wavelength and surface pattern of experimental Formation. implants. Thesis defended 13.6.2007. External examiner: Professor Lyndon F. Cooper. Thesis defended 19.11.2001. External examiner: Professor S. Lundgren. 37. Pär-Olov Östman DDS, 2007. On various protocols for direct loading of implant- 26. Young-Taeg Sul DDS, 2002. On the Bone Response to Oxidised Titanium Implants: supported fixed prostheses. The role of microporous structure and chemical composition of the surface oxide in Thesis defended 21.12.2007. External examiner: Professor B. Klinge. enhanced osseointegration. Thesis defended 7.6.2002. External examiner: Professor J.E. Ellingsen. 38. Kerstin Fischer DDS, 2008. On immediate/early loading of implant supported prostheses in the maxilla. 27. Victoria Franke Stenport DDS, 2002. On Growth Factors and Titanium Implant Thesis defended 8.2.2008. External examiner: Professor K. Arvidsson Fyrberg. Integration in Bone. Thesis defended 11.6.2002. External examiner: Associate Professor E. Solheim. 39. Alf Eliasson 2008. On the role of number of fixtures, surgical technique and timing of loading. 28. Mikael Sundfeldt MD, 2002. On the Aetiology of Aseptic Loosening in Joint Thesis defended 23.5.2008. External examiner: Professor K. Arvidsson Fyrberg. Arthroplasties and Routes to Improved cemented Fixation. Thesis defended 14.6.2002. External examiner: Professor N. Dahlén. 40. Victoria Fröjd DDS, 2010. On Ca2+ incorporation and nanoporosity of titanium surfaces and the effect on implant performance. 29. Christer Slotte CCS, 2003. On Surgical Techniques to Increase Bone Density and Thesis defended 26.11.2010. External examiner: Professor J.E. Ellingsen. Volume. Studies in Rat and Rabbit. Thesis defended 13.6.2003. External examiner: Professor C.H.F. Hämmerle. 41. Lory Melin Svanborg DDS, 2011. On the importance of nanometer structures for implant incorporation in bone tissue. Thesis defended 01.06.2011. External examiner: Associate professor C. Dahlin. 30. Anna Arvidsson MSc, 2003. On Surface Mediated Interactions Related to Chemomechanical Caries Removal. Effects on surrounding tissues and materials. Thesis defended 28.11.2003. External examiner: Professor P. Tengvall. 42. Byung-Soo Kang MSc, 2011. On the bone tissue response to surface chemistry modifications of titanium implants. 31. Pia Bolind DDS, 2004. On 606 retrieved oral and cranio-facial implants. An analysis Thesis defended 30.09.2011. External examiner: Professor J. Pan. of consequently received human specimens. Thesis defended 17.12.2004. External examiner: Professor A. Piattelli. 43. Kostas Bougas DDS, 2012. On the influence of biochemical coating on implant bone incorporation. 32. Patricia Miranda Burgos DDS, 2006. On the influence of micro- and macroscopic Thesis defended 12.12.2012. External examiner: Professor T. Berglundh. surface modifications on bone integration of titanium implants. Thesis defended 1.9.2006. External examiner: Professor A. Piattelli. 44. Arne Mordenfeld DDS, 2013. On tissue reaction to and adsorption of bone substitutes. 33. Jonas P. Becktor DDS, 2006. On factors influencing the outcome of various Thesis defended 29.5.2013. External examiner: Professor C. Dahlin. techniques using endosseous implants for reconstruction of the atrophic edentulous and partially dentate maxilla. 45. Ramesh Chowdhary DDS, 2014. On efficacy of implant thread design for bone Thesis defended 17.11.2006. External examiner: Professor K.F. Moos. stimulation. Thesis defended 21.05.2014. External examiner: Professor Flemming Isidor. 34. Anna Göransson DDS, 2006. On Possibly Bioactive CP Titanium Surfaces. Thesis defended 8.12.2006. External examiner: Professor B. Melsen. 46. Anders Halldin MSc, 2015. On a biomechanical approach to analysis of stability and load bearing capacity of oral implants. Thesis defended 28.05.2015. External examiner: Professor J. Brunski.

8 9 24. Magne Røkkum MD, 2001. On Late Complications with HA Coated Hip 35. Andreas Thor DDS, 2006. On platelet-rich plasma in reconstructive dental implant Arthroplasties. surgery. Thesis defended 12.10.2001. External examiner: Professor P. Benum. Thesis defended 8.12.2006. External examiner: Professor E.M. Pinholt.

8 9

47. Francesca Cecchinato MSc, 2015. On magnesium-modified titanium coatings and magnesium alloys for oral and orthopaedic applications: in vitro investigation. Thesis to be defended 20.11.2015. External examiner: Professor C. Stanford.

See www.mah.se/muep

See www.mah.se/muep TABLE OF CONTENTS

LIST OF PAPERS...... 13 ABSTRACT...... 15 ACRONYMS AND SYMBOLS...... 19 INTRODUCTION...... 22 Bone replacement and repair...... 22 Peri-implant bone healing...... 27 Magnesium in bone metabolism...... 30 Magnesium-based biomaterials...... 31 AIMS...... 37 MATERIALS AND METHODS...... 38 Specimen preparation...... 38 Surface characterization...... 41 Magnesium release...... 44 Material degradation parameters...... 45 Cells...... 46 Cell isolation and expansion...... 47 Cell seeding and culture...... 48 Cell morphology...... 49 In vitro cytotoxicity...... 50 Cell adhesion...... 51 Alkaline phosphatase...... 53 Cell mineralization...... 53 RNA extraction ...... 54 Gene expression techniques...... 56 Statistics...... 59

10 RESULTS...... 60 Magnesium as a bioactive substance for mesoporous titania implant coatings...... 60 Magnesium alloys as bioresorbable metals for bone applications...... 72 DISCUSSION...... 79 Magnesium-modified mesoporous titania films to enhance peri-implant osteogenesis - studies I, II and III...... 80 Magnesium alloys as tailored biodegradable implant materials for bone regeneration – study IV...... 87 CONCLUSIONS AND FUTURE PERSPECTIVES...... 91 Mesoporous titania film as a carrier for magnesium at the peri-implant site...... 91 Magnesium alloys as biodegradable metals for bone tissue regeneration...... 92 ACKNOWLEDGEMENTS...... 95 REFERENCES...... 98 PAPERS I-IV...... 113 LIST OF PAPERS

This thesis is based on the following papers, which are referred to in the text by their Roman numerals.

I. Cecchinato F, Xue Y, Karlsson J, He W, Wennerberg A, Mustafa K, Andersson M, Jimbo R. In vitro evaluation of human foetal

osteoblast response to magnesium loaded mesoporous TiO2 coating. J Biomed Mater Res A. 2013; 102(11) 3862-71.

II. Cecchinato F, Karlsson J, Ferroni L, Gardin C, Galli S, Wennerberg A, Zavan B, Andersson M, Jimbo R. Osteogenic potential of human adipose-derived stromal cells on 3-

dimensional mesoporous TiO2 coating with magnesium impregnation. Material Science and Engineering 2015; 225-234.

III. Cecchinato F, Atefyekta, Wennerberg A, Andersson M, Jimbo R, Davies J. Modulation of the nanometer pore size improves magnesium adsorption into mesoporous titania coatings and promotes bone morphogenic protein 4 expression in adhering osteoblasts. Manuscript.

IV. Cecchinato F, Agha NA, Martinez-Sanchez AH, Luthringer BJC,

Feyerabend F, Jimbo R, Willumeit-Römer R, Wennerberg A. Influence of magnesium alloy degradation on undifferentiated human cells. Submitted.

Paper I reprinted by permission from John Wiley and Sons, License Number 3670081040368. Paper II reprinted by permission from Elsevier, License Number 3670090088607.

13 THESIS AT A GLANCE

STUDY HYPOTHESIS ILLUSTRATION KEY FINDINGS I Magnesium can be loaded Magnesium adsorbs to mesoporous titania films In vitro evaluation of in mesoporous films, and with a 6-nm average pore size and is released human foetal osteoblast its release enhancesosteob- within 24 hours of contact with bone cells,

response to magnesium- last activity. demonstrating a positive effect on initial osteob-

loaded mesoporous TiO2 last viability. coatings. II Magnesium exhibitsosteo- Magnesium released from 6-nm mesoporous films Osteogenic potential of conductive potential by promotes (i) stromal cell differentiation along the human adipose-derived guiding undifferentiated osteogenic lineage; and (ii) osteopontin expression stromal cells on 3-dimen- cell differentiation toward in particular. sional mesoporous TiO2 the osteoblast phenotype. coatings with magnesium impregnation. III A 1nanometre increase The nano-pore dimensions of mesoporous films Modulation of nanometre in pore size increases modulate magnesium adsorption. An increase of pore size improves magne- magnesium content in the 1 nanometre in pore size increases Mg content, sium adsorption in meso- mesoporous structure, which once released, significantly promotes bone porous titania coatings and thus improving magnesium morphogenic protein 4 expression at later stages promotes bone morphoge- release and its osteogenic of osteoblast proliferation. nic protein 4 expression in potential in vitro. adhering osteoblasts. IV The degradation behaviour Degradation rate and degradation parameters are Influence of magnesium and the surface chemistry higher for Mg4Y3RE compared to Mg2Ag and alloy degradation on undif- and topography of three Mg10Gd. However, cellular adhesion structures ferentiated human cells. magnesium alloys dif- are better developed on Mg10Gd, which pos - ferently influence early sesses homogeneous degradation and, therefore, adhesion and spreading of represents a suitable alternative to biodegradable undifferentiated cells. metal in bone. TiO2=Titanium dioxide; Mg2Ag=Mg-silver; Mg10Gd=Mg-gadolinium; Mg4Y3RE=Mg-yttrium-rare-earth In Study II, Tokyo, Japan) the cultured cells point in a Balzers CPD 030 ( hydrated the samples in ethanol, and dried them at the critical (BalTec,dryer dried them at thecritical point in a BalzersCPD 010 critical point into specimens, dehydrated the cells in graded ethanol, and then In Study I, Study In imaging. to prior tive conductive pow Sam this allows cells to maintain their m their maintain to cells allows this tem ambient 49 ethanol. Cellsthen underwent critical 100% with cells the inside water to the replace necessary was dure before dehydration in a graded series of ethanol. Dehy strate - sub a to cellsadhering microscopy visualizing for electron of form Scanning electron microscopy (SEM) is the most commonly used Scanning electronmicroscopy morphologyCell F

fluid, such as CO replaces the liquid inside biological structures with a suitable inert seeding i g

u r ples e . After culturing, cells culturing, . After

4. on

were then mounted Schematic pre hFOB ADSC ADSC -incubated perature. perature.

. Pfä der to make the cell monolayer electrically conduc-

were fixed were were taken with taken were illustration Last, Switzerland). ffikon ZH,

were fixed after after fixed were 2 , whose critical temperature is just above above just is temperature critical , whose Mg Mg Thus, nosurface tension effect occurs, and all oys

were immobilized were after and covered them with a thin layer of of for for the Leica MicrosystemLeica Wetzlar, AB,

- the for 1 and 24 hours Studies I, II, and and IV II, I, Studies

experimental SEM (JSM-

5 and 24 hours orphology without distortion. in

- vitro point drying. This metho

investigations

procedure in a fixation proce 7400F; JEOL Ltd., Ltd., JEOL 7400F; micrographs of

of culture of cell culture culture cell of

(Studies for for dra HUCPV , de tion IV). d

- -

ABSTRACT

In dentistry and orthopaedic surgery, research to find and develop improved biomaterials is progressing rapidly. Of specific interest is to accelerate bone formation around the implant surface, which could improve the reliability of the implant even in patients with compromised situations. Although the surface modification of the implant has been proven to certain extent to promote osseointegration, the lack of bone in the patient remains a major issue and bone augmentation is commonly conducted prior to implant insertion. Synthetic and naturally derived resorbable materials are commonly used. However, problems such as the lack of optimal mechanical properties or the undesirable material resorption kinetics still exist and there still remain possibility for improvement. Clinical approaches for orthopaedic trauma require the use of non-resorbable screws, plates and pins made of metallic materials such as titanium, cobalt-chrome and stainless steel alloys. The major drawback of these materials is the need of implant removal at re-entry. Therefore, the research of bioresorbable materials that could withstand the mechanical stresses is an ongoing topic.

Based on this clinical reality, the aim of this thesis was to investigate the suitability of magnesium (Mg) as a biomaterial for regenerative bone applications. Namely, Mg as a doping material for engineered mesoporous titanium implant surfaces (Studies I, II and III), and as a bioresorbable metal alloy for bone regeneration in bone trauma and bone defects conditions (Study IV).

15 15 Study I, II, III Mesoporous titania films produced with evaporation-induced self- assembly (EISA) technique and applied as implant surface coatings are under investigation as a release system for the controlled administration of several substances, such as osteoporotic drugs, to enhance early bone anchorage to the implant. Modulating the pore size of such films though the selection of EISA parameters permits to control the adsorption of such substances into the mesoporous matrix and their subsequent release into the peri-implant region. Studies I, II and III analysed the effect of Mg incorporation into mesoporous titania coatings towards two cellular models during early and later stages of cell activity.

Study I characterized the morphology, chemistry, and topography of mesoporous titania coatings and the effects of Mg-loading on surface micro- and nano-structures. Mg release was determined and its effect was evaluated on human foetal osteoblast populations. It was shown that mesoporous films possessed a smooth surface with pores that faced outward. Mg adsorption did not substantially alter the mesoporous surface roughness both at micro- and nano- levels. Mg was released within 24 hours of incubation in cell culture conditions, thus its bioactive effect only occurred during initial osteoblasts activity.

Study II evaluated the ability of Mg-loaded mesoporous coatings to modulate multipotent adipose-derived stromal cell differentiation toward the osteoblast phenotype. The results demonstrated that Mg release had a strong impact on this cellular model, promoting osteoblast marker expression in standard cell culture conditions. Interestingly, Mg significantly promoted the expression of osteopontin, a protein that is essential for early biomaterial-cell osteogenic interaction.

In study III, the reagents and EISA parameters in the mesoporous deposition were varied to generate three mesoporous titania coatings with 2-, 6- and 7-nm average pore size, to increase Mg content in the interconnected porosity of the films. The effect of various Mg contents released from the three mesoporous structures

16 16 was tested on human foetal osteoblasts populations with pre-designed osteogenic PCR arrays and real-time polymerase chain reaction. It was shown that Mg release affected osteogenesis and was controlled by tuning the pore dimensions of the mesoporous films. Increasing pore size by 1 nm (from 6 nm to 7 nm) significantly enhanced the bioactivity of the film without altering the surface roughness.

Study IV In orthopaedics Mg alloys has received increasing attention as bioresorbable metals for bone regeneration. However, localized material degradation is too fast and provokes the premature loss of mechanical properties, preventing correct cellular development and bone healing in vivo . For this reason, various alloying elements are combined with high-purity Mg to modulate and optimize degradation behaviour.

Study IV of this thesis investigated the degradation parameters of Mg2Ag, Mg10Gd, and Mg4Y3RE alloys and how the alloys differently affect human umbilical cord perivascular cell adhesion and spreading. Mg4Y3RE showed the highest degradation rate and, thereby, the highest trend in increases in pH and osmolality of the surrounding fluid. However, both Mg4Y3RE and Mg10Gd allowed cells to better adhere and spread across their degraded surfaces; in comparison, surface degradation of Mg2Ag was more aggressive with weak or no visible cellular structures on it.

Conclusions In summary, the results of the present thesis explored the potential of Mg for its application in bone tissue regeneration. Titanium implant surfaces coated with mesoporous TiO2 thin films and further loaded with Mg enhanced bone cell activity and osteoprogenitor development into mature osteoblasts. Thus, mesoporous deposition followed by Mg loading may be a suitable alternative to existing implant surface treatments.

17 17 Bioresorbable materials must degrade slowly and uniformly in order to simulate the tissue healing process. Mg10Gd possesses reduced content of alloying element and a suitable homogenous degradation pattern in vitro that allows proper adhesion of undifferentiated cells. Mg10Gd thus represents a biodegradable Mg-based material with promising mechanical and biological properties for use in dental and orthopaedic fields.

18 18 ACRONYMS AND SYMBOLS

ACRONYMS 3-D Three-dimensional ADSC Adipose-derived stromal cells AFM Atomic force microscopy ASTM American Society for Testing and Materials ATP Adenosine triphosphate cDNA Complementary deoxyribonucleic acid

CO2 Carbon dioxide Co-Cr Cobalt-chromium CPD Critical point drying CP Ti Commercially pure titanium DAPI 2-(4-Amidinophenyl)-1H-indole-6-carboxamidine DMEM/MEM Dulbecco’s modified eagle medium DNA Deoxyribonucleic acid ECM Extracellular matrix EDTA Ethylenediaminetetraacetic acid EDX Energy-dispersive X-ray spectroscopy EISA Evaporation-induced self-assembly FBS Foetal bovine serum FITC Fluorescein isothiocyanate HA Hydroxyapatite hFOB Human foetal osteoblasts HUCPV Human umbilical cord perivascular MAO Micro-arc oxidation Mg2Ag Magnesium-silver Mg4Y3RE Magnesium-yttrium-rare-earth Mg10Gd Magnesium-gadolinium MP Mesoporous MPMg Mg-loaded mesoporous

19 19 MSCs Mesenchymal stem cells MTT 3-(4,5-Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide NaCl Sodium chloride NP Non-porous PIIID Plasma immersion ion implantation and deposition pNPP p-nitrophenyl phosphate PTFE Polytetrafluoroethylene PTH Parathyroid hormone QCM-D Quartz crystal microbalance with dissipation monitoring RNA Ribonucleic acid RT-PCR Real-time polymerase chain reaction SEM Scanning electron microscopy TEOT Titanium (IV) ethoxide

TiO2 Titanium dioxide, or titania TRITC- Tetramethylrhodamine isothiocyanate-conjugated phalloidin phalloidin International Standardization Organization, Biological ISO 10993 Evaluation of Medical Devices WJ Wharton’s jelly XPS X-ray photoelectron spectroscopy

ELEMENT SYMBOLS Ag Silver Al Aluminium Ar Argon C Carbon Ca Calcium Ce Cerium F Fluor Gd Gadolinium La Lanthanum Li Lithium Mg Magnesium Mn Manganese N Nitrogen O Oxygen P Phosphorus S Sulphur Ti Titanium

20 20 V Vanadium Y Yttrium Zn Zinc Zr Zirconium

GENE SYMBOLS ALP Alkaline phosphatase B2M Beta-2-microglobulin BMP Bone morphogenic protein BMP2 Bone morphogenic protein 2 BMP4 Bone morphogenic protein 4 BSP Bone sialoprotein COL1 Collagen, type 1 GAPDH Glyceraldehyde 3-phosphate dehydrogenase MMP Matrix metalloproteinase MMP8 Matrix metalloproteinase 8 OCN (BGLAP) Osteocalcin (bone gamma-carboxyglutamate [gla] protein) ONC Osteonectin OPN Osteopontin OSX Osterix RUNX2 Runt-related transcription factor 2 TGFB1 Transforming growth factor, beta 1

21 21 INTRODUCTION

Bone replacement and repair Overview Worldwide, the annual number of prostheses utilized in patients has risen dramatically in recent decades. Increasing life expec- tancies, which have generated more patients in need of treatment, is one factor in this trend [1-3]. About 90% of the population over age 40 years suffers from degenerative bone and joint diseases, such as osteoporosis and osteoarthritis [4]. In osteoporosis, reduced bone mineral density weakens bone structure, potentially causing severe fractures. Approximately 30% of postmenopausal women suffer from osteoporosis in the United States and Europe [5, 6]. Thus, all populations around the world are expected to experience increasing incidences of degenerative bone diseases due to aging.

According to the 2014 Annual Report of the Activities of the Bone and Joint Decade, the huge impact of musculoskeletal conditions on global health represents the second greatest cause of disability, increasing 45% between 1990 and 2010 [7]. Development of new and modified biomaterials is thus needed to improve quality of life, particularly in bone replacement and repair. Use of non-resorbable pins, screws, plates, and rods of metallic materials is common in fracture management, whilst joint replacement includes permanent devices for the hip, knee, shoulders, ankle, and elbow. The non- resorbable characteristic of these materials can be a drawback if re- entry into the surgical site, which is invasive and costly, is needed [8, 9].

22 22 The pursuit of alternative bioresorbable materials that function as well as traditional non-resorbable materials has been of some interest in recent decades. For instance, polymers and ceramics have replaced some types of metal implants. These newer materials possess excellent biofunctionality, but they do not possess the strength and mechanical properties that load-bearing implants require [10].

Metallic devices have been widely used to rehabilitate edentulism in the past 50 years [11]. One of the factors proposed by Albrektsson and co-workers that strongly influences the integration of an implant in bone is the implant surface quality [12]. Wennerberg et al. demonstrated that surfaces with moderate microroughness yielded the highest implant success rates in terms of direct bone-to-implant contact [13, 14]. Excessively rough surfaces, such as thermal spray implant surfaces, have been largely discussed, and their use is hesitant due to the risk of negative biological reactions around the implant [15-18].

Cell recruitment and initial interaction seem to occur more rapidly on nano-structured surfaces, which appear to enhance surface bioactivity [19, 20]. Nano-scale features, such as surface nano-pits or isotropic nano-porosity, reportedly influence initial attachment of osteoprogenitor cells positively [21-23]. Modification of an implant surface at the nano level generally includes modulation of the topography and also the chemistry, providing a synergistic effect that promotes new bone formation [24, 25]. Thus, improving the bioactive properties of the implant surface is a clever strategy for overcoming negative clinical outcomes and improving bone-implant surface adherence in the early stages of bone formation, which is essential for early loading of the implant.

Common metals for orthopaedic applications The fundamental requirements of orthopaedic materials are optimal wear and corrosion resistance and a combination of mechanical and physical properties that allow them to withstand constant loading during function. Currently, the best materials for bone replacement and fixation are considered to be metals: stainless steel (316L stainless steel, ASTM F-593), cobalt-

23 23 chromium alloys (ASTM F-75), commercially pure (CP) titanium (Ti) (ASTM F-76), and Ti alloys (ASTM F-136). Stainless steel is still used in applications ranging from cardiovascular to otorhinology. Cobalt-chromium (Co-Cr) alloys have demonstrated excellent wear resistance and, thereby, better functionality in total joint and femoral head replacement. These applications are subjected to continuous load during function, thus accurate transfer of the loading pressure to the host bone is necessary.

Another concern is the release of metallic non-physiological ions and nano-particles from conventional metals, which may occur at the peri-implant site and lead to inflammatory progression, thus reducing implant stability [26]. The high difference in elastic modulus of these materials compared to bone results in stress shielding events caused by the transfer of normal stresses through the implant to the bone [27]. CP Ti and Ti alloys are the materials of choice in failed joint replacement or fractured bone fixation, due to their light weight, lower density, and better ductility. Moreover, Ti and its alloys have widely demonstrated to be the most compatible materials for tissue healing when inserted into the human body [28]. The bioinert properties of Titanium make the metal ideally suitable for use in dentistry for tooth replacement, orthodontics, and maxillofacial surgery applications.

Materials for oral and maxillofacial applications Surgical repair of maxillofacial bones requires that the plates and screws are able to withstand varying stresses, especially in the mandibular and maxillary regions. In 1976, Champy introduced mini-plates and mini-screws in facial surgery [29]. These components are produced in various sizes and lengths for use in cortical bone of varying thicknesses in the orofacial regions.

Another challenge is the treatment of large bony defects. Localized lack of bone in the jaw is generally caused by infection, trauma, tumours, and skeletal abnormalities [30]. Nowadays, the common strategy for regenerating bone in the oral and maxillofacial region, often involves barrier membranes. The advantage of barrier membranes is their ability to generate space that allows bone

24 24 forming cells to repopulate the damaged site and thus enhance the healing. Currently, a large range of membranes is available on the market. They are made of various materials, such as non- resorbable synthetic polymers or metals, or naturally-derived materials like collagen. Initially, non-resorbable membranes made of polytetrafluoroethylene (PTFE) and reinforced with a Ti frame were adapted to the defect margins to provide a rigid space- keeping effect [32]. The use of Ti meshes was subsequently introduced. These non-resorbable applications provided more stable space maintenance and, thereby, generated more adequate bone volume in the defect than PTFE membranes [33-35]. Regenerative treatment with non-resorbable membranes requires fixation of the device with Ti mini-pins. The entire system must be removed once its role is completed. Therefore, other materials have been used to produce membranes that degrade upon the healing of the injured site, eliminating the need of a second surgery. For instance, pure porcine collagen membranes completely and safely resorb, promoting the natural healing of the bone [36-38]. The problematic relies on its low rigidity and, thereby, poor space- keeping effect. For this reason collagen membranes are normally used together with allogeneic or xenogeneic bone [39].

Dental implants: current trends of implant surface modifications Ti implants were introduced in the 1960s when Brånemark and his colleagues accidentally discovered that Ti successfully integrated in bone, and they are currently used in tooth restoration with a clinical success rate greater than 95%. However, stimulation of initial bone apposition at the peri-implant interface during healing to achieve firm implant stability remains a prominent goal. One of the factors proposed by Albrektsson and co-workers that strongly influenced the integration of implant in bone is implant surface quality [12]. The topographical, chemical, physical, and mechanical characteristics of implant surfaces have been extensively modified to develop implants with predictable tissue- integrative properties. Wennerberg and Albrektsson have proposed guidelines for developing and analysing new Ti implant surfaces with optimal topography [40]. Experimental evidence

25 25 demonstrated that optimal bone response occurred with a moderately roughened implant surface (Sa about 1.5 µm) [13, 14].

Many studies, utilizing different techniques, have attempted to modify the chemical composition of implant surfaces. The principle is based on the assumption that a specific surface chemistry would induce faster bonding of the surrounding cells and proteins to the implant, converting the inert nature of Ti into a bioactive one.

Techniques for modifying the surface of dental implants during manufacture include etching with various acid solutions to obtain a surface roughness of specific dimensions. Ellingsen’s group first introduced this technique using hydrofluoric acid (HF) to etch Ti implants, which they then placed in rabbit bone; after 4 and 8 week of healing, implant retention was increased four-fold compared with non-etched Ti surfaces [41]. Acid treatment produces defined surface textures, with controlled micro- and spontaneously formed nano-scale structures. In some cases, the surface chemistry can be modified by choice of acid solution [41- 44]. Such treatment generates surfaces that incorporate specific electrolytes and hydroxyl groups, thereby changing the surface chemistry. This occurs also in case of electrochemical oxidation treatments, which generates porous structures on the implant surface [43].

Ti implant surfaces can also be modified by grit-blasting. This technique generates defined micro- and nano-structures when metallic or hard ceramic particles are projected at high speeds onto the implant surface. Even here, surface chemistry may be modified due to the incorporation of residual particles [45]. Another strategy for modifying implant surfaces is the deposition of sol-gel coatings, which have been shown to yield good results in vitro and in vivo. Sol-gel coating is the deposition of a gel onto the implant surface that generates various topographical and chemical properties depending on choice of sol-gel solution, period of immersion, and heat treatment parameters [46, 47].

26 26 Another interesting surface coating is mesoporous (MP) titania coating, which is generated using the evaporation-induced self- assembly (EISA) technique. This advanced surface modification combines the benefits of an engineered three-dimensional (3-D) matrix characterized by homogeneous porosity and a nano-structured surface topography [48]. The presence of interconnected nano-porosity that is evenly distributed along the thickness of the coating allows this surface modification to act as a drug delivery system [49, 50]. Although this implant surface is not yet com- mercialized, recent research has demonstrated that the MP matrix can be loaded with drugs, such as Raloxifene and Alendronate, which can be released in a controlled and sustained manner immediately after implant placement into animal tissues. Clearly, this coating may be a suitable alternative for improving the bioactivity of implant surfaces; loading it with bioactive substances or molecules should enhance integration of the implant with peri- implant bone.

Peri-implant bone healing When inserted into host tissue, a biomedical device first makes contact with blood, which excites a series of biological events at the peri-implant site, such as protein adsorption, coagulum formation, inflammation, and new tissue formation. In peri- implant bone healing, osteogenesis occurs as contact osteogenesis and distance osteogenesis. Osborn and Newesly, who defined these two phenomena, described contact osteogenesis as the process by which bone cells build newly formed bone on the implant surface, and distance osteogenesis as the process by which new bone is formed on the surface of the host bone in apposition to the implant. In contact osteogenesis, the adsorption of proteins such as fibrinogen, fibronectin, and vitronectin to the implant surface facilitates cell adhesion to the device [51-54]. Mesenchymal stem cells (MSCs) migrate to the damage site after platelet and specific protein adhesion [55]. MSCs are unspecialized cells (no specific function) that are capable of renewing themselves through cell division. They can differentiate into tissue-specific cells, such as adipocytes, chondrocytes, osteoblasts, neurons, and muscle cells under certain physiological or experimental conditions [56]. MSCs

27 27 are classified according to their origin as embryonic, foetal, or adult: embryonic stem cells derive from the early embryo, such as human umbilical cord stem cells; foetal stem cells derive from the developing foetus; and adult stem cells derive from adult mature organs, such as bone marrow, adipose tissue, the heart, and dental pulp [57].

One of the prerequisites for successful osseointegration around an implant is the correct adhesion and proliferation of MSCs onto the implant surface (Figure 1), which concludes with the complete maturation of an osteoblast population, producing a mineralized matrix and thus the desired conditions for osteoid formation [55, 58]. Biodegradable implants that gradually degrade from the surface must be manufactured from a material that does not shed overly large fragments and products from the outer layer but instead possesses a surface that initially degrades slowly and homogeneously in order to achieve the desired cell adhesion [59]. Thus, choice of material – and its topographical and chemical, but also physical and mechanical properties – has been proven to be crucial for guiding cell colonization and differentiation [19, 20, 60-62].

Figure 1. Schematic illustration of MSC morphology during adhesion, proliferation, and differentiation on a material surface [63].

Cell adhesion, proliferation, and differentiation MSCs adhere to a surface through focal adhesions and cytoskeleton bonding [64]. Once the cells adhere to the implant surface, a specialized group of factors initiates commitment of the MSCs to a specific cellular phenotype (Figure 2). Runt-related transcription factor 2 (RUNX2) plays a central role in the

28 28 differentiation of MSCs into osteoblasts, in part by inhibiting differentiation into adipocytes [65, 66]. After MSC commitment to the osteoblast lineage, MSCs evolve into preosteoblasts prior to developing into immature osteoblasts. Early bone markers that are weakly expressed by preosteoblasts are abundantly produced by immature osteoblasts, thus functioning to indicate the progression of osteogenic differentiation [66]. Immature osteoblasts are spindle-shaped cells that produce bone-related proteins such as osteocalcin (OCN), osteopontin (OPN), and bone sialoprotein (BSP); these non-collagenous proteins are further produced throughout osteoblast development. Mature osteoblasts possess a cuboidal shape and are capable of forming a mineralized bone matrix, due to the expression of collagen type I (COL I); its synthesis is one of the primary activities of differentiated bone cells [67]. Non-collagenous proteins and COL I are major components of extracellular matrix deposition [67]. This matrix stores factors that are involved in bone remodelling, and thus in the activation of bone resorption. Transforming growth factor beta 1 (TGFBI) is a growth factor produced by osteoblasts, but it is also involved in osteoclast-activated bone resorption. Metalloproteinases (MMP) are another class of proteins that plays a crucial role in later stages of bone remodelling. These proteins are collagenases and gelatin- ases that degrade collagen structures in mature bone [55, 68, 69].

29 29

Figure 2. Schematic illustration of mesenchymal stem cell (MSC) morphology and gene expression during differentiation into mature osteoblasts. Runt- related transcription factor 2 (RUNX2) initiates MSC osteogenic differentiation and inhibits adipogenesis. After commitment, MSCs develop into preosteoblasts that express osteogenic factors such as distal-less homeobox 5 (Dlx5). Immature osteoblasts express bone morphogenic protein 2 (BMP2), osterix (OSX), bone sialoprotein, and osteopontin. Mature osteoblasts are characterized by the expression of collagen type I [70].

Magnesium in bone metabolism Magnesium (Mg) is an essential component of the human body. The Mg ion (Mg2+) is a cofactor in more than 300 enzymatic reactions: it is a determinant in the synthesis of proteins and of DNA and RNA; in ion transportation; in cell migration and function; and in intracellular energy production via the adenosine triphosphate (ATP) system [71-73]. About 60% of the body’s Mg is located in bone; Mg is thus highly influential in mineral and matrix bone metabolism, affecting both bone cell function and hydroxyapatite (HA) crystal formation [74]. Mg is involved in the mineralization process of bone formation, particularly in immature stages, and is suggested to be responsible for the regulation of osteoblast and osteoclast activation [75]. Many studies have demonstrated a correlation between Mg and bone mineral density. Mg deficiency in various animal models leads to impaired bone growth, induced osteopenia, and increased

30 30 skeletal fragility [76, 77]. Reduced Mg content was observed in women suffering from osteoporosis, while larger HA crystals formed in trabecular bone becoming brittle, thus reducing bone stiffness [78, 79]. Low Mg intake also affects cartilage and bone differentiation [80]. Mg deficiency is associated with a reduction in parathyroid hormone (PTH) secretion and, consequently, with a decrease in vitamin D levels [81]. Indeed, Sahota et al. found lower levels of PTH and Mg in post-menopausal women who were vitamin-D defi- cient than in post-menopausal women who were not [82]. This strong correlation between Mg and bone health has led researchers to consider the element as a potential material for strengthening newly formed bone or fastening its apposition around a medical device.

Magnesium-based biomaterials Clinical applications of Mg and Mg alloys are most probably in the reconstruction of bone defects and the repair of bone fractures, or possibly as a bioactive incorporated substance for oral implant surfaces.

Magnesium and magnesium alloys for medical applications Metals that can degrade in the physiological environment have been recently introduced as potential materials for oral maxillofacial and orthopaedic applications. As previously described, biodegradable polymer devices are generally considered to have insufficient mechanical properties, and non-biodegradable metal implants must often be removed when they are no longer necessary.

Mg and Mg alloys have shown promising potential as tailored biodegradable metals for medical implants, changing the paradigm of metal implant use in medical conditions [83]. In 1907, Lambotte was the first to attempt to use a pure Mg plate in lower leg fracture healing, comparing its performance to that of a gold-plated steel nail – the standard treatment at that time. The Mg plate attempts were unsuccessful due to rapid degradation in vivo, disappearing only 8 days after implantation and generating a high amount of hydrogen gas. However, Lambotte demonstrated that pure Mg within cortical bone constituted a biocompatible system [84]. Mg

31 31 and its alloys are attractive because they possess the biodegradable nature of polymers, which were the earliest and most commonly used materials for such applications, and mechanical properties similar to cortical bone, such as density and elastic modulus. Table 1 lists the mechanical properties of some of the materials being used for orthopaedic implants today, according to ASTM specifications.

Table 1. Mechanical properties of cortical bone and some of the materials being used in orthopaedic applications today. (ASTM = American Society for Testing and Materials).

Tissue/Material Density Elastic Ultimate Yield (g/cm3) modulus strength strength (GPa) (MPa) (MPa)

Cortical bone 1.8–2.0 5–23 283 240

Stainless steel (316L) 7.9 193 558 290 Cobalt-Chrome alloy 8.3 210–250 655 450 (ASTM F-75)

Commercially pure 4.5 105 550 483–655 titanium, grade 4 (ASTM F-67)

Titanium alloy 4.4 114 860 795 Ti6Al4V (ASTM F-136) Polymethyl 1.1–120 1.8–3.3 38–80 45–107 methacrylate (ISO/ASTM 51276)

High-purity 1.7 45 90–190 20–115 magnesium Note: table compiled from references [85, 86] [87].

In aqueous environments, however, one concern in the biomedical use of Mg and its alloys is the evolution of hydrogen gas that accompanies degradation, according to the following reaction:

+ 2 ( ) +

푀푔 퐻2푂 → 푀푔 푂퐻 2 퐻2 32 32 Mg degradation alters the surface chemistry of the implant by producing degradation products of varying natures that dissolve and increase the pH of the peri-implant region, but also protect the material from further degradation [88].

Use of biodegradable Mg implants for osteosynthesis requires adequate maintenance of the device during the healing period. If the implant degrades too rapidly, hydrogen bubbles interpose between the implant surface and the surrounding tissues, impeding attachment of proteins and cells to the material surface and, thereby, pro- voking premature failure of the implant. The ideal course of degradation begins slowly and increases with time once the damaged tissue has healed sufficiently.

Several alloying elements have been combined with pure Mg to achieve moderate and homogeneous degradation behaviour. Witte and co-workers demonstrated that implant degradation depends on alloy composition and, in Guinea pig bones, the device should be present for at least 12 weeks to allow the healing [89]. The researchers alloyed pure Mg with cadmium to produce plates and screws for securing fractures. Of 34 cases, only 9 failed, possibly due to infection, while the rest were successfully tolerated by the host, preserving higher mechanical integrity than pure Mg. McBride et al. reported positive bone tissue reactions to magnesium-aluminium-manganese (Mg-Al-Mn) screws and pegs in the repair of fractures and bone grafts [90].

Aluminium (Al) has been considered to improve the strength and wear resistance of Mg, however, Al leads to the depletion of phosphate in tissues, a factor in the progression of senile dementia [89, 91]. Essential physiological elements, such as calcium (Ca) and zinc (Zn) have also been used in Mg alloys, but their effect on reducing the degradation rate and improving mechanical strength is mild. Nanda and Saravanan designed and tested binary Mg-silver [Ag] alloys with varying percentages of Ag in vitro and in vivo [92]. Ag proved to tailor the degradation process and the mechanical properties of Mg, showing no cytotoxic effect on mammalian cells [93, 94]. Other elements,

33 33 such as lithium (Li) and zirconium (Zr) have been combined with Mg in attempts to improve the mechanical properties of Mg [95].

The best outcomes in terms of bone response were obtained from in vivo applications of Mg implants containing rare earth elements (REE) [89]. REEs are a group of 17 elements which are mainly used to improve ductility, degradation resistance, and grain boundary strength in Mg. Johnson et al. reported that bone marrow-derived mesenchymal stem cells adhered to magnesium- yttrium (Mg-Y) surfaces in a homogenous and healthy state [96]. Feyerabend and co-workers elucidated the effect of some REEs on viability, apoptosis, and inflammatory cytokine expression of four cell types in vitro. The results revealed that lanthanum (La) and cerium (Ce) induced the highest cytotoxicity in the group while gadolinium (Gd) and Y seemed to be suitable for promoting cell growth [97]. Witte et al. explored the performances of two Mg- REE rods, comparing them to two Mg-Al-Zn rods inserted in Guinea pig femur and retrieved at 6 and 18 weeks of healing. Analyses of the degradation layer within the bone found that the REEs were localized in the degradation layer and had not diffused into the surrounding tissue. Moreover, high levels of Ca and P were detected at the interface between the degradation layer and the newly formed bone, indicating the possible formation of amor- phous apatite [89].

The initial degradation rate of Mg alloys, however, remained too high and localized in vivo, even in cases of Mg4Y3RE which, up to now, represents the most promising Mg-based biodegradable material for bone applications. Thus, demand for new Mg alloy compositions continues to increase, along with a need for a fundamental understanding of their degradation process in physiological conditions, of their in vitro cytocompatibility, and of their in vivo tissue response.

Some Mg alloys have been tested as possible open-porous scaffolds for load-bearing applications in tissue engineering. Witte et al. cast one Mg-Al-Zn alloy and one Mg-Al-Zn-Mn alloy as porous metallic scaffolds, which they tested in vivo; however, degradation

34 34 was still too rapid, and traces of Al and Zn triggered inflammatory reactions [98]. In addition, Mg alloys have been produced in particle form for use as matrix composites in musculoskeletal and dental treatments. Some examples are an Mg matrix composite reinforced with hydroxyapatite (HA) [99] and injectable Mg- phosphorous [P] cements [100].

Magnesium-modified titanium dental implants Chemical surface modification of Ti dental implants is a promising strategy for improving the biochemical bonding of implants with bone [41, 101, 102]. Bioactive ions, such as Ca, Mg, P, and fluorine (F), have been incorporated into Ti surfaces with different techniques [103-106]. In particular, Ca and Mg proved to be crucial in many biochemical mechanisms that occur at the bone- implant interface.

Several research groups have focused on the bioactivity of Mg- incorporated surfaces to elucidate the biochemical role of Mg ions in bone generation. Zreiquat et al. reported that osteoblast adhesion to biomaterials depends on an integrin-mediated mechanism, and that Mg supplementation to bioceramic substrates appeared to be involved in this molecular process [107]. It was shown in another report that Ti alloy modification with Mg contributed to human osteoblast function and differentiation [108].

A few years later, Revell et al. analysed interfacial bonding between bone and Mg-ion embedded HA coatings deposited on Ti implants, revealing the positive effect of Mg in bone cell activity compared with non-enriched HA surfaces [109]. Sul and co- workers examined in depth the influence of Mg on bone-cell responses, incorporating it into Ti implants via micro-arc oxidation (MAO) and, subsequently, with plasma immersion ion implantation and deposition (PIIID). In vivo experimental evidence demonstrated that Mg-incorporated implants promoted significantly higher bonding strength and faster osseointegration compared to non-incorporated CP Ti [110-112].

35 35 Park et al. investigated the impact of Mg-incorporated nanoporous Ti oxide surfaces on osteoblasts in vitro. The results showed that Mg-incorporated surfaces increased not only initial cellular adhesion but also subsequent cellular events such as alkaline phosphatase (ALP) activity and RNA expression of integrins and transcription factors [113].

36 36 AIMS

The overall aim of Studies I, II, and III was to assess Mg as a bioactive substance when loaded into a MP titania thin films produced via EISA, and considered as a novel surface modification for Ti implants. The effect of Mg release was tested on osteoprogenitors and osteoblast populations in vitro, with the aim of enhancing initial cell apposition, proliferation, and differentiation.

Specific aims (Studies I, II, and III): . To estimate Mg release profile in cell culture conditions (Study I). . To evaluate if Mg loading alters the surface morphology and topography of native MP coatings (Study I and III) . To investigate the osteoconductive potential of Mg release toward human foetal osteoblasts (hFOB) at early and late stages of growth (Studies I). . To investigate the ability of Mg to promote adipose-derived stromal cell (ADSC) differentiation toward the osteogenic lineage (Study II). . To investigate if increasing MP pore dimensions improves Mg loading and release and, thus, enhances hFOB development at even later stages of their proliferation (Study III).

The aim of Study IV was: . To evaluated the degradation behaviour of Mg2Ag, Mg10Gd, and Mg4Y3RE alloys as well as high purity Mg in controlled cell culture conditions. The short-term cell response of human umbilical cord perivascular cells (HUCPV) was investigated in terms of cell viability and adhesion structures, providing a preliminary indication of the suitability of these alloys as biodegradable metals for medical applications.

37 37 MATERIALS AND METHODS

Specimen preparation

Magnesium-loaded mesoporous TiO2 coatings (Studies I, II, and III) CP Ti (grade 4; Zimmer Holdings, Warsaw, IN, USA) discs were used in Studies I, II, and III. In Studies I and II, the discs were 15 mm in diameter and 1 mm in thickness; and in Study III, 12 mm and 1 mm. All specimens were soaked in absolute ethanol (≥ 99.8% )and rinsed in an ultrasonic bath before coating deposition. After each coating and Mg loading procedure, the discs underwent gamma radiation for sterilization.

Evaporation-induced self-assembly

MP TiO2 thin films were synthesized via EISA. Various structure- directing agents, amphiphilic surfactants or block-copolymers with defined chain lengths, tuned the pore size of the films. Together with an inorganic precursor, the agents assembled into micelles to generate ordered structures. The nature and ratio of each component as well as the self-assembly parameters, such as aging time and calcination, are fundamental to guiding the periodicity of pore distribution as well as their size.

In Studies I and II, MP films with an average pore size of 6.0 nm were generated onto non-porous (NP) Ti discs. To do this, a structure-directing agent, the amphiphilic block copolymer Pluronic®

P123 (triblock copolymer EO20PO70EO20; Sigma-Aldrich, St. Louis, MO, USA), were mixed with an inorganic precursor, Ti (IV) ethoxide (TEOT). A drop of final solution was deposited onto the Ti discs and spin-coated them at 7000 rpm for 1 minute. After spin-coating, the discs were stored at room temperature to allow

38 38 self-assembly to complete. Then the structure-directing agent was removed through calcination, which also increased condensation and the crosslinking density of the TiO2 matrix. Figure 3 illustrates the mesoporous TiO2 films deposited via evaporation-induced self- assembly and subsequently loaded with Mg.

Figure 3. Illustration of evaporation-induced self-assembly to form mesoporous TiO2 films (MP) and magnesium (Mg) loading through immersion in a magnesium chloride solution to yield the final magnesium-loaded mesoporous TiO2 film (MPMg).

In Study III, the synthesis parameters in EISA were varied to obtain average pore sizes of 2 nm (MP1) and 7 nm (MP3) in the MP films; this modification was done based on the results of Studies I and II for MP films with average pore sizes of 6 nm (MP2; see Results). For all films, the precursor was a TEOT solution, but the structure- directing agent that was mixed with the precursor solution varied.

Briefly: MP1 films were generated using Brij® S10 C18H37-

(OCH2CH2)nOH; Sigma-Aldrich); MP2 films using Pluronic P123 (Sigma-Aldrich); and MP3 films in a mixture of P123 with an organic additive, polypropylene glycol (PPG), which increased pore size. Table 2 describes the recipes for each group in detail, including component volumes. To achieve uniformity, all films were spin-coated at 7000 rpm for 1 minute. MP1 and MP2 films were than aged at room temperature, followed by calcination at 350°C; MP3 films were aged for 1 day prior to calcination in a sealed chamber containing saturated NaCl solution.

39 39 Table 2. Types and amounts of components used to synthesize the mesoporous (MP) titania films in Study III.

Group Surface-directing agent TEOT HCL Ethanol

(g) (g) (g) (g) MP1 Brij-S10 0.52 2.1 0.7 12 MP2 P123 0.5 2.1 1.6 8.5 MP3 P123+PPG 0.5 2.1 1.6 8.5

Brij® S10 = C18H37(OCH2CH2)nOH; P123 = Pluronic® P123; PPG = polypropylene glycol; TEOT = Ti (IV) ethoxide; HCL = hydrochloric acid

Magnesium loading Half of the coated discs of each group were soaked for 1 hour in a magnesium chloride (MgCl2) solution to load the discs with Mg and then dried in the oven at 100°C to evaporate the liquid. Therefore, Studies I and II discuss MPMg (the magnesium-loaded

MP TiO2 films of 6-nm pore size), whilst Study III discusses

MP1Mg, MP2Mg, and MP3Mg (the magnesium-loaded MP TiO2 films of 2-, 6-, and 7-nm pore size, respectively).

Magnesium alloys (Study IV) Three Mg-based materials were selected as test materials in Study IV: Mg2Ag (1.89% Ag, the remainder was Mg), Mg10Gd (8.4% Gd, the remainder was Mg), and Mg4Y3RE (3.45% Y, 2.03% Nd, 0.84% Ce, the remainder was Mg). High-purity Mg (99.97%) was used as a control. The discs had a diameter of 1 cm, thicknesses of 1.5 mm, and average weights of 0.2 g; the production steps were casting, homogenizing heat treatment, extrusion, turning, and cutting.

Casting and heat treatment Permanent mould gravity casting was used to produce Mg alloys (Mg2Ag, Mg10Gd, and Mg4Y3RE). After melting pure Mg, Mg was kept at 720°C and the preheated alloying elements were added under continuous stirring. The melt was poured into a pre-heated (550°C) permanent steel mould treated with boron nitride and T4 was selected as heat treatment to homogenize the alloys prior to

40 40 extrusion in an argon (Ar) atmosphere at 550°C (Mg10Gd and Mg4Y3RE) or 420°C (Mg2Ag) for 6 hours. High-purity Mg was cast by permanent mould direct chill cast.

Extrusion, turning, and cutting The alloys were extruded indirectly with an extrusion ratio of 4:25 (Strangpreßzentrum Berlin, Germany). The extrusion chamber was set to 370°C and the billets (diameter = 30 mm) were pre-heated for 1 hour at 370°C (Mg2Ag), 390°C (Mg4Y3RE) or 430°C (Mg10Gd). Extrusion speed was between 3 and 4.5 mm/sec.

The cast billet (diameter = 110 mm) of high-purity Mg was extruded indirectly with an extrusion ratio 1/84. The temperature of the billet was 340°C and extrusion speed was 0.7 mm/sec. Discs (10 mm in diameter and 1.5-mm thickness) were machined from the extruded bars (Henschel KG, Munich, Germany) and sterilized with gamma radiation.

Surface characterization Scanning electron microscopy (Studies I and III) In scanning electron microscopy (SEM), an electron source generates an electron beam that then scans the sample surface. When the electrons penetrate the outer layer of the material, they transfer energy to electrons inside the sample (secondary electrons) that scatter in various directions depending on the topography of the sample. Detection of the emitted electrons generates a topographical image of the material surface.

In the present thesis, high-resolution SEM images (LEO Ultra 55 FEG scanning electron microscope; Zeiss, Oberkochen, Germany) under an accelerating voltage of 5 KV were obtained to evaluate the thickness and pore direction of the MP TiO2 coatings in Studies I and III.

Optical interferometry – Study I Optical interferometry uses a light beam (λ 550 nm) with detection accuracy within a few millimetres. The beam is split into two separate beams: one beam hits the material surface and reflects; the

41 41 other beam passes through a reference plane. Irregularities in the experimental surface being investigated cause a phase shift in the reflected light; by comparing the reflected beam with the reference beam, a topographical image of the experimental surface can be generated. This technique is suitable for analysing surface microstructures due to its limited resolution in the spatial direction (vertical resolution of 0.005 µm; lateral resolution of 0.3 µm).

In Study I, MicroXamTM interferometer (ADE Phase shift technology, Inc., Arizona, USA) was used to examine micrometre differences in the topographies of MP and MPMg coatings and processed images with MountainsMap® 6 software (Digital Surf, Besançon, France), and a 50 x 50 µm Gaussian filter.

Atomic force microscopy – Studies I, II, III, and IV Atomic force microscopy (AFM) uses a flexible cantilever to measure forces between the cantilever tip and the sample surface. The attractive or repulsive force between the tip and the sample causes the cantilever to deflect, changing the angle of the laser beam. A photodetector records the angles of the laser beam and calculates the deflection signal. Resolution with this methodology is high for both height and spatial direction, thus it is often used to measure nanometre surface topographies.

All studies in the present thesis used AFM (XE-100; Park Systems Corp, Suwon, Korea) to investigate surface nanotopography. The 1x1 µm and 10x10 µm images underwent levelling and Gaussian filtering with cut-offs at 0.25 µm and 2.5 µm, respectively, in MontainsMap 6 software (Digital Surf). The following 3-D parameters were extracted from the data and analysed in order to de- scribe surface topography [40]: the arithmetic mean of the height deviation from the mean plane (Sa), the number of peaks per unit of area (Sds), and the ratio between the developed surface area and a flat reference area (Sdr).

X-ray photoelectron spectroscopy – Study I, III X-ray photoelectron spectroscopy (XPS) analysis is a technique that permits chemical elements in the outermost 5–10-nm thick

42 42 layer of a surface to be identified and quantified. XPS works by irradiating a material with X-rays, causing electrons to be ejected. The kinetic energy and the intensity of the ejected photoelectrons allow the identification of the relative concentration of the element in the sample outer layer.

For the present thesis, a Quantum 2000 scanning XPS microprobe (Physical Instrument, Marlborough, USA) was used to determine the chemical composition of Mg-loaded MP TiO2 films in Studies I and III. The monochromatic Al Kα X-ray source in the Quantum 2000 had a beam size of 100 μm and a take-off angle of 45°. The outer layer was analysed to a depth of about 5 nm with analysis point diameter of 100 μm.

Energy-dispersive X-ray spectroscopy – Study IV Energy-dispersive X-ray spectroscopy (EDX) is an analytical technique for the elemental characterization of a sample. When an X-ray beam that is focused on a sample surface excites the electrons in the atoms of the sample outer layer, the electrons pass from an inner to an outer atomic shell. The difference in energy between the higher-energy (inner) shell and the lower-energy (outer) shell is released as an X-ray, which corresponds to a specific element peak in the EDX spectrum.

Study IV investigated the chemistry of corroded Mg alloy surfaces. EDX detector (Apollo XP, EDX; Ametek GmbH, Wiesbaden, Ger- many) was used to analyse incubated samples after 72 hours of immersion in cell culture medium and in cell culture incubator. The Apollo XP was equipped with an SEM (Auriga; Zeiss, Oberkochen, Germany). An accelerating voltage of 10 keV with a secondary electron detector (SE2) was used to make surface measurements; the aim was to analyse the corroded surface and avoid the high contribution of Mg in the bulk of the sample. EDX quantification of the samples found eight common elements of interest; Mg, O, P, C have been considered despite of the alloying elements in each alloy.

43 43 Magnesium release Magnesium detection kit – Study I In Study I, extracts from NP, MP and MPMg were obtained, following ISO 10993-12-2012 (Biological Evaluation of Medical Devices - Part 12: sample preparation and reference materials) guidelines. Samples were incubated in Dulbecco’s modified eagle medium (DMEM; PAA Laboratories GmbH, Austria) with a surface/volume ratio of 3 cm2/1 mL for these time points: 3, 12, 24, 48, and 72 hours. Incubation was carried out at 37°C under constant shaking to simulate conditions in the human body. Mg assay kit (Abcam, Cambridge, UK) quantified the Mg that was released. The assay utilizes an enzyme-linked reaction that forms an intensely coloured product (450 nm) proportional to the concentration of Mg. Sample extracts and standard solutions were mixed with a Mg reaction mix containing Mg enzyme and the plate was read at λ=450 nm.

Quartz crystal microbalance with dissipation monitoring – Study III Quartz crystal microbalance with dissipation monitoring (QCM-D) is a mass-sensitive analytical technique that is used for in-situ monitoring of the adsorption and release of, for example, ions, molecules, and proteins, onto an oscillating quartz crystal. Voltage applied to two gold electrodes, one on either side of the crystal, causes it to oscillate. This technique provides information on nanogram-sized changes in mass, which are recorded as frequency shifts (∆f). These frequency changes over time can be converted into mass adsorption-desorption with the Sauerbrey equation:

= × × ퟏ where: ∆풎 −푪 ∆풇 ∆m = adsorbed-desorbed mass풏 C = Mass sensitivity constant (17.7ng/(cm2xHz) ∆f = change in frequency (Hz) n = selected overtone

44 44 In Study III, quartz crystal discs were coated with the three MP titania coatings (MP1, MP2, MP3) and used QCM-D (Q-Sense AB, Gothenburg, Sweden) to monitor the adsorption and release of Mg. Controls were non-coated quartz crystals. A Mg adsorption- release profile was created by loading MgCl2 solution into the system to quantify adsorption and changing the rinsing solution to Milli-Q water to quantify release. All frequency shifts were based on data recorded at the 7th overtone; data were processed using Q- Tools software (Q-Sense AB).

Material degradation parameters Pre-incubation – Study IV To evaluate the degradation parameters and cellular behaviour and adherence to the materials in varying degradation states, Mg alloys as well as hp Mg were first pre-incubated for 24, 48, and 72 hours. Extraction volume was calculated according to ISO 10993-12- 2012 (Biological Evaluation of Medical Devices - Part 12: sample preparation and reference materials) guidelines; 3 mL of alpha modified eagle’s medium (MEM, Thermo Fisher Scientific, Waltham, MA, USA) supplemented with 15% foetal bovine serum (FBS, Thermo Fisher Scientific, Waltham, MA, USA) and 5% penicillin/streptomycin were added to each sample. The samples were then incubated at 37°C in a 5% CO2 and 95% controlled humidity atmosphere.

Mass loss method – Study IV After pre-incubation the degradation products were removed by soaking the samples in chromic acid (VWR International, Darmstadt, Germany) at room temperature. The degradation rate for Mg alloys (mm per year) was calculated as follows:

. × = . . ퟒ ퟖ ퟕퟔ ퟏퟎ 휟품 where: 푫푹 Δ = weight change (g)푨 풕 흆 = sample surface area (cm2) 푔 = immersion time (hours) ρ퐴 = density (g/cm3) 푡

45 45 The advantage of this method is that measurements at different time points allow variations in degradation rates to be determined.

Osmometer and pH meter – Study IV An osmometer measures the total osmolality of aqueous solutions by making comparative measurements of the freezing points of pure water (0°C) and of solutions.

A pH-meter measures the electro-chemical potential between a known liquid inside the membrane and an unknown liquid outside.

In Study IV, the pH (Sentron Argus X pH-meter; Fisher Scientific GmbH, Schwerte, Germany) and osmolality (Osmomat 030; Gonotec GmbH, Berlin, Germany) of the corrosion medium were measured before and after each pre-incubation.

Cells Human foetal osteoblasts - Studies I and III American Type Culture Collection (ATCC, Manassas, VA USA) supplied the human foetal osteoblasts (hFOB 1.19; ATCC® CRL- 11372TM) used in Studies I and III. These cells are immortalized with a temperature-sensitive mutant that allows them to rapidly divide at the permissive temperature of 33.5°C. After reaching confluence, the hFOB undergo differentiation to form mineralized nodules in the final stages. Thus, this cell line provides a suitable and rapidly proliferating model for studying various stages of normal human osteoblast proliferation and differentiation.

Human adipose-derived stromal cells – Study II Tissue samples obtained from abdominoplasty procedures in the same group of patients (age 40–55; body mass index: 30–38) supplied the human adipose-derived stromal cells (ADSC) used in Study II. All patients signed informed-consent forms in accordance with the Helsinki Declaration before their inclusion in the study. The Ethics Committee of Padova Hospital (Italy) approved the research protocol. Human adipose tissue represents a particularly good source of multipotent stem cells that show the same fibro- blastic morphology, phenotype, and in vitro differentiation poten-

46 46 tial of stem cells isolated from bone marrow cavity. Thus, from a clinical perspective, human adipose tissue is a more accessible source of stromal cells, due to a less invasive approach.

Human umbilical cord perivascular cells – Study IV Wharton’s jelly (WJ) of umbilical cord samples supplied undifferentiated HUCPV cells after approval from the local ethics committee Ethik-Kommission der Ärztekammer Hamburg (Hamburg, Germany), and following the protocols of Sarugaser et al. [114]. These cells belong to mesenchymal stem cell populations that give rise to WJ connective tissue. They possess pluripotent plasticity and proliferate faster than adult MSCs.

Cell isolation and expansion Before the start of the studies, the primary human-derived cells were isolated and expanded under appropriate cell culture conditions to achieve optimal densities. Cell line was expanded according to the manufacturer´s protocol.

For Studies I and III, hFOB were expanded in calcium- and Mg- free medium DMEM/high glucose 1x (Thermo Fisher Scientific), completed with 10% FBS, 4 mM L-glutamine and supplemented with 1% penicillin/streptomycin solution in humidified atmosphere of 5% CO2 at 33.5°–34°C.

For Study II, ADSC were extracted from human adipose tissue. Samples of human fat tissue were pooled and soaked in DMEM (Euroclone, Milan, Italy). They were then digested with a solution of collagenase from Clostridium histolyticum Type II (C6885-1G; Sigma-Aldrich) in Hank's Balanced Salts Solution and stirred at room temperature. After digestion, DMEM was used to quench collagenase activity, and the sample was centrifuged at 1200 rpm.

Centrifugation yielded separation into four phases: • a yellow oily surface layer • a layer of undigested material • an aqueous phase (DMEM and digestion solution) • a cell pellet consisting of ADSC and red blood cells

47 47 To remove any trace of undigested material, the cell suspension was filtered and the flowthrough was centrifuged to obtain a cell pellet to be seeded in flasks with DMEM. The seeded cells were left for at least 2 days in an incubator to join ADSC.

For Study IV, HUCPV were isolated from umbilical cord samples following the protocols of Sarugaser et al. [114]. The cord was cut into pieces of about 5 cm. The vessels were then isolated and tied together at the ends, creating a vessel loop. The loops were placed in cell culture flasks and cultured for 10 days in α-MEM supplemented with 15% FBS and 5 mL antibiotics. After outgrowth of cells from the tissues, the medium was changed every 2–3 days.

Cell seeding and culture In Studies I and III, upon reaching confluence, hFOB were trypsinized with trypsin-EDTA solution and seeded at a density of 104/cm2/100 µL on the samples. A small volume of cell suspension was added to each sample, and the samples were incubated for 30 minutes to allow early cell adhesion. Fresh medium was then added, and the cells were further cultured in an incubator, with medium renewal every 2–3 days.

In Study II, ADS cells were cultured at confluence; thereafter, they were detached with trypsin-EDTA, seeded at a density of 5 x 103/cm2/100 µL, and cultured in an incubator.

For all assays in Study IV, we first transferred the pre-incubated discs into agarose pre-coated well plates, added 50,000 cells in 50 µL to each sample, and then incubated the samples for 30 min to achieve early cell adhesion. Subsequently, we added fresh medium to the samples and cultured the cells for 24 h. Figure 4 illustrates HUCPV seeding on pre-incubated samples.

48 48

Figure 4. Schematic illustration of the experimental procedure for HUCPV seeding on pre-incubated Mg alloys for the in vitro investigations (Studies IV).

Cell morphology Scanning electron microscopy - Studies I, II, and IV Scanning electron microscopy (SEM) is the most commonly used form of electron microscopy for visualizing cells adhering to a substrate. After culturing, cells were immobilized in a fixation procedure before dehydration in a graded series of ethanol. Dehydration was necessary to replace the water inside the cells with 100% ethanol. Cells then underwent critical-point drying. This method replaces the liquid inside biological structures with a suitable inert fluid, such as CO2, whose critical temperature is just above ambient temperature. Thus, no surface tension effect occurs, and this allows cells to maintain their morphology without distortion. Samples were then mounted and covered them with a thin layer of conductive powder to make the cell monolayer electrically conductive prior to imaging.

In Study I, hFOB were fixed after for 1 and 24 hours of cell culture into specimens, dehydrated the cells in graded ethanol, and then dried them at the critical point in a Balzers CPD 010 critical point dryer (BalTec, Pfäffikon ZH, Switzerland). Last, micrographs of the cultured cells were taken with SEM (JSM-7400F; JEOL Ltd., Tokyo, Japan).

In Study II, ADSC were fixed after 5 and 24 hours of culture, dehydrated the samples in ethanol, and dried them at the critical

49 49 point in a Balzers CPD 030 (Leica Microsystem AB, Wetzlar, Germany). Micrographs were the took in an SEM (JEOL JSM- 6490; JEOL Ltd.). The cell area of 5-hour-old cells was then calculated at 500x magnification for each sample type with a Fiji- MacOSX program.

In Study IV, HUCPVC after 72 hours of cell culture onto the samples, dehydrated the cells in a graded series of propanol, and dried the cells at the critical point in the CDP chamber of a Leica EM CPD300 (Leica Microsystem AB, Wetzlar, Germany). Images of the cells grown on samples were taken with SEM (Auriga, Carl Zeiss AG, Jena, Germany).

In vitro cytotoxicity In vitro cytotoxicity tests are widely used for evaluating medical devices and biomaterials. ISO 10993-5-2009 (Biological Evaluation of Medical Devices – Part 5: Tests for in vitro cytotoxicity) specifies the criteria and methods for determining the biological response of mammalian cells to medical devices in vitro.

MTT colorimetric assay - Studies I and II The assay determines the viability of cells via metabolic activity. Proliferating cells will reduce yellow, water-soluble 3-(4,5- Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) to blue-violet insoluble formazan, which absorbs at 570 nm. The number of viable cells correlates to spectrophotometric assessment of colour intensity.

In Study I, hFOB were treated with MTT after 24 hours of cell growth onto the samples and adsorbance was measured at 570 nm with a micro-plate spectrophotometer (BMG Labtech GmbH, Ortenberg, Germany); the reference wavelength was 650 nm.

In Study II, ADSC were incubated with a working solution of MTT after 24 and 48 hours of cell culture onto the samples and adsorbance at 570 nm was measured with a spectrophotometer (Victor multilabel plate readed, PerkinElmer, Waltham, MA, USA).

50 50 LIVE/DEAD staining - Study IV LIVE/DEAD staining allows simultaneous detection of living and dead cells, via staining with a mixed solution of Calcein AM (LIVE) and Ethidium homodimer-1 (DEAD). Samples were washed with PBS to eliminate non-adherent cells, and immersed in staining solution. The staining solution was then replaced with -αMEM and visualized the samples under a fluorescent microscope (Nikon GmbH, Düsseldorf, Germany). The applied filters were FITC (Ex: 460-500 nm; Em: 510-560 nm; Mirror at 505 nm) and Texas red (Ex: 540-580nm; Em: 600-660; Mirror at 595 nm). Images were taken in large-image mode (4X4) at 20x magnification. The intensity profile of FITC (green) and Texas red (red) was quantified with the microscope imaging software NIS-Element AR (Nikon, GmbH, Düsseldorf, Germany). The results were normalized as follow:

live cells sample Viability (%) = live cells + dead cells × 100 live cells control �live cells + dead cells� � � Cell adhesion Immunocytochemical staining makes it possible to visualize the distribution and localization of specific cellular components within cells in the proper tissue context. The procedure has two parts: sample preparation and labelling (Figure 5).

51 51

Figure 5. Fluorescence image of human umbilical cord perivascular cells adhered onto a metal surface. Cells display nuclei (blue-2-(4-Amidinophenyl)- 1H-indole-6-carboxamidine DAPI staining) pointed by arrows, and actin filaments (yellow-FITC+ Tetramethyl-rhodamine isothiocyanate conjugated phalloidin TRITC stainings) that elongate along each direction.

TRITC-phalloidin staining - Study I In Study I, TRITC-phalloidin (Sigma-Aldrich) was used to label cells after 48 hours of cell culture onto the samples. TRITC- phalloidin is a fluorescent phalloidin conjugate that binds to actin filaments of fixated cells, thus making the filaments visible under fluorescent microscopy at 550 nm.

Actin cytoskeleton and focal adhesion staining - Study IV Study IV used the Actin Cytoskeleton and Focal Adhesion Staining Kit FAK100 (Millipore, Billerica, MA USA), which includes three components (TRITC-conjugated phalloidin, anti-Vinculin, and 2- (4-Amidinophenyl)-1H-indole-6-carboxamidine [115]) for staining actin filaments in the cytoskeleton, the focal contacts, and cell nuclei respectively. After 24 hours of cell culture onto each pre- incubated sample, the samples were carefully transferred to a new plate and stained the cells according to the manufacturer’s protocol.

52 52 Alkaline phosphatase Alkaline phosphatase activity assay - Study II ALP catalyses the hydrolysis of phosphate esters, yielding organic radicals or inorganic phosphates. ALP is an important enzyme in hard tissue formation, highly expressed in mineralized tissue cells during early developmental stages. The assay uses a colorimetric reaction to measure ALP activity directly inside biological samples.

Study II measured intracellular and extracellular ALP activity after 2 weeks of cell culture onto the specimens, to evaluate the initial differentiation of ADSC into immature osteoblasts. Alkaline phosphate kit (ab83369 kit, Abcam, Cambridge, UK) was used in this investigation. The kit uses p-nitrophenyl phosphate (pNPP) as a phosphatase substrate. which adsorbs at 405 nm when dephosphorylated by ALP. Optical density was read at 405 nm in a microplate reader (Victor, PerkinElmer). PNP standard curves was plotted to identify the pNPP concentration and then calculated ALP activity in the samples as follows:

ALP activity (U/mL) = A/V/T where: A = amount of pNPP generated by samples (in μmol) V = amount of sample added in the assay well (in mL) T = reaction times (in minutes)

Cell mineralization Alizarin red staining - Study I Differentiated osteoblasts generate extracellular calcium deposits in vivo and in vitro. Thus, calcium deposits indicate that differentiation of immature osteoblasts into mature cells was successful.

In Study I, alizarin red solution was added to each sample where cells were grown for 3 weeks, and incubated for a minimum of 20 minutes. Thereafter, the excess dye solution was removed, examined the cells under an optical microscope (Eclipse TS100,

53 53 Nikon Instruments, Amsterdam, Netherland) and took images. The alizarin red solution stained differentiated cells containing mineral deposits.

RNA extraction In Studies I, II and III, total RNA was extracted for gene expression analyses.

In Study I, total RNA was isolated from 2-week-old hFOB cultures using the Maxwell® 16 LEV simplyRNA Tissue kit (Promega Corp., Madison, WI USA). In brief, cells were disrupted by lysis buffer and homogenized. The lysate was loaded into a Maxwell® 16 instrument and the system automatically isolated the cells. Thereafter, RNA was automatically eluted with nuclease-free water and quantified with NanoDrop ND-1000 Spectrophotometer (Thermo Fisher Scientific).

In Study II, ADSC differentiation was assessed following two protocols: Protocol 1. For ADSC cultured for MTT analysis, each coating was scraped off to obtain a powder, which it was assumed had previously adsorbed the osteogenic markers during cell culturing for the cytotoxicity assay. We then added the scrapings to regular medium (medium without osteogenic additives) and cultured the ADS cells for 2 weeks. Controls were ADS cells cultured on plastic wells with complete osteogenic differentiation medium, a condition where differentiation factors were expressed 100%. Table 3 presents medium compositions for test and control conditions.

54 54 Table 3. Medium composition of osteogenic (control) and non-osteogenic (test) culture conditions.

Control Test (osteogenic) (non-osteogenic) DMEM High Glucose DMEM High Glucose 10% FBS 10% FBS 1% antibiotic/antimycotic 1% antibiotic/antimycotic + + Dexamethasone (10nM) powdered mesoporous Ascorbic acid (50 μg/mL) coating (from MP and Beta-glycerophosphate MPMg groups) after cell (10mM) growth FGF-2 (10ng/mL) DMEM = Dulbecco’s modified eagle medium; FBS = foetal bovine serum; FGF-2 = Fibroblast growth factor-2; MP = mesoporous; MPMg = mesoporous with added magnesium.

Protocol 2. ADSC were seeded directly onto the disc surface and cultured for 3 weeks to determine effective osteogenic differentiation. Controls were ADSC cultured on non-porous discs, to evaluate both the effect of added Mg on the MP surface and of surface topography. We isolated total RNA using the RNase Mini Kit (Qiagen, Inc., Hilden, Germany) following the manufacturer’s protocol. Briefly, cells were lysed with RLT lysis buffer, which contains guanidine thiocyanate, which is a general denaturant, and β- mercaptoethanol, which reduces sulphide bridges in protein structures. Samples were then transferred to the RNase Spin column, one for each sample, and extracted the RNA according to the manufacturer´s instructions. Concentration (OD at 260 nm) and purity (OD260/280) of the eluted RNA were quantified with a NanoDrop (Agilent Technologies, Santa Clara, CA, USA).

In Study III, RNA was isolated from the hFOB populations after 14 and 21 days of cell culture using the AurumTM total RNA fatty and fibrous tissue kit (Bio-Rad Laboratories, Inc., CA, USA), according to the spin format (column) protocol. Cell lysis was done with PureZOL, a ready-to-use reagent, designed to isolate high-

55 55 quality RNA (also DNA and proteins) from liquid samples of various natures. PureZOL maintains the integrity of the RNA due to the highly selective inhibition of RNase activity during sample homogenization. Chloroform was then added to each lysate, incubated the samples at room temperature, and centrifuged (Eppendorf 2420R, Hamburg, Germany) at 12,000 rpm. After centrifugation, the colourless aqueous phase at the top of each column (RNA interphase) was transferred into an RNA binding column and processed according to the manufacturer’s protocol. RNA was eluted with the elution solution and quantified the RNA in a UV spectrophotometer (Thermo Fisher Scientific).

DNase digestion was done for each RNA extraction to eliminate any contamination from genomic DNA.

Gene expression techniques Polymerase chain reaction (PCR) is a sensitive analytical technique that amplifies a DNA segment or a gene that codes for a specific protein, growth factor, enzyme, receptor, and many other molecules from a pool of DNA.

The PCR reaction mixture is made of a single-strand DNA template that derives from denaturation of the double-strand DNA in the sample, specific primers, nucleotides, and DNA polymerase enzyme; these synthesize the PCR products through assembly of single nucleotides. The DNA primers are short fragments of single- strand DNA that are complementary to DNA sequences in the target DNA, amplifying the specific product of interest. A defined volume of this reaction mixture is then loaded in 96-well plates or in tubes and inserted in a thermal cycler. This machine modulates temperature in different steps, according to defined PCR run protocols. In the first step, denaturation, the temperature rises above the DNA melting point, in order to separate the DNA into its two single strands. In the second step the temperature is decreased to an optimal annealing temperature that allows the primers to bind to specific regions of the target DNA. In the third step, DNA polymerase extends the double strand at higher temperatures. Figure 5 illustrates the PCR principle. The final

56 56 results indicate how much of a specific gene is expressed in one sample compared to another. We used two analytical techniques to analyse gene expression in cell populations: Real-time polymerase chain reaction in Studies I, II and III (RT-PCR) and pre-designed osteogenesis super array in Study III.

Figure 5. Schematic representation of the polymerase chain reaction process: PCR steps correspond to specific temperatures and specific times.

Real-time polymerase chain reaction - Studies I, II, and III In Studies I, II, and III, we used RT-PCR to investigate the expression of the main genes involved in osteogenesis. After extraction, RNA was reverse transcribed using the M-MLV Reverse Transcription kit (Invitrogen, Waltham, MA, USA) in Study II, and with iScript Advanced cDNA Synthesis kit (Bio-Rad Laboratories) in Study III, according to the manufacturer’s instructions.

In Study I, primers for OCN, ALP, RUNX2, COL I, ITGα1 were selected and mixed them with components in the TaqMan® pre- developed gene expression kit to obtain the PCR solution. PCR was performed in a StepOnePlus PCR system (Applied Biosystems,

57 57 Waltham, MA, USA) in triplicate for each sample and determined gene expression by the comparative Ct method, normalizing expression to glyceraldehyde 3-phosphate dehydrogenase (GAPDH), the housekeeping gene, which was selected as reference gene.

In Study II, primers of OCN, OPN, osteonectin (ONC), RUNX2, and COL I were analyzed in the PCR system (Corbett Research Australia, Mortlake, Australia) according to the manufacturer´s instructions and using FastStart SYBR Green Master kit components (Roche, Basel, Switzerland) for the PCR solution. Relative expression levels of the target genes were determined by the comparative Ct method, normalizing them to GAPDH, the housekeeping gene.

In Study III, primers of BMP2, BMP4, MMP9 and TGFβ1 to verify super array data and analysed in an RT-PCR thermocycler (CFX ConnectTM Real-Time PCR System; Bio-Rad Laboratories). PCR solution was prepared using SsoAdvancedTM Universal SYBR® Green Supermix (Bio-Rad Laboratories, Inc.) and gene expression was determined by the comparative Ct method, normalizing each expression to GAPDH.

Predesigned osteogenesis super array - Study III PCR super arrays allow the expression of all genes in a pathway to be analysed in one PCR run. In Study III, 84 specific primers corresponding to the genes involved in osteogenesis were already distributed and dried into the wells of a predesigned 96-well plate. We used osteogenesis-related genes arrays (Osteogenesis [SAB Target List] H96 predesigned panel with SYBR Green; Bio-Rad Laboratories) to track the expression of genes, which could vary depending on the amount of Mg release from the three MP titania films. The PCR run was set at an activation temperature of 95°C for 2 min followed by 40 cycles of 95°C for 5 sec and 60°C for 30 sec in the RT-PCR thermocycler. Data were analysed using the comparative cycle threshold method with normalization of the raw data to five housekeeping genes: β-actin (ACTB), β2M, ribosomal

58 58 protein, large, P0 (RPLP0), hypoxanthine phosphori- bosyltransferase 1 (HPRT1), and GAPDH.

Statistics Each study compared the means obtained from at least three groups. Thus, the one-way ANOVA was applied to compare means among homogenous distributions of data, followed by the post-hoc Tukey test. Non-parametric Kruskal-Wallis test was applied in comparisons of means obtained from non-homogeneous distributions, followed by pair-wise comparison with the Mann- Whitney U-test. Means from homogenous populations were presented as the mean ± the standard deviation of the mean; in cases of non-homogeneous distributions, data were presented as the mean ± the standard error of the mean. In all studies, statistical analyses were done with the Statistical Package for the Social Sciences (SPSS 21 Inc., Chicago, IL, USA). In Study IV, some of the analyses related to the corrosion investigation were done with the SigmaStat package (version 11.0; Systat software GmbH, Erkrath, Germany), applying Dunn’s multiple comparison as a post hoc test for the analysis of ranks; the results form LIVE/DEAD assay were compared with one-way ANOVA test followed by post-hoc Tukey test for multiple comparison using with OriginLab Package (Origin software, version 9.0, MA, USA).

59 59 RESULTS

Magnesium as a bioactive substance for mesoporous titania implant coatings In Studies I and II, the influence of Mg adsorption on MP titania film morphology and topography with 6 nm average pore size was investigated. Mg release effectiveness was assessed on two cell populations to elucidate at which cell stage Mg presented a more prominent osteogenic effect. To adjust the Mg content, MP films with various pore sizes were produced and characterized in Study III, to evaluate if modification of pore dimension at a nano level could modulate Mg adsorption into and release from the MP matrix, further accelerating bone cell development on the surface.

Study I

Preliminary outcomes demonstrated that MP TiO2 coatings (MP) had an average pore size of 6.0 nm, and a porosity that was highly homogeneous and evenly distributed throughout the 200-nm thickness of the film (Figures 6 B and C).

60 60

Figure 6. Scanning electron microscopy images of a non-porous titanium surface (NP) (A), a mesoporous TiO2 film (MP) (B) and a mesoporous TiO2 film with a scrape (C). Note: pore distribution is homogeneous; average pore size = 6.0 nm; thickness = 200 nm.

Table 4 lists the atomic percentages of the surface elements. XPS demonstrated that Mg successfully adsorbed to the MP surfaces in the MPMg group with an element percentage of 8.6. Optical interferometry measurements (Table 5) confirmed that MP films possessed a smooth surface with an average Sa of 0.3 µm that did not alter significantly with Mg adsorption. However, the surface enlargement (Sdr) and the peak density of the sampling area (Sds) were significantly higher at the micr-scale level for MP and MPMg than the control (NP).

Table 4. X-ray photon spectroscopy data for mesoporous TiO2 films with Mg (MPMg) and without Mg (MP). Data are presented as relative atomic %.

Group C Ti O Mg Cl

MPMg 21.5% 13.0% 46.3% 8.6% 10.6%

MP 14.8% 24.5% 60.7% – –

61 61 Table 5. Mean values and standard deviations for interferometry measurements of roughness and p-values calculated with the one-way ANOVA and the post-hoc Tukey test; the significance level was set at p = 0.05.

Group Sa (µm) Sdr (%) Sds (1/µm) MP 0.32 (0.02) 9.64 (0.6) 115328.10 (8031.56) MPMG 0.33 (0.02) 10.19 (1.09) 105869.32 (7256.67) NP 0.31 (0.05) 7.79 (0.8) 102253.54 (9083.11) p-value 0.263 0.0* 0.008*

Figure 7 shows representative AFM 3-D reconstructions of MP and MPMg surfaces. Both coated surfaces appear highly smooth with no visible differences in average height deviation (Sa) from the added Mg (Table 6). However, the surface enlargement (Sdr) at 1x1 was significantly higher for MP and MPMg than the control.

62 62

Figure 7. AFM reconstructions of MP coating and MPMg coating, scan area = 1x1 µm. AFM = atomic force microscopy; MP = mesoporous; MPMg = Mg- loaded mesoporous

63 63 Table 6. Mean values and standard errors for AFM measurements of roughness at 1x1 and 10x10 µm and p-values calculated with the Kruskal-Wallis test and Mann-Whitney U-test for multiple comparisons; the significance level was set at p = 0.05.

-2 Group Sa (nm) Sdr (%) Sds (µm ) 10x10 1x1 10x10 1x1 10x10 1x1 MP 16.7 (8.0) 1.2 (0.7) 1.3 (1.4) 2.2 (1.1) 10.9 (4.3) 3728(944) MPMG 20.5 (5.4) 2.1 (1.1) 2.5 (1.7) 1.9 (1.5) 7.7 (3.1) 2966 (1511) NP 5.4 (5.2) 1.1 (0.5) 1.7 (1.9) 1.5 (0.3) 3.1 (1.2) 2424 (582) p-value 0.260 0.086 0.299 0.004* 0.001* 0.052

Mg release was measured after 3, 12, 24, 48, and 72 hours of incubation in culture medium, and its effect was subsequently tested on hFOB. Mg concentration was determined with a Mg- specific colorimetric assay. To obtain reliable data from the small amounts of Mg released from the Mg-impregnated matrices, the values at each time point were normalized to the normal Mg concentration in the medium and to the sample surface that was exposed to the medium during incubation. In the MPMg group, a peak in Mg release occurred after 24 hours of incubation in culture medium (Figure 8). This Mg release enhanced hFOB viability, which was significantly higher for MPMg compared to the control cells and not significantly lower than the surface control (NP). At 48 hours of hFOB growth, the effect of Mg showed no positive effect, although cells grown on native and Mg-loaded MP surfaces showed a cytoskeletal organization characterized by larger sizes and a higher distribution compared to NP surfaces.

64 64 Figure 8. Magnesium release from MPMg after 3, 12, 24, 48 and 72 hours of incubation in cell culture medium. (MPMg = mesoporous with Mg.)

hFOB viability after 24 hours of culturing on NP, MP, and MPMg. The control is repre- sented by cell cultures on plastic wells. (hFOB = human foetal osteoblasts; NP = non-porous; MP = mesoporous.)

The outcomes of long-term cell response showed that the release of Mg from 6-nm MP matrices did not significantly increase the expression of RUNX2, OCN, ONC, OPN, ITGα1 and collagen type-1 (COL1). Images obtained with phase microscopy indicated red-stained mineralized cell monolayers on all types of surfaces after 3 weeks of culture, and the red-stained layers covered more extensive areas on MP and MPMg surfaces than NP surfaces.

Study II

Mg release potential from 6-nm MP TiO2 films was examined on ADSC activity to elucidate whether it possessed the osteoconductive ability to promote osteogenic differentiation of undifferentiated cells.

SEM images show how cells attached to the surfaces of the different samples (Figure 9). High magnification images show that cells developed more and longer filopodia, particularly on MP surfaces than on the control. Images at 100x give a wider view of cell coverage on the different surfaces, showing that ADSC covered both coated surfaces better than the non-coated one.

65 65

Figure 9. SEM images of ADSC cultured on NP, MP and MPMg (from left to right) for 24 h. Images were obtained at 5000x and 100x. (ADSC = adipose- derived stromal cells; NP = non-porous; MP = mesoporous; MPMg = magnesium-loaded mesoporous).

The effect of pore capacity in adsorbing cellular growth factors was tested by isolating RNA according to protocol 1 (see Materials and Methods). At 2 weeks, besides COL1 expression, no significant enhancement occurred in any of the three groups; OCN, ONC, and in particular OPN (p-value = 0.055), however, showed a tendency for upregulation in test culture conditions, especially in the MPMg group, compared to the control group (Figure 10). COL1 was significantly downregulated in cells cultured under test conditions compared to the control group. At 3 weeks of cell growth according to protocol 2 to explore the osteoconductive potential of the surface topography of the samples, the obtained results corroborated the outcomes from protocol 1, particularly regarding the expression of OPN; at 3 weeks of direct cell contact with the MPMg surface, OPN expression was significantly higher compared to the MP and NP groups.

66 66

Figure 10. p-values (Kruskal-Wallis and Mann-Whitney U-test for multiple comparisons; significance levels, p = 0.05) for mRNA levels of osteocalcin (OCN), osteonectin (ONC), osteopontin (OPN), runt-related transcription factor 2 (RUNX2) and collagen type-1 (COL1) after 2 weeks (left, cultured according to protocol 1) and 3 weeks (right, protocol 2) of growth of adipose- derived stromal cells (ADSC).

Study III It appeared that improved Mg adsorption in the mesopores was necessary to maximize the osteogenic effect. To further explore mesoporous Mg-delivery property, the average pore size of the MP films was tuned to modulate the amount of Mg that can be hosted into the mesoporous structure. By using different structure- directing agents and by the absence or presence of a swelling agent (PPG) in the precursor solution it was possible to generate films with 2-nm (MP1) and 7-nm (MP3) pore sizes, comparing their Mg- adsorbing properties with 6-nm pores films (MP2) previously investigated (Study I and II). Surface examinations revealed morphological and topographical similarities among the three MP surfaces. All films covered the Ti disc surfaces completely and possessed high degrees of periodically and homogeneously distributed porosity.

67 67

Figure 11. Scanning electron (A, B, C) and atomic force microscopy reconstructions (D, E, F) of mesoporous films with average pore sizes of 2 nm (MP1: A, D), 6 nm (MP2: B, E) and 7 nm (MP3: C, F).

QCM-D monitoring of the dynamic adsorption and release of Mg into and from the three MP structures showed higher Mg adsorption into the 2- and 7-nm films in comparison to 6-nm films (Figure 12).

AFM measurements analysis of surface nano-topographies revealed no significant differences among the three native surfaces, nor between any native film and its Mg-loaded counterpart. Mg loading had no measurable effect on the average surface roughness

Sa.

68 68 Figure 12. Quartz crystal microbalance with dissipation monitoring (QCM- D) results of adsorption and release of magnesium from mesoporous thin films with average pore sizes of 2 nm (MP1), 6 nm (MP2), and 7 nm (MP3) and non-porous (control) films. Changes in resonance frequency (∆f) as a function over time were converted into mass adsorption and desorption (∆m) with the Sauerbrey equation.

To evaluate the correlation between Mg release and osteogenesis, 84 genes involved in osteogenic differentiation and mineralization were examined. After 1 week of hFOB proliferation onto MP films with and without Mg, Mg released from the 7-nm pore matrix produced much higher variability in gene expression compared to the other films (Figure 13). Moreover, this Mg-impregnated surface was the one that enhanced osteocalcin (OCN or BGLAP) expression, a non-collagenous protein secreted by mature osteoblasts.

69 69

70 70

Figure 13. Comparison of relative expression for 84 osteogenic genes between hFOB cultured on mesoporous (MP) titania films without and with magnesium ions and of 2-, 6-, and 7-nm pore size (MP1 vs MP1Mg, MP2 vs MP2Mg, MP3 vs MP3Mg, respectively) for 1 week. The relative difference in gene expression between native titania MP conditions (x-axis) and magnesium-loaded titania MP conditions (y-axis) are displayed in scatter plots. The red and green lines indicate a 3-fold change in gene expression threshold.

The gene profile of the MP3 matrix at 3 weeks of cell culture showed enhanced osteoblast activity in the MP3Mg group with upregulation of four genes that encode for two growth factors (insulin-like growth factor 1 [IGF1]) and growth differentiation factor 10 [GDF10]), one transcription factor (supernatant protein factor 7 [SPF7]), and one metalloproteinase (MMP8); no upregulation was observed in the native MP counterpart.

RT-PCR then quantified relative gene expression of four genes selected from super array outcomes. Figure 14 presents comparisons of MP1Mg, MP2Mg, and MP3Mg, showing the effectiveness of different amounts of Mg on long-term (3 weeks) hFOB activity. The results demonstrated that BMP4 was strongly

71 71 upregulated in cells grown on MP3Mg and significantly downregulated on MP1Mg.

Figure 14. Effect of magnesium (Mg) release on mRNA levels of four bone- related genes in human foetal osteoblasts (hFOB) cultured on Mg-loaded mesoporous thin films of 2-, 6-, and 7-nm average pore size (MP1Mg, MP2Mg and MP3Mg, respectively) for 3 weeks. The relative mRNA expression of each gene is expressed as fold-change compared with MP2Mg, designated as the calibrator sample; independent sample t-tests calculated p- values; the significance level was set at p < 0.05. (BMP = bone morphogenic protein; TGFB1 = transforming growth factor, beta 1)

Magnesium alloys as bioresorbable metals for bone applications Study IV The degradation parameters of Mg2Ag, Mg10Gd and Mg4Y3RE alloys were analysed after 24, 48, and 72 hours of incubation in cell culture medium under cell culture conditions. Mg4Y3RE and high-purity Mg showed high degradation rates during the first 24 h, which correlated with higher increases in pH and osmolality compared to Mg2Ag and Mg10Gd. After 48 h, the initial interaction between the material and the surrounding environment decreased, leading to a significant reduction in degradation rate by > 50% for the high-purity Mg samples. Mg10Gd showed a stable degradation rate over time with moderate increase of pH and osmolality. This homogenous

72 72 degradation behaviour was observed also for Mg2Ag samples but with lower values and lower effects on the surrounding environment (Figure 15).

Figure 15. Degradation rate (mm/year), pH and osmolality (osm/kg) after immersion for 24, 48, and 72 hours in corrosion medium under cell culture conditions. The basic solution is -αMEM with addition of 15% FBS and 1% penicillin/streptomycin. Bars mark significant differences in groups (significance level p <=0.05). At 72 hours, EDX spectra of all sample degradation layers shared the same trend in concentrations; basically, a high oxygen (O) concentration with contributions from carbon (C), phosphorous (P) and Mg in the outer degradation layer (Figure 16).

73 73

Figure 16. Scanning electron microscopy images with the corresponding energy dispersive x-ray spectroscopy analysis for magnesium (Mg) alloys and high-purity Mg (hp Mg). Analysis was done after 72 hours pre-incubation with 10 KeV accelerating voltage. (Mg2Ag = Mg-silver; Mg10Gd = Mg- gadolinium; Mg4Y3RE = Mg-yttrium-rare-earth). To explore the degradation modalities occurring at the surface, the surface topography was examined after each degradation period. The overall trend was characterized by a decrease in surface roughness for Mg2Ag and Mg10Gd between 24 and 48 hours, followed by a significant increase with a further 24 hours of degradation. In

74 74 particular, initial degradation of the Mg2Ag surface occurred with significant reductions in average roughness (Sa) and surface area

(Sdr) between 24 and 48 hours. Conversely, Mg4Y3RE and high- pure Mg showed no significant decreases in Sa, although they exhibited the same surface degradation trend as the other two alloys.

At 24 hours of culture, LIVE/DEAD staining cytotoxicity and actin cytoskeleton and focal adhesion staining were used to explore correlations of degradation state with HUCPV cell behaviour and the Mg alloys, compared with high-purity Mg. HUCPV cells showed good viability on all three Mg alloys as well as on Mg. Few red cells were detected on any sample, at any incubation time (Figure 17). The LIVE/DEAD staining quantification is shown in figure 18 and demonstrated that no significant differences were detected among the samples at each pre-incubation time. However, cell viability slightly increased for Mg2Ag and Mg4Y3RE between 24 and 48 hours pre-incubated states, with a reduction below 60% for Mg2Ag at 72 hours of incubation. Stable cell viability was observed on Mg10Gd as the pre-incubation time increased.

75 75 Figure 17. Fluorescent images showing viable cell populations of human umbilical cord perivascular cells cultured on high-purity magnesium (Mg), and on magnesium-silver (Mg2Ag), Mg-gadolinium (Mg10Gd), and Mg- yttrium-rare-earth (Mg4Y3RE) alloys. Fluorescent LIVE (green)/DEAD (red) staining was done after 24 hours of cell culture on 24, 48, and 72 hour pre- incubated samples. Monochromes: large image mode (4x4) at 20x magnification.

76 76

Figure 18. FITC intensity quantification for HUCPV cultured for 24 h on Mg, Mg2Ag, Mg10Gd and Mg4Y3RE at 24 (1 day), 48 (2 days) and 72 (3 days) hours pre-incubation states. Measurements were obtained from images taken form 3 samples of each material at each pre-incubation condition after LIVE/DEAD assay. Significance level was set at p < 0.05; n=3 per group. Micrographs at 20x magnification were used to compare cell growth and adhesion structures (Figure 19) of the cells on 24- and 48-hour pre-incubated samples. On Mg2Ag, actin structures appeared more developed after 48 hours of degradation, while only weak yellow spots were visible in the 24-hour image. A similar situation was observable for Mg10Gd: HUCPV cells covered a wider surface area in the 48-hour image while their growth was limited in the 24-hours image. However, cells showed a well- defined cytoskeletal structure and ovoid nuclei when grown on 24- hours pre-incubated Mg10Gd. Cells cultured on Mg4Y3RE exhibited a different behaviour; they appeared to decrease in density after a further 24 hour of degradation, but with elongated actin filament architecture still characterized cell structure. On high-purity Mg, cells were clustered together, and no obvious differences between the 24- and 48-hour cultures were visible.

77 77

Figure 19. Confocal fluorescence microscopy of focal adhesion and the actin cytoskeleton in human umbilical cord perivascular (HUCPV) cells cultured for 24 hours on 24- and 48-hours pre-incubated samples. Monochrome images at 20x were overlaid to display the triple labelling.

78 78 DISCUSSION

The cell–biomaterial interaction is a complex phenomenon that requires thorough understanding for predicting the safe use of new biomaterials. In vitro investigations of cellular adhesion to a material surface, cell proliferation and differentiation modalities are essential experimental steps to investigate for the improvement and future development of biomaterials. Irrespective of a material’s nature, the surface properties ultimately influence initial cellular events, determining cell fate and cell specific functions [12, 19, 20]. Cell culture models are routinely used to examine molecular and cellular responses to biomaterials, under highly controlled laboratory conditions.

In the present thesis, four in vitro studies analysed the impact of Mg on cell adhesion, proliferation, and osteogenic differentiation to determine whether this bioactive element could be considered as substance to be implemented in Ti implant surfaces, and as biodegradable alloys for possible alternative to existing metallic bone applications.

Studies I, II, and III investigated the osteogenic potential of local delivery of Mg from mesoporous titania films toward osteoprogenitors and immature osteoblasts. In addition, the mesoporous design was modulated to achieve sustained release of Mg over time.

Study IV examined the degradation properties of three biodegradable Mg alloys in cell culture conditions. The biological impact of material degradation was tested on the undifferentiated human cell response.

79 79 Magnesium-modified mesoporous titania films to enhance peri-implant osteogenesis - studies I, II and III One strategy to accelerate the bone-forming capacity of Ti implants is to enhance surface bioactivity by modulating the surface chemistry and nanotopography, without excessively altering surface microtopography [20-25]. The deposition of mesoporous TiO2 thin films was recently proposed as a novel method for delivering various molecules to the peri-implant site after implant placement in host tissue. By definition, a mesoporous material is characterized by an interconnected framework of pores with diameters in the range of 2–50 nm [116]. Self-assembly of a structure-directing agent and an inorganic precursor produces an ordered cubic porous structure. By changing the process parameters and the volume ratio of the self-assembly components, it is possible to generate mesoporous structures of various pore dimensions and distributions. The amount of substance loaded into the mesoporous matrices varies according to pore dimension and surface area [48]. In one ex vivo study, Alendronate and Raloxifene were selected to assess the capability of mesoporous films deposited on Ti implants to act as drug carriers at the bone- implant interface. A slow and sustained release of these drugs was confirmed with the QCM-D analysis [117]. In addition, significant increases in bone density were demonstrated to occur around implants coated with mesoporous films and loaded with the two drugs compared to non-loaded mesoporous implants [49].

The present thesis selects Mg as a promising bioactive element for incorporation into a mesoporous matrix and delivery to the bone- implant interface. The first three studies explored the biochemically active nature of Mg, especially in bone metabolism, with particular focus on the effect of its release in various cell models at their short and long developmental stages [118, 119]. The primary objective was to examine pore size and distribution with microscopic techniques, such as the SEM (study I) and, as previously shown by Karlsson et al., with transmission electron microscopy [120]. An average pore diameter of 6.0 nm was measured, with pores that faced out from the surface, permitting Mg to be adsorbed both into the pores and onto the surface. Pore distribution was narrow and

80 80 ordered, which is fundamental for controlling the adsorption and release of substances [120].

Chemical analysis is a first step in characterizing a biomaterial surface, since initial interaction with cells and proteins occurs at the bone-implant interface. Several in vivo investigations on the bone- forming capacity of Mg-loaded implants have defined this interaction as “biomechanical bonding” between bone and the Mg oxide layer on the Ti surface [101, 112, 121]. Studies I and III used XPS in surface chemical analyses to confirm the successful loading of Mg in mesoporous films. This adsorption phenomenon was due to the EISA parameters and reagent concentrations that were selected for forming the coating, which resulted in an interconnected porous cubic structure. Figure 20A illustrates the cubic structure of mesoporous films.

Study I detected an average Mg percentage of 8.6. This result agreed with the findings of Sul et al., who investigated the biomechanical bonding around implants loaded with various atomic concentrations of Mg that were placed in rabbit tibia; the researchers concluded that the bone response was Mg-dose dependent, with a Mg concentration of about 9% being most favourable [122]. More recently, a similar percentage of Mg was loaded in Mg-mesoporous Ti implants made according to the same procedure described in this thesis. It was demonstrated that the amount of Mg released from such implants significantly improved implant retention in the early stages of healing in rabbit tibiae, enhancing bone-implant contact and gene expression dynamics at the peri-implant site [123, 124].

81 81

Figure 20. Schematic illustration of the formation of mesoporous titania thin films. EISA = evaporation-induced self-assembly [125].

Modification of surface chemistry often alters surface topography, particularly at the nano level [126]. Thus, discriminating between positive bioactive effects resulting from surface chemistry modification and effects related to nanotopography alteration is currently difficult, even under controlled laboratory conditions. For this reason, a systematic investigation of the surface micro- and nanotopography of the mesoporous films before and after Mg impregnation was done in all studies to reliably ascertain surface properties and their impact on cellular development. The interferometry investigation confirmed a smooth mesoporous surface microtopography (Sa about 0.3 µm), which did not change after Mg loading. Both interferometry and AFM results displayed a larger developed surface area parameter (Sdr) due to the addition of Mg. Possibly, this can be attributed to Mg oxides that adsorbed to the surface and were rapidly dissolved when placed in the culture medium, producing a transitory topographical effect at the nano level.

After confirmation of Mg loading on the mesoporous film, all samples were incubated in cell culture medium in conditions of

82 82 constant shaking that closely mimic the physiological dynamic environment (Study I). Maximum Mg release was observed within 24 hours of incubation, thus positive effects of Mg were expected after 24 hours of cell growth on the Mg-loaded surface. The rapid release of Mg from the mesoporous surface may be related to the nature of the Mg–TiO2 layer bond, which probably occurred through physical deposition of Mg(OH)2 onto the Ti surface.

Study I investigated the mesoporous surface with and without Mg loading with respect to hFOB. Uncoated Ti surfaces were included in each investigation as negative surface controls, and cells cultured on plastic wells were included as positive cell controls. Cell response was tested after short and long culture periods under Mg- free conditions in order to quantify the trace amounts of Mg released from the mesoporous matrix. Initial hFOB viability (24- hour cell growth) was significantly affected by a Mg-enriched surface compared to the positive cell control. Nevertheless, Mg potential was only detected in 24-hour cell populations while no differences were observed in presence of Mg in the longer-term cell experiments. However, native TiO2 mesoporous coatings significantly promoted mRNA expression of RUNX2, ITGA1, OCN, and ALP compared with uncoated samples.

To better understand the mechanisms behind the mesoporous capacity of the titania coating to adsorb and release Mg, and the osteogenic potential of Mg, the dynamics of gene expression were investigated in ADSC, which are unspecialized cells able to differentiate into various cellular phenotypes, including osteoblasts (Study II). The same experimental groups in Study I were used in Study II, to highlight the cell type and its ability to administrate Mg in various cellular stages. A complete cell culture medium with standard ion and saline concentrations was used to mimic the physiological state. The emphasis on Mg-release functionality was to evaluate if Mg was able to promote osteogenic differentiation in undifferentiated cells alone, without additional osteogenic promo- ters in the cell culture system. Real-time PCR showed that the effect of Mg seems to be more prominent than previous outcomes with hFOB, even in later stages of ADSC development. Real-time

83 83 PCR data sets clearly demonstrated (i) the osteoconductive potential of interconnected nano-porosity, which appeared to be able to adsorb growth and osteogenic factors (protocol 1, Materials and Methods), and (ii) the beneficial effect of Mg when cells were cultured on Mg-loaded surfaces (protocol 2, Materials and Methods). The two experimental set-ups allowed investigation of the osteogenic properties of the ordered nano-pores as well as the capability of these cells to administrate Mg in the expression of osteogenic markers. It was demonstrated that OPN, OCN, and ONC reached their highest levels in Mg condition at 2 weeks and were as well expressed as in the osteogenic control where, it was assumed, the osteogenic medium allows each gene to express 100%. OPN in particular exhibited a strong tendency of upregulation at 2 weeks in when cells were cultured in the presence of Mg; this tendency became significant at 3 weeks in ADSC cultured on Mg-loaded surfaces. This is an intriguing outcome since OPN is an abundant protein in the cement lines between new and old bone [127]. Moreover, experimental evidence demonstrated that OPN mediates integrin pathways [128]. Integrins are those proteins involved in cell–cell and cell–substrate adhesion; thus it can be speculated that the high expression of OPN in ADSC cultured in Mg-enriched conditions may be one indication of the potential contribution of Mg in osteogenic differentiation [129, 130]. COL I remained down-regulated in both test conditions compared to the osteogenic control at 2 weeks but at 3 weeks it reached a level comparable to in the control. This agrees with our understanding that mature osteoblasts normally express collagen before mineralization; osteoprogenitors and pre-osteoblasts do not [131, 132].

In summary, Study I and II outcomes show that Mg was easily incorporated into the mesoporous titania matrix and, subsequently, released at first contact with the cellular environment. The in vitro release test demonstrated a maximum Mg release during the first 24 hours of incubation; thus, positive effect of the Mg was only observed on osteoblast viability during the first day of growth on the specimens. Nevertheless, the sustained release of Mg guided undifferentiated cells toward the osteoblastic phenotype, as the gene expression analyses in Study II

84 84 showed. OPN, OCN, and ONC are fundamental proteins in osteogenic activity, so their expression in the presence of Mg- loaded MP surface (MPMg) is a strong indication of the osteogenic differentiation occurring in adipose stromal cells.

Mg content and release modality, however, are two parameters that most likely could be further explored and optimized to improve the osteogenic potential of Mg.

The pore size in a MP matrix can be fine-tuned by varying the self- assembly parameters and the structure-directing agent nature and concentration [48]. Tuning the pore dimensions of titania films on implants allows the total loading capacity of adsorbing bioactive molecules and substances to be optimized. Study III varied the chemicals and their concentrations to generate average pore sizes of 2- and 7- nm in MP films, allowing their Mg loading and releasing properties to be compared with the previously investigated 6-nm pore size MP films.

The morphological analysis of the three native TiO2 films showed the same narrow, ordered distribution of pores, regardless of pore size. For the surface chemical analysis, XPS analysed Mg content on the surface, while QCM-D, a more accurate analytical technique, analysed Mg content in the pores, monitoring the dynamic flow of Mg into the MP films and its subsequent release after pore saturation. XPS results showed that Mg content was highest on 6-nm MP surfaces. This is inconsistent with the QCM-D analysis, which, conversely, demonstrated that the highest Mg adsorption occurred in the pores of the 7-nm MP films, followed by the 2-nm films, with the 6-nm titania films exhibiting the lowest Mg adsorbing capacity. The topographical parameters of the 6-nm titania films might explain these discrepancies between the XPS and QCM-D findings; AFM, used to characterize the surface topography, showed that the native surface of the 6-nm films had a greater surface area (Sdr) than the 7-nm films, and the highest number of summits per unit of area (Sds). Thus, the 6-nm film surface was highly exposed to the MgCl2 soaking solution, than either of the other films, which increased the amount of Mg able to

85 85 physically adsorb to the film. Additionally, differences in the Mg- loading procedures for the two analyses might be another reason for the inconsistencies in the two data sets: a dynamic flow of

MgCl2 to load the films was applied in the QCM-D procedure, while a static soaking procedure in MgCl2 preceded XPS analysis. Possibly, Mg dried on the 6-nm surface in oxide form, which immediately dissolved in contact with fluids. This suggests that the Mg on the 6-nm MP films is released in an initial burst, which would be more effective in stimulating cells in the early stages of growth, as Studies I and II showed [118].

It is well known that the osseointegrative properties of Ti implants are highly dependent on both surface nano-features and surface chemistry, which together, have a synergic effect and enhance osteogenesis around implants [11, 24, 25, 133, 134]. The absence of significant differences in average height among the native MP films and in comparison with their Mg-loaded counterparts allowed to correlate the variation in hFOB mRNA levels only to the Mg released from the three mesoporous films. In this context, a marked higher variability in the expression of the 84 osteogenic genes was observed in the comparison of 7 nm mesoporous matrix and its Mg-impregnated counterpart (MP3 vs MP3Mg) after one week of hFOB growth. It is worthwhile pointing out that OCN was highly up-regulated solely in osteoblasts cultured on MP3Mg and not in the other two Mg-loaded titania films. This is important, because this non-collagenous protein is selectively expressed by mature osteoblasts; its expression is thus an indication of properly developing hFOB [135, 136].

Other relevant proteins were observed, such as COL14AI, which underwent a nearly four times higher fold-change in cells cultured on Mg-impregnated MP3 films compared with in cells cultured on the native titania matrix. Two metalloproteinases, MMP2 and MMP9, were also observed; these are essential in bone turnover with the specific role of degrading extracellular matrix components, thus permitting cell migration and deposition of new extracellular structures [137, 138].

86 86 Considerable attention in bone morphogenesis investigations is addressed to bone morphogenic proteins (BMPs). These are multi- functional growth factors that belong to the transforming growth factor beta (TGFB) superfamily. Once secreted, these proteins bind to their receptors and induce the mesenchymal condensation pre- figuring osteoblast phenotype and, thus, osteoblast differentiation [139-141]. BMP2 and BMP4 were observed in MP1Mg and MP2Mg, respectively, at 1 week of osteoblast proliferation. The great variability in osteogenic gene expression observed in cells growing on MP3Mg was visibly reduced at 3 weeks of cell proliferation. However, the real-time PCR results showed that BMP4 was significantly up-regulated in osteoblasts grown on MP3Mg at 3 weeks of culture, while the same gene was significantly downregulated when grown on MP1Mg. These biological differences between the two MP films that differed slightly in Mg loading capacity may be due to the ability of the 7-nm MP film to more slowly administrate Mg, which would also have a positive effect on cell bioactivity in later stages of tissue formation.

Magnesium alloys as tailored biodegradable implant materials for bone regeneration – study IV Ti is a commonly used metal for the production of permanent endosseous implants [28, 142, 143]. However, many permanent devices must be removed due to rejection by the host, failure, pain, or discomfort [144]. Conversely, polymers resorb when tissue heals, but they are not biomechanically comparable to metals and, therefore, are not considered an optimal choice for weight-bearing bone applications [145].

Mg alloys are receiving increasing attention in medical research as implantable materials for skeletal conditions [146, 147]. Results to date demonstrate that these materials degrade safely in living tissue with the advantages of higher ultimate strength than polymers and an elastic modulus close to that of cortical bone [59, 89]. Hence, they represent valuable alternative materials for regenerative skeletal applications. Many studies have shown that biodegradable Mg implants enhance bone formation [89, 148, 149] and osteoblast activity [107, 150] during degradation due to the bioactive

87 87 nature of Mg. The most problematic issue in the use of Mg-based materials, however, is controlling their degradation behaviour in an aqueous environment, such as occurs in the human body. Mg degradation is accompanied by hydrogen gas evolution and chemical surface alteration, which might further irritate the injured site and do not properly allow to match the bone healing in vivo [88, 89, 148, 151]. To tailor the biodegradable properties of such materials, various alloying elements have been combined with bulk Mg, both physiological and non-physiological elements [152]. However, many drawbacks arose due to possible toxic effects of some non-physiological elements.

In Study IV of the present thesis, three Mg alloys were selected and their degradation properties analysed under standard cell culture conditions, and compared with high-purity Mg. The influence of the materials’ degradation on the early response of HUCPV stem cells in terms of cell viability and adhesion structures was observed to provide a preliminary indication of the suitability of these alloys as biodegradable metals that can be further proceeded in vivo. As previously described, the influence of surface chemistry and topography is highly important in the development and testing of new biomaterials, regardless of their biodegradable or non-degradable nature. When a foreign material is implanted in living tissue, it is important that the initial interaction between the implant surface and the surrounding biological entities, (e.g., platelets, proteins and cells) is correct and drives the subsequent cellular and molecular events that induce adherent cells to build newly formed tissue. [58]. Thus, in the case of Mg alloys, changes in the pH and osmolality of the environment surrounding the implant, as well as a too fast/slow degradation rate are critical factors for material cyto/tissue- compatibility [59, 60, 62, 88, 145, 153].

The actin cytoskeleton is a highly dynamic network composed of actin polymers and other associated proteins. The cytoskeleton mediates a variety of essential biological functions in all eukaryotic cells, including intra- and extra-cellular movement and structural support. To perform these functions, the organization of the actin cytoskeleton must be tightly regulated, both temporally and

88 88 spatially. Thus, the orientational distribution of actin filaments is an important determinant of cellular shape and motility on any kind of substrate. Focal adhesion and adherent junctions are membrane-associated complexes that serve as nucleation sites for actin filaments and as cross-linkers between the cells, plasma membrane, and actin cytoskeleton. The present investigations demonstrated that HUCPV differently adhered and spread onto Mg alloys and high-purity Mg. As degradation increased, cell density on Mg4Y3RE was reduced. The higher degradation rate of Mg4Y3RE might explain the low cell density compared to Mg2Ag and Mg10Gd. Accordingly, the material degradation analysis for Mg4Y3RE showed that pH and osmolality of the medium were higher after 72 hours of immersion than at baseline. It has been demonstrated that hypertonic medium (a medium with higher extra- than intracellular solute concentration) is a critical condition for cells and depresses cell growth [154-156]. Hypertonic conditions cause water to diffuse out of the cells due to osmotic pressure. Thus, cell shrinkage might be one cause of cell reduction on the Mg4Y3RE surface.

Changes in pH may also affect the initial phase of cell proliferation [157-159]. In their first contact with Mg-based materials, cells are exposed to rapid increases in pH [88, 146, 153, 160, 161]. Focal adhesions, for example, are more stable in an environment at pH values below 6.0 and above 7.2 [162, 163]. The alkalizing effect of material degradation raises the extracellular pH, which becomes acidic during wound healing and skeletal regeneration [164]. This buffering effect might preserve an alkaline environment, which is favourable for the development of cellular adhesion structures. Mg4Y3RE induced a too strong increase in pH in the incubation medium, which probably reduced cell density, but allows a well development of actin structures.

Actin structures were poorly developed and sometimes invisible on Mg2Ag compared with the other alloys. Conversely, cytoskeletal structures were well-defined and -developed on Mg10Gd.

89 89 AFM measurements of materials surfaces at each degraded state showed that the surface texture of Mg2Ag samples became smoother between 24 and 48 hours of degradation, with significant reductions in surface area (Sdr) as pre-incubation time increased. After 48 hours of immersion, a significant increase in average height deviation (Sa) occurred. It is known that the degree of cellular spreading on a substrate also depends on the surface area of the substrate that is exposed to the surrounding [165]. Thus, the strong change in the surface topography of the Mg2Ag samples, with the initial reduction in specific surface area, may be another factor that contributed to weak cellular adhesion. Conversely, surface degradation was more homogeneous on the Mg4Y3RE and Mg10Gd samples, which suggests that these alloys are somewhat less susceptible to surface pitting degradation, thus possibly permitting correct initial cell recruitment and adhesion in living tissue.

In conclusion, these observations indicate that Mg10Gd may better induce filopodia and lamellopodia development, and thus a more stable cell layer, compared with Mg2Ag that reduced cell activity, and with Mg4Y3RE that possessed the highest degradation rate.

90 90 CONCLUSIONS AND FUTURE PERSPECTIVES

Mesoporous titania film as a carrier for magnesium at the peri-implant site From Studies I, II and III it can be concluded that

. Magnesium can be successfully loaded into mesoporous films with an average pore size of 6 nm; within 1 hour of incubation in cell conditions, magnesium is released in an initial burst.

. Magnesium loading does not alter the surface morphology or microtopography of the mesoporous film, but loading does provoke a transitory topographical effect at the nano level, possibly due to the rapid dissolution of magnesium oxides when in contact with the culture medium.

. The osteoconductive potential of magnesium is detectable after 24 hours of hFOB growth, most likely due to the rapid release of magnesium. However, no magnesium- positive effect is observable in late stages of osteoblast proliferation compared to non-loaded mesoporous films.

. Magnesium enhances the expression of osteoblast phenoytpic markers (OPN, OCN, ONC, RUNX2) when osteoprogenitors (ADSC) are grown in contact with Mg- enriched mesoporous films.

. Increasing the pore dimensions of mesoporous films (from 6-nm to 7-nm) by adding a swelling agent such as poly(propylene) glycol (PPG) in the precursor solution

91 91 allows to improve the control of magnesium release, giving a beneficial osteogenic effect on adherent osteoblasts even at later stages of their osteogenic activity.

Despite the high clinical success rate of modern dental implants, there still remain possibilities for improvement in order to expand the application spectra. In the future there may be an implant surface modification that could significantly reduce the risk of initial implant failure, and magnesium as surface modification agent could represent one of the alternatives.

The findings reported in this thesis suggest that an engineered mesoporous titania matrix produced in EISA process and physically loaded with Mg functions as drug (Mg in this thesis) releasing system for titanium implants. The intriguing characteristic of this engineered surface relies on its ability to adsorb Mg without altering the surface roughness, thus producing an initial transitory enhancement of progenitor cells osteogenic differentiation, osteoblasts proliferation and bone apposition to the surface given by the bioactivity of this element. Moreover, the nano-scale dimensions of these mesoporous films are important, as witnessed by the significant increase in Mg osteogenic potential at the cellular level when pore size increased by 1 nm.

Future research topics may include further increases in nano-pore dimensions of mesoporous films, maintaining the native surface microstructures, with the aim to optimize Mg content and, thereby, its osteogenic potential in more complex in vitro systems as well as in higher organisms.

Magnesium alloys as biodegradable metals for bone tissue regeneration Study IV showed how degradation parameters, surface topography, and surface chemistry of Mg-based materials affect initial cellular adhesion. It can be concluded that:

92 92 . No adverse cell reactions were observed on any Mg alloy at any measured point of degradation (24, 48 and 72 h). Of the four tested materials (three Mg alloys and high- purity Mg), Mg4Y3RE showed the highest degradation rate, and the Mg4Y3RE medium exhibited the highest increases in pH and osmolality as degradation proceeded. HUCPV cell density was likely reduced when grown on this alloy. However, cellular adhesion structures appeared to be better developed and more evenly distributed in each direction over Mg4Y3RE and Mg10Gd than high-purity Mg. Additionally, Mg4Y3RE and Mg10Gd had a more homogeneous surface degradation profile compared to Mg2Ag on which cellular actin structures were less developed and less elongated.

When the resorption of the alveolar ridge after tooth extraction occurs, reconstruction or augmentation of the bone defect is often applied. Non-resorbable devices such as membranes made of titanium are used in some cases. The main disadvantage is that these materials must be removed at re-entry. In contrast, biodegradable synthetic or natural polymers do not require removal due to their degradable nature, however their mechanical stability has been of some discussion.

Mg alloys represent a possible alternative in oral, maxillofacial, and orthopaedic surgery due to their biodegradable but metallic (mechanically stable) nature. The main objective is to develop a prototype Mg- based implant with optimal and safely tailored biodegradable properties. To do this, various alloying elements are still being explored for their potential to modulate the degradation behaviour of Mg. The aim is to reduce the content of the alloyants while optimizing the degradation rate for cellular adhesion, proliferation and differentiation, thus, mimicking the tissue healing process in higher organisms as closely as possible. The present findings are a preliminary biological indication of the suitability of Mg10Gd as a biomedical material. They show how the degradation parameters, surface topography, and surface chemistry

93 93 of the Mg-based materials affect initial cellular adhesion. Although Mg4Y3RE is a well-tolerated alloy that safely degrades both in vitro and in vivo, its initial degradation is still too fast and localized. Examining cellular and tissue responses to Mg alloys with varying percentages of Gd as a suitable alloying element may be a rewarding path.

94 94 ACKNOWLEDGEMENTS

It has been a unique journey for me, the so-called “starting point”, and it would not have been possible without the precious help from a lot of people. I wish to express my sincere gratitude to:

Associate Professor Ryo Jimbo, my supervisor and friend, for teaching me how to “risk” in research, without having the fear to fail. You made me more self-confident, professional and independent, guiding me and, at the same time, giving me the freedom of choice. Thank you for motivating me, even when I felt lost.

Professor Ann Wennerberg, my co-supervisor, for your inspiration, your valuable knowledge, your incitement and encouragement, your kind understanding and solidarity when I needed it. It has been an honour to be part of your team.

Professor Thomas Albrektsson, for the fruitful discussions and tricky questions even along the corridors of the school! Thank you also for giving me this great opportunity. It has been a pleasure to meet you.

I am deeply grateful to Professor Jan Lindhe, without whom I would never be here. Thank you.

Professor Regine Willumeit-Römer, head of the division Metallic Biomaterial at Helmoltz-Zentrum Geesthacht, my external co- supervisor in the European Marie-Curie network. Thank you for your endless positive energy and enthusiasm in motivating us, for

95 95 the fruitful discussions and for the nice time together during our workshops and meetings around Europe.

Professor Julia Davies and the staff at the Department of Oral Bi- ology for the outstanding discussions and the fruitful collaboration. Thank you also for your help in the practical things during the PhD journey.

My colleagues: Silvia Galli, co-author, travel-mate, room-mate, compatriot and friend, that made my staying far away from home more pleasant and funny!

Yoshihito Naito for your knowledge, kindness and for the awesome time together in Padova and Venice. It has been a pleasure to have you there.

Marco Toia, who deeply supported me in many occasions within and outside work and with whom I understood a lot from a clinical perspective.

Mariko, Yohei, Michele, Ali, Pär, Anna, Ramesh, Ricardo, Bruno for collaboration, timely help, support and awesome party time!

The entire staff at the Faculty of Odontology, Malmö University and the staff of the Department of Prosthodontics, with special thanks to Ulf Persson and Monica Lotfinia for the administrative things, and to Deyar Mahmood and Evaggelia Papia for your use- ful tips during the preparation of the thesis.

The ESR´s, my colleagues within the European Marie-Curie Train- ing Network, with special thanks to Helvia and Nezha, my co- authors, for fruitful brainstorming and technical support and nice time; Gabor for the material production and for making the night shift during the monitoring of the synchrotron radiation tomogra- phy (somehow) funny! Andrea and Inigo for the lovely weekend in Prague; Maria for your warm italian attitude; Marian, Anastasia, Sepideh, Sriveena, Maryam, Frantisek and Olga.

96 96 Post-Doc Frank Feyerabend and Berengere Julie Christine Lu- thringe, co-authors and valuable researchers at Helmoltz-Zentrum. You have inspired me a lot.

The “Chalmers Crew” in Göteborg University, Professor Martin Andersson and his staff Johan, Wenxiao and Saba at the Depart- ment. It has been a pleasure to collaborate with you and I have learnt a lot at you Department.

Professor Kamal Mustafa and the valuable staff of Department of Clinical Dentistry at the University of Bergen in Norway.

Assistant Professor Barbara Zavan and the girls Letizia and Chiara at the Department of Histology, Microbiology, and Medical Bio- technologies at the University of Padova in Italy. Thank you very much for giving me the chance to work in my hometown.

My wonderful friends in Padova, and particularly to Pamela for being here today, and to my best friend Claudia with whom the friendship has never changed for 26 years. Thanks for you for being part of the family.

My family, especially my grandmother nonna Lina who under- stands me as anyone else; my uncle Fabrizio for taking care of me as the brother that I´ve never had.

Finally the deepest thanks to my parents, Denis and Tiziana, for their endless support and love. Thank you for believing in me.

Grazie di cuore a tutti!

This reasearchreasearch work has been mademad possible by funding from the People Programme (Marie Curie Actions) ofof the the European European Union’s Union ́s Seventh Framework Programme FP7/2007-2013/FP7/2007-2013/ under REA grant agreement n°n289163.

95 97 REFERENCES

[1] Dudley J. Implants for the ageing population. Australian dental journal 2015;60 Suppl 1:28-43. [2] Tullo E, Khoo TK, Teodorczuk A. Preparing to meet the needs of an ageing population - a challenge to medical educators globally. Medical teacher 2015;37:105-7. [3] contributors W. Population ageing. Wikipedia, The Free Encyclopedia; 2015. [4] Franks CM. Handbook of Clinical Behavior-Therapy with the Elderly Client - Wisocki,Pa. J Behav Ther Exp Psy 1993;24:269-. [5] Cooper C, Campion G, Melton LJ, 3rd. Hip fractures in the elderly: a world-wide projection. Osteoporosis international : a journal established as result of cooperation between the European Foundation for Osteoporosis and the National Osteoporosis Foundation of the USA 1992;2:285-9. [6] Reginster JY, Burlet N. Osteoporosis: a still increasing prevalence. Bone 2006;38:S4-9. [7] http://bjdonline.org/. [8] Johansson B, Krekmanov L, Thomsson M. Miniplate osteosynthesis of infected mandibular fractures. Journal of cranio-maxillo-facial surgery : official publication of the European Association for Cranio-Maxillo-Facial Surgery 1988;16:22-7. [9] Bos RR, Rozema FR, Boering G, Nijenhuis AJ, Pennings AJ, Verwey AB. Bio-absorbable plates and screws for internal fixation of mandibular fractures. A study in six dogs. International journal of oral and maxillofacial surgery 1989;18:365-9. [10] Hermawan H, Mantovani D. Degradable metallic biomaterials: the concept, current developments and future directions. Biomaterials 2009;21:207-16. [11] Wennerberg A, Albrektsson T. Effects of titanium surface topography on bone integration: a systematic review. Clinical oral implants research 2009;20 Suppl 4:172-84.

98 98 [12] Albrektsson T, Branemark PI, Hansson HA, Lindstrom J. Osseointegrated titanium implants. Requirements for ensuring a long- lasting, direct bone-to-implant anchorage in man. Acta orthopaedica Scandinavica 1981;52:155-70. [13] Wennerberg A, Albrektsson T, Andersson B. Bone tissue response to commercially pure titanium implants blasted with fine and coarse particles of aluminum oxide. The International journal of oral & maxillofacial implants 1996;11:38-45. [14] Wennerberg A, Albrektsson T, Johansson C, Andersson B. Experimental study of turned and grit-blasted screw-shaped implants with special emphasis on effects of blasting material and surface topography. Biomaterials 1996;17:15-22. [15] Simion M, Gionso L, Grossi GB, Briguglio F, Fontana F. Twelve- Year Retrospective Follow-Up of Machined Implants in the Posterior Maxilla: Radiographic and Peri-Implant Outcome. Clin Implant Dent Relat Res 2014. [16] Albouy JP, Abrahamsson I, Persson LG, Berglundh T. Spontaneous progression of peri-implantitis at different types of implants. An experimental study in dogs. I: clinical and radiographic observations. Clinical oral implants research 2008;19:997-1002. [17] Albouy JP, Abrahamsson I, Persson LG, Berglundh T. Spontaneous progression of ligatured induced peri-implantitis at implants with different surface characteristics. An experimental study in dogs II: histological observations. Clinical oral implants research 2009;20:366-71. [18] Quirynen M, Bollen CM. The influence of surface roughness and surface-free energy on supra- and subgingival plaque formation in man. A review of the literature. J Clin Periodontol 1995;22:1-14. [19] Bressan E, Sbricoli L, Guazzo R, Tocco I, Roman M, Vindigni V, et al. Nanostructured surfaces of dental implants. International journal of molecular sciences 2013;14:1918-31. [20] Biggs MJ, Richards RG, Gadegaard N, McMurray RJ, Affrossman S, Wilkinson CD, et al. Interactions with nanoscale topography: adhesion quantification and signal transduction in cells of osteogenic and multipotent lineage. Journal of biomedical materials research Part A 2009;91:195-208. [21] Dalby MJ, Gadegaard N, Tare R, Andar A, Riehle MO, Herzyk P, et al. The control of human mesenchymal cell differentiation using nanoscale symmetry and disorder. Nature Materials 2007;6:997-1003. [22] Dalby MJ, Macintyre A, Roberts JN, Yang J, Lee LC, Tsimbouri PM, et al. Nanoscale titanium surface treatments for marrow progenitor culture. Nanomedicine (Lond) 2012;7:20-1.

99 99 [23] Mendonca G, Mendonca DB, Simoes LG, Araujo AL, Leite ER, Duarte WR, et al. The effects of implant surface nanoscale features on osteoblast-specific gene expression. Biomaterials 2009;30:4053-62. [24] Wennerberg A, Albrektsson T. Structural influence from calcium phosphate coatings and its possible effect on enhanced bone integration. Acta Odontol Scand 2009;67:333-40. [25] Jimbo R, Sotres J, Johansson C, Breding K, Currie F, Wennerberg A. The biological response to three different nanostructures applied on smooth implant surfaces. Clinical oral implants research 2012;23:706-12. [26] Cohen J. Current concepts review. Corrosion of metal orthopaedic implants. The Journal of bone and joint surgery American volume 1998;80:1554. [27] Nagels J, Stokdijk M, Rozing PM. Stress shielding and bone resorption in shoulder arthroplasty. Journal of shoulder and elbow surgery / American Shoulder and Elbow Surgeons [et al] 2003;12:35-9. [28] Proceedings of the 1st Symposium on Degradable Metals for Biomedical Applications, part of the International Conference on Processing & Manufacturing of Advanced Materials (THERMEC). August 25-29, 2009. Berlin, Germany. Acta biomaterialia 2010;6:1679- 877. [29] Champy M, Lodde JP, Schmitt R, Jaeger JH, Muster D. Mandibular osteosynthesis by miniature screwed plates via a buccal approach. Journal of maxillofacial surgery 1978;6:14-21. [30] Dimitriou R, Jones E, McGonagle D, Giannoudis PV. Bone regeneration: current concepts and future directions. BMC medicine 2011;9:66. [31] Dahlin C, Linde A, Gottlow J, Nyman S. Healing of bone defects by guided tissue regeneration. Plastic and reconstructive surgery 1988;81:672-6. [32] Piette E, Alberius P, Samman N, Linde A. Experience with e-PTFE membrane application to bone grafting of cleft maxilla. International journal of oral and maxillofacial surgery 1995;24:327-32. [33] Van Steenberghe D, Johansson C, Quirynen M, Molly L, Albrektsson T, Naert I. Bone augmentation by means of a stiff occlusive titanium barrier. Clinical oral implants research 2003;14:63-71. [34] Pinho MN, Roriz VL, Novaes AB, Jr., Taba M, Jr., Grisi MF, de Souza SL, et al. Titanium membranes in prevention of alveolar collapse after tooth extraction. Implant dentistry 2006;15:53-61.

100 100 [35] Otto K, Schopper C, Ewers R, Lambrecht JT. [Guided bone regeneration beneath titanium foils]. Schweizer Monatsschrift fur Zahnmedizin = Revue mensuelle suisse d'odonto-stomatologie = Rivista mensile svizzera di odontologia e stomatologia / SSO 2004;114:996-1002. [36] Sanz M, Lorenzo R, Aranda JJ, Martin C, Orsini M. Clinical evaluation of a new collagen matrix (Mucograft prototype) to enhance the width of keratinized tissue in patients with fixed prosthetic restorations: a randomized prospective clinical trial. J Clin Periodontol 2009;36:868-76. [37] Schwarz F, Bieling K, Latz T, Nuesry E, Becker J. Healing of intrabony peri-implantitis defects following application of a nanocrystalline hydroxyapatite (Ostim) or a bovine-derived xenograft (Bio-Oss) in combination with a collagen membrane (Bio-Gide). A case series. J Clin Periodontol 2006;33:491-9. [38] Wallace SS, Froum SJ, Cho SC, Elian N, Monteiro D, Kim BS, et al. Sinus augmentation utilizing anorganic bovine bone (Bio-Oss) with absorbable and nonabsorbable membranes placed over the lateral window: histomorphometric and clinical analyses. The International journal of periodontics & restorative dentistry 2005;25:551-9. [39] Botticelli D, Berglundh T, Lindhe J. The influence of a biomaterial on the closure of a marginal hard tissue defect adjacent to implants. An experimental study in the dog. Clinical oral implants research 2004;15:285-92. [40] Wennerberg A, Albrektsson T. Suggested guidelines for the topographic evaluation of implant surfaces. The International journal of oral & maxillofacial implants 2000;15:331-44. [41] Ellingsen JE, Johansson CB, Wennerberg A, Holmen A. Improved retention and bone-to-implant contact with fluoride-modified titanium implants. Int J Oral Max Impl 2004;19:659-66. [42] Fandridis J, Papadopoulos T. Surface characterization of three titanium dental implants. Implant dentistry 2008;17:91-9. [43] Kang BS, Sul YT, Oh SJ, Lee HJ, Albrektsson T. XPS, AES and SEM analysis of recent dental implants. Acta biomaterialia 2009;5:2222-9. [44] Massaro C, Rotolo P, De Riccardis F, Milella E, Napoli A, Wieland M, et al. Comparative investigation of the surface properties of commercial titanium dental implants. Part I: chemical composition. Journal of materials science Materials in medicine 2002;13:535-48. [45] Sul YT, Jonsson J, Yoon GS, Johansson C. Resonance frequency measurements in vivo and related surface properties of magnesium- incorporated, micropatterned and magnesium-incorporated TiUnite, Osseotite, SLA and TiOblast implants. Clinical oral implants research 2009;20:1146-55.

101 101 [46] Mendes VC, Moineddin R, Davies JE. The effect of discrete calcium phosphate nanocrystals on bone-bonding to titanium surfaces. Biomaterials 2007;28:4748-55. [47] Hamlet S, Ivanovski S. Inflammatory cytokine response to titanium chemical composition and nanoscale calcium phosphate surface modification. Acta biomaterialia 2011;7:2345-53. [48] Grosso D, Cagnol F, Soler-Illia GJdAA, Crepaldi EL, Amenitsch H, Brunet-Bruneau A, et al. Fundamentals of Mesostructuring Through Evaporation-Induced Self-Assembly. Adv Funct Mater 2004;14:309-22. [49] Harmankaya N, Karlsson J, Palmquist A, Halvarsson M, Igawa K, Andersson M, et al. Raloxifene and alendronate containing thin mesoporous titanium oxide films improve implant fixation to bone. Acta biomaterialia 2013;9:7064-73. [50] Vallet-Regi M. Ordered mesoporous materials in the context of drug delivery systems and bone tissue engineering. Chemistry 2006;12:5934- 43. [51] Berglundh T, Abrahamsson I, Lang NP, Lindhe J. De novo alveolar bone formation adjacent to endosseous implants. Clinical oral implants research 2003;14:251-62. [52] Franchi M, Fini M, Martini D, Orsini E, Leonardi L, Ruggeri A, et al. Biological fixation of endosseous implants. Micron 2005;36:665-71. [53] Webster TJ, Ergun C, Doremus RH, Siegel RW, Bizios R. Specific proteins mediate enhanced osteoblast adhesion on nanophase ceramics. Journal of biomedical materials research 2000;51:475-83. [54] Osborn JS, Newesly H. Bonding osteogenesis induced by calcium phosphate ceramic implants. Biomaterials 1980. [55] Davies JE. Understanding peri-implant endosseous healing. Journal of dental education 2003;67:932-49. [56] Pittenger MF, Mackay AM, Beck SC, Jaiswal RK, Douglas R, Mosca JD, et al. Multilineage potential of adult human mesenchymal stem cells. Science 1999;284:143-7. [57] Wagner W, Wein F, Seckinger A, Frankhauser M, Wirkner U, Krause U, et al. Comparative characteristics of mesenchymal stem cells from human bone marrow, adipose tissue, and umbilical cord blood. Experimental hematology 2005;33:1402-16. [58] Anselme K. Osteoblast adhesion on biomaterials. Biomaterials 2000;21:667-81. [59] Castellani C, Lindtner RA, Hausbrandt P, Tschegg E, Stanzl-Tschegg SE, Zanoni G, et al. Bone-implant interface strength and osseointegration: Biodegradable magnesium alloy versus standard titanium control. Acta biomaterialia 2011;7:432-40.

102 102 [60] Le Guehennec L, Soueidan A, Layrolle P, Amouriq Y. Surface treatments of titanium dental implants for rapid osseointegration. Dental materials : official publication of the Academy of Dental Materials 2007;23:844-54. [61] Albrektsson T, Johansson C. Osteoinduction, osteoconduction and osseointegration. European spine journal : official publication of the European Spine Society, the European Spinal Deformity Society, and the European Section of the Cervical Spine Research Society 2001;10 Suppl 2:S96-101. [62] Thalji G, Gretzer C, Cooper LF. Comparative molecular assessment of early osseointegration in implant-adherent cells. Bone 2012. [63] Lavenus S, Louarn G, Layrolle P. Nanotechnology and dental implants. International journal of biomaterials 2010;2010:915327. [64] Balaban NQ, Schwarz US, Riveline D, Goichberg P, Tzur G, Sabanay I, et al. Force and focal adhesion assembly: a close relationship studied using elastic micropatterned substrates. Nat Cell Biol 2001;3:466-72. [65] Stein GS, Lian JB, van Wijnen AJ, Stein JL, Montecino M, Javed A, et al. Runx2 control of organization, assembly and activity of the regulatory machinery for skeletal gene expression. Oncogene 2004;23:4315-29. [66] Hu Y, Cai K, Luo Z, Zhang Y, Li L, Lai M, et al. Regulation of the differentiation of mesenchymal stem cells in vitro and osteogenesis in vivo by microenvironmental modification of titanium alloy surfaces. Biomaterials 2012;33:3515-28. [67] Stein GS, Lian JB. Molecular mechanisms mediating proliferation/differentiation interrelationships during progressive development of the osteoblast phenotype. Endocrine reviews 1993;14:424-42. [68] Carmeli E, Moas M, Reznick AZ, Coleman R. Matrix metalloproteinases and skeletal muscle: a brief review. Muscle & nerve 2004;29:191-7. [69] Lombard C, Saulnier J, Wallach J. Assays of matrix metalloproteinases (MMPs) activities: a review. Biochimie 2005;87:265- 72. [70] Zhang Y, Khan D, Delling J, Tobiasch E. Mechanisms underlying the osteo- and adipo-differentiation of human mesenchymal stem cells. TheScientificWorldJournal 2012;2012:793823. [71] Gromova OA, Torshin I, Kalacheva AG, Kuramshina DB. [Molecular-biological basics of neuroprotection effects of magnesium]. Zhurnal nevrologii i psikhiatrii imeni SS Korsakova / Ministerstvo zdravookhraneniia i meditsinskoi promyshlennosti Rossiiskoi Federatsii, Vserossiiskoe obshchestvo nevrologov [i] Vserossiiskoe obshchestvo psikhiat 2011;111:90-101.

103 103 [72] Iseri LT, French JH. Magnesium: nature's physiologic calcium blocker. American heart journal 1984;108:188-93. [73] Yang L, Arora K, Beard WA, Wilson SH, Schlick T. Critical role of magnesium ions in DNA polymerase beta's closing and active site assembly. Journal of the American Chemical Society 2004;126:8441-53. [74] Wallach S. Effects of magnesium on skeletal metabolism. Magnesium and trace elements 1990;9:1-14. [75] Jones G. Early life nutrition and bone development in children. Nestle Nutrition workshop series Paediatric programme 2011;68:227-33; discussion 33-6. [76] Rude RK, Kirchen ME, Gruber HE, Meyer MH, Luck JS, Crawford DL. Magnesium deficiency-induced osteoporosis in the rat: uncoupling of bone formation and bone resorption. Magnesium research : official organ of the International Society for the Development of Research on Magnesium 1999;12:257-67. [77] Rude RK, Singer FR, Gruber HE. Skeletal and hormonal effects of magnesium deficiency. Journal of the American College of Nutrition 2009;28:131-41. [78] Sojka JE, Weaver CM. Magnesium supplementation and osteoporosis. Nutrition reviews 1995;53:71-4. [79] Clark I, Belanger L. The effects of alterations in dietary magnesium on calcium, phosphate and skeletal metabolism. Calcified tissue research 1967;1:204-18. [80] Schwartz R, Reddi AH. Influence of magnesium depletion on matrix- induced endochondral bone formation. Calcified tissue international 1979;29:15-20. [81] Rude RK, Adams JS, Ryzen E, Endres DB, Niimi H, Horst RL, et al. Low serum concentrations of 1,25-dihydroxyvitamin D in human magnesium deficiency. The Journal of clinical endocrinology and metabolism 1985;61:933-40. [82] Sahota O, Mundey MK, San P, Godber IM, Hosking DJ. Vitamin D insufficiency and the blunted PTH response in established osteoporosis: the role of magnesium deficiency. Osteoporosis international : a journal established as result of cooperation between the European Foundation for Osteoporosis and the National Osteoporosis Foundation of the USA 2006;17:1013-21. [83] Claes LE. Mechanical characterization of biodegradable implants. Clinical materials 1992;10:41-6. [84] Lambotte AL. Utilisation du magnesium comme materiel perdu dans l’osteosynthese. Bulletins et Mémoires de la Société de Chirurgie de Paris 1932;28:1325-59.

104 104 [85] Biomaterials, Part II. Symposium organized by the Scandinavian Orthopedic Association. Ystad, Sweden, September 29-October 1, 1986. Abstracts. Acta orthopaedica Scandinavica 1988;59:342-57. [86] Biomaterials, Part I. Symposium organized by the Scandinavian Orthopedic Association. Ystad, Sweden, September 29-October 1, 1986. Abstracts. Acta orthopaedica Scandinavica 1988;59:197-225. [87] International A. Materials for medical devices. In: Davis JR, editor. USA: ASM International; illustrated edition edition (December 1, 2003); 2003. [88] Willumeit R, Fischer J, Feyerabend F, Hort N, Bismayer U, Heidrich S, et al. Chemical surface alteration of biodegradable magnesium exposed to corrosion media. Acta biomaterialia 2011;7:2704-15. [89] Witte F, Kaese V, Haferkamp H, Switzer E, Meyer-Lindenberg A, Wirth CJ, et al. In vivo corrosion of four magnesium alloys and the associated bone response. Biomaterials 2005;26:3557-63. [90] McBride ED. Absorbable metal in bone surgery - A further report on the use of magnesium alloys. J Amer Med Assoc 1938;111:2464-6. [91] Wren CD. A review of metal accumulation and toxicity in wild mammals. I. Mercury. Environmental research 1986;40:210-44. [92] Nanda A, Saravanan M. Biosynthesis of silver nanoparticles from Staphylococcus aureus and its antimicrobial activity against MRSA and MRSE. Nanomedicine : nanotechnology, biology, and medicine 2009;5:452-6. [93] Tie D, Feyerabend F, Muller WD, Schade R, Liefeith K, Kainer KU, et al. Antibacterial Biodegradable Mg-Ag Alloys. European cells & materials 2013;25:284-98. [94] Huang CF, Chiang HJ, Lan WC, Chou HH, Ou KL, Yu CH. Development of silver-containing austenite antibacterial stainless steels for biomedical applications part I: microstructure characteristics, mechanical properties and antibacterial mechanisms. Biofouling 2011;27:449-57. [95] Mueller WD, de Mele MF, Nascimento ML, Zeddies M. Degradation of magnesium and its alloys: dependence on the composition of the synthetic biological media. Journal of biomedical materials research Part A 2009;90:487-95. [96] Johnson I, Perchy D, Liu H. In vitro evaluation of the surface effects on magnesium-yttrium alloy degradation and mesenchymal stem cell adhesion. Journal of biomedical materials research Part A 2012;100:477- 85.

105 105 [97] Feyerabend F, Fischer J, Holtz J, Witte F, Willumeit R, Drucker H, et al. Evaluation of short-term effects of rare earth and other elements used in magnesium alloys on primary cells and cell lines. Acta biomaterialia 2010;6:1834-42. [98] Witte F, Ulrich H, Rudert M, Willbold E. Biodegradable magnesium scaffolds: Part 1: appropriate inflammatory response. Journal of biomedical materials research Part A 2007;81:748-56. [99] Witte F, Feyerabend F, Maier P, Fischer J, Stormer M, Blawert C, et al. Biodegradable magnesium-hydroxyapatite metal matrix composites. Biomaterials 2007;28:2163-74. [100] Mestres G, Ginebra MP. Novel magnesium phosphate cements with high early strength and antibacterial properties. Acta biomaterialia 2011;7:1853-61. [101] Sul YT, Kang BS, Johansson C, Um HS, Park CJ, Albrektsson T. The roles of surface chemistry and topography in the strength and rate of osseointegration of titanium implants in bone. Journal of biomedical materials research Part A 2009;89:942-50. [102] Yang GL, He FM, Song E, Hu JA, Wang XX, Zhao SF. In vivo comparison of bone formation on titanium implant surfaces coated with biomimetically deposited calcium phosphate or electrochemically deposited hydroxyapatite. The International journal of oral & maxillofacial implants 2010;25:669-80. [103] Albrektsson T, Wennerberg A. The impact of oral implants - past and future, 1966-2042. Journal 2005;71:327. [104] Buser D, Broggini N, Wieland M, Schenk RK, Denzer AJ, Cochran DL, et al. Enhanced bone apposition to a chemically modified SLA titanium surface. J Dent Res 2004;83:529-33. [105] Nishiguchi S, Kato H, Fujita H, Kim HM, Miyaji F, Kokubo T, et al. Enhancement of bone-bonding strengths of titanium alloy implants by alkali and heat treatments. Journal of biomedical materials research 1999;48:689-96. [106] Frojd V, Wennerberg A, Franke Stenport V. Importance of Ca(2+) modifications for osseointegration of smooth and moderately rough anodized titanium implants - a removal torque and histological evaluation in rabbit. Clin Implant Dent Relat Res 2012;14:737-45. [107] Zreiqat H, Howlett CR, Zannettino A, Evans P, Schulze-Tanzil G, Knabe C, et al. Mechanisms of magnesium-stimulated adhesion of osteoblastic cells to commonly used orthopaedic implants. Journal of biomedical materials research 2002;62:175-84.

106 106 [108] Zreiqat H, Valenzuela SM, Nissan BB, Roest R, Knabe C, Radlanski RJ, et al. The effect of surface chemistry modification of titanium alloy on signalling pathways in human osteoblasts. Biomaterials 2005;26:7579-86. [109] Revell PA, Damien E, Zhang XS, Evans P, Howlett CR. The effect of magnesium ions on bone bonding to hydroxyapatite coating on titanium alloy implants. Key Eng Mat 2004;254-2:447-50. [110] Sul YT, Johansson C, Albrektsson T. Which surface properties enhance bone response to implants? Comparison of oxidized magnesium, TiUnite, and osseotite implant surfaces. Int J Prosthodont 2006;19:319- 28. [111] Sul YT, Johansson C, Albrektsson T. A novel in vivo method for quantifying the interfacial biochemical bond strength of bone implants. J R Soc Interface 2010;7:81-90. [112] Sul YT, Johansson P, Chang BS, Byon ES, Jeong Y. Bone tissue responses to Mg-incorporated oxidized implants and machine-turned implantsin the rabbit femur. Journal of applied biomaterials & biomechanics : JABB 2005;3:18-28. [113] Park JW, Kim YJ, Jang JH, Song H. Osteoblast response to magnesium ion-incorporated nanoporous titanium oxide surfaces. Clinical oral implants research 2010;21:1278-87. [114] Sarugaser R, Lickorish D, Baksh D, Hosseini MM, Davies JE. Human umbilical cord perivascular (HUCPV) cells: a source of mesenchymal progenitors. Stem Cells 2005;23:220-9. [115] Trejo CG, Lozano D, Manzano M, Doadrio JC, Salinas AJ, Dapia S, et al. The osteoinductive properties of mesoporous silicate coated with osteostatin in a rabbit femur cavity defect model. Biomaterials 2010;31:8564-73. [116] McCusker LB, Liebau F, Engelhardt G. Nomenclature of structural and compositional characteristics of ordered microporous and mesoporous materials with inorganic hosts (IUPAC recommendations 2001). Micropor Mesopor Mat 2003;58:3-13. [117] Karlsson J, Harmankaya N, Allard S, Palmquist A, Halvarsson M, Tengvall P, et al. Ex vivo alendronate localization at the mesoporous titania implant/bone interface. J Mater Sci-Mater M 2015;26. [118] Cecchinato F, Xue Y, Karlsson J, He W, Wennerberg A, Mustafa K, et al. In vitro evaluation of human fetal osteoblast response to magnesium loaded mesoporous TiO coating. Journal of biomedical materials research Part A 2013.

107 107 [119] Cecchinato F, Karlsson J, Ferroni L, Gardin C, Galli S, Wennerberg A, et al. Osteogenic potential of human adipose-derived stromal cells on 3-dimensional mesoporous TiO2 coating with magnesium impregnation. Materials science & engineering C, Materials for biological applications 2015;52:225-34. [120] Karlsson J, Jimbo R, Fathali HM, Schwartz-Filho HO, Hayashi M, Halvarsson M, et al. In vivo biomechanical stability of osseointegrating mesoporous TiO(2) implants. Acta biomaterialia 2012;8:4438-46. [121] Sul YT, Johansson C, Byon E, Albrektsson T. The bone response of oxidized bioactive and non-bioactive titanium implants. Biomaterials 2005;26:6720-30. [122] Sul YT, Johansson C, Wennerberg A, Cho LR, Chang BS, Albrektsson T. Optimum surface properties of oxidized implants for reinforcement of osseointegration: surface chemistry, oxide thickness, porosity, roughness, and crystal structure. The International journal of oral & maxillofacial implants 2005;20:349-59. [123] Galli S, Naito Y, Karlsson J, He W, Andersson M, Wennerberg A, et al. Osteoconductive potential of mesoporous titania implant surfaces loaded with magnesium: an experimetal study in the rabbit. Clinical Implant Dentistry and Related Research 2013;Epub Ahead of Printing. [124] Galli S, Naito Y, Karlsson J, He W, Miyamoto I, Xue Y, et al. Local release of magnesium from mesoporous TiO2 coatings stimulates the peri- implant expression of osteogenic markers and improves osteoconductivity in vivo. Acta biomaterialia 2014;10:5193-201. [125] Pan JH, Zhao XS, Lee WI. Block copolymer-templated synthesis of highly organized mesoporous TiO2-based films and their photoelectrochemical applications. Chem Eng J 2011;170:363-80. [126] Albrektsson T, Wennerberg A. Oral implant surfaces: Part 1 - Review focusing on topographic and chemical properties of different surfaces and in vivo responses to them. Int J Prosthodont 2004;17:536-43. [127] Sodek J, Brunette DM, Feng J, Heersche JNM, Limeback HF, Melcher AH, et al. Collagen-Synthesis Is a Major Component of Protein- Synthesis in Periodontal-Ligament in Various Species. Archives of oral biology 1977;22:647-53. [128] Katagiri YU, Sleeman J, Fujii H, Herrlich P, Hotta H, Tanaka K, et al. CD44 variants but not CD44s cooperate with beta 1-containing integrins to permit cells to bind to osteopontin independently of arginine- glycine-aspartic acid, thereby stimulating cell motility and chemotaxis. Cancer research 1999;59:219-26. [129] Merry K, Dodds R, Littlewood A, Gowen M. Expression of Osteopontin Messenger-Rna by Osteoclasts and Osteoblasts in Modeling Adult Human Bone. Journal of cell science 1993;104:1013-20.

108 108 [130] Lin Z, Rios HF, Volk SL, Sugai JV, Jin Q, Giannobile WV. Gene expression dynamics during bone healing and osseointegration. Journal of periodontology 2011;82:1007-17. [131] Termine JD, Kleinman HK, Whitson SW, Conn KM, McGarvey ML, Martin GR. Osteonectin, a bone-specific protein linking mineral to collagen. Cell 1981;26:99-105. [132] Zhu B, Lu Q, Yin J, Hu J, Wang Z. Alignment of osteoblast-like cells and cell-produced collagen matrix induced by nanogrooves. Tissue engineering 2005;11:825-34. [133] Wennerberg A, Albrektsson T. On implant surfaces: a review of current knowledge and opinions. The International journal of oral & maxillofacial implants 2010;25:63-74. [134] Jimbo R, Xue Y, Hayashi M, Schwartz-Filho HO, Andersson M, Mustafa K, et al. Genetic responses to nanostructured calcium-phosphate- coated implants. J Dent Res 2011;90:1422-7. [135] Sila-Asna M, Bunyaratvej A, Maeda S, Kitaguchi H, Bunyaratavej N. Osteoblast differentiation and bone formation gene expression in strontium- inducing bone marrow mesenchymal stem cell. The Kobe journal of medical sciences 2007;53:25-35. [136] Patti A, Gennari L, Merlotti D, Dotta F, Nuti R. Endocrine actions of osteocalcin. International journal of endocrinology 2013;2013:846480. [137] Armstrong DG, Jude EB. The role of matrix metalloproteinases in wound healing. Journal of the American Podiatric Medical Association 2002;92:12-8. [138] Zhu XY, Fang CH, Zhang LY, Zuo DM, Chen ZL. [Expressions and role of endogenetic matrix metalloproteinases-9 and transforming growth factor-beta during wound healing of blast injury]. Di 1 jun yi da xue xue bao = Academic journal of the first medical college of PLA 2005;25:844-6. [139] Chen G, Deng C, Li YP. TGF-beta and BMP signaling in osteoblast differentiation and bone formation. International journal of biological sciences 2012;8:272-88. [140] Derynck R, Akhurst RJ. Differentiation plasticity regulated by TGF- beta family proteins in development and disease. Nat Cell Biol 2007;9:1000-4. [141] Tachi K, Takami M, Sato H, Mochizuki A, Zhao B, Miyamoto Y, et al. Enhancement of bone morphogenetic protein-2-induced ectopic bone formation by transforming growth factor-beta1. Tissue engineering Part A 2011;17:597-606.

109 109 [142] De Long WG, Jr., Einhorn TA, Koval K, McKee M, Smith W, Sanders R, et al. Bone grafts and bone graft substitutes in orthopaedic trauma surgery. A critical analysis. The Journal of bone and joint surgery American volume 2007;89:649-58. [143] Niinomi M. Mechanical biocompatibilities of titanium alloys for biomedical applications. J Mech Behav Biomed Mater 2008;1:30-42. [144] Meslemani D, Kellman RM. Recent advances in fixation of the craniomaxillofacial skeleton. Current opinion in otolaryngology & head and neck surgery 2012;20:304-9. [145] Engelberg I, Kohn J. Physicomechanical Properties of Degradable Polymers Used in Medical Applications - a Comparative-Study. Biomaterials 1991;12:292-304. [146] Hort N, Huang Y, Fechner D, Stormer M, Blawert C, Witte F, et al. Magnesium alloys as implant materials--principles of property design for Mg-RE alloys. Acta biomaterialia 2010;6:1714-25. [147] Witte F. The history of biodegradable magnesium implants: a review. Acta biomaterialia 2010;6:1680-92. [148] Witte F, Ulrich H, Palm C, Willbold E. Biodegradable magnesium scaffolds: Part II: peri-implant bone remodeling. Journal of biomedical materials research Part A 2007;81:757-65. [149] Xu L, Yu G, Zhang E, Pan F, Yang K. In vivo corrosion behavior of Mg-Mn-Zn alloy for bone implant application. Journal of biomedical materials research Part A 2007;83:703-11. [150] Castiglioni S, Leidi M, Carpanese E, Maier JA. Extracellular magnesium and in vitro cell differentiation: different behaviour of different cells. Magnesium research : official organ of the International Society for the Development of Research on Magnesium 2013;26:24-31. [151] Witte F, Hort N, Vogt C, Cohen S, Kainer KU, Willumeit R, et al. Degradable biomaterials based on magnesium corrosion. Current Opinion in Solid State & Materials Science 2008;12:63-72. [152] Persaud-Sharma D, McGoron A. Biodegradable Magnesium Alloys: A Review of Material Development and Applications. J Biomim Biomater Tissue Eng 2012;12:25-39. [153] Nguyen TY, Garcia S, Liew CG, Liu H. Effects of magnesium on growth and proliferation of human embryonic stem cells. Conference proceedings : Annual International Conference of the IEEE Engineering in Medicine and Biology Society IEEE Engineering in Medicine and Biology Society Conference 2012;2012:723-6. [154] Dezengotita VM, Kimura R, Miller WM. Effects of CO2 and osmolality on hybridoma cells: growth, metabolism and monoclonal antibody production. Cytotechnology 1998;28:213-27.

110 110 [155] Kajimura M, O'Donnell ME, Curry FE. Effect of cell shrinkage on permeability of cultured bovine aortic endothelia and frog mesenteric capillaries. The Journal of physiology 1997;503 ( Pt 2):413-25. [156] Ozturk SS, Palsson BO. Effect of medium osmolarity on hybridoma growth, metabolism, and antibody production. Biotechnology and bioengineering 1991;37:989-93. [157] Kohn DH, Sarmadi M, Helman JI, Krebsbach PH. Effects of pH on human bone marrow stromal cells in vitro: implications for tissue engineering of bone. Journal of biomedical materials research 2002;60: 292-9. [158] Mackenzie CG, Mackenzie JB, Beck P. The effect of pH on growth, protein synthesis, and lipid-rich particles of cultured mammalian cells. The Journal of biophysical and biochemical cytology 1961;9:141-56. [159] Beaumont KG, Mrksich M. The mechanostability of isolated focal adhesions is strongly dependent on pH. Chem Biol 2012;19:711-20. [160] Janning C, Willbold E, Vogt C, Nellesen J, Meyer-Lindenberg A, Windhagen H, et al. Magnesium hydroxide temporarily enhancing osteoblast activity and decreasing the osteoclast number in peri-implant bone remodelling. Acta biomaterialia 2010;6:1861-8. [161] Seuss F, Seuss S, Turhan MC, Fabry B, Virtanen S. Corrosion of Mg alloy AZ91D in the presence of living cells. J Biomed Mater Res B 2011;99B:276-81. [162] Beaumont KG, Mrksich M. The Mechanostability of Isolated Focal Adhesions Is Strongly Dependent on pH. Chem Biol 2012;19:711-20. [163] Denker SP, Barber DL. Cell migration requires both ion translocation and cytoskeletal anchoring by the Na-H exchanger NHE1. J Cell Biol 2002;159:1087-96. [164] Schneider LA, Korber A, Grabbe S, Dissemond J. Influence of pH on wound-healing: a new perspective for wound-therapy? Archives of dermatological research 2007;298:413-20. [165] Chen CS, Mrksich M, Huang S, Whitesides GM, Ingber DE. Geometric control of cell life and death. Science 1997;276:1425-8.

111 111