surface functionalisation of magnesium alloys for use as bio...

Surface Functionalisation of Magnesium Alloys for Use as Bio-implants

by

Dolly Mushahary

M.Sc. (Biotechnology)

Submitted in fulfilment of the requirements for the degree of

Doctor of Philosophy

Deakin University

July, 2014

parisr

Redacted stamp

……….Dedicated to the Almighty

& my parents

Acknowledgements

i

Acknowledgements

I am forever indebted to my spiritual guru Shri Shri Ravi Shankar for his blessings and

inspiration, without which I would not have been what I am today.

I am thankful to the Director of the Institute for Frontier Materials (IFM), Deakin University,

Prof. Peter Hodgson, for giving me this research opportunity and helping me realize my

dream. This research work was supported by grants from the Department of Industry,

Innovation, Science, Research and Tertiary Education (DIISRTE), Australia, through the

Australia-India Strategic Research Fund (AISRF) and the Department of Biotechnology

(DBT), Govt. of India, Australia India Biotechnology Fund (AIBF) and Australian Research

Council. I thank my supervisors in Deakin University — Dr. Yuncang Li, Prof Cuie Wen,

and Prof. Peter Hodgson, without whom this PhD would not have been possible. I am

thankful to Dr. Gopal Pande, my mentor and co-supervisor in the collaborating Indian

institute, CSIR-Centre for Cellular and Molecular Biology (CSIR-CCMB), who provided the

necessary funds to carry out my research and has helped me grow intellectually. I am also

thankful to Ms. Helen Woodall, partnerships coordinator of Deakin-India, for making me

comfortable during my stay in Deakin.

Many thanks to Jixing Lin for providing the alloys for my research work; my colleague, Ms.

Ragamouni Sravanthi, for her generous contributions in initiating the histology experiments

and her constant scientific feedback; my lab-mates in IFM- Yunfei, Dongmei, Ali and Yu, for

their warmth and support. Yunfei’s assistance in completing the pending experiments will

always be appreciated. I offer special thanks to my senior colleagues, Dr. Tripura

Chaturvedula and Dr. Mahalaxmi Yadla, for constantly supporting and encouraging me

during my stay in CSIR-CCMB — their help and kind words have always lifted my spirit. I

will always cherish the time spent with my previous lab members in CSIR-CCMB —Dr.

Abhilasha Tiwari, Ms. Samiksha Wasnik, Ms. Vibudha Nanduri, Dr. Subhash Sista, Dr.

Smitha Mathews, Dr. Ushasri Chilakamarthi, Ms. Renu Kapoor, Ms. Pavani Gauri and Mr.

Surendra Mohan. Dr. Tiwari — a good critic that she has been throughout my stay in CCMB

and Ms. Wasnik for her patience during my first visit to Deakin in September 2011, will

always be appreciated. I am grateful to the technical staff of my previous lab in CSIR-

CCMB, Mr. Srinivas Gunda and Mr. Avinash Raj, for their feedback and assistance in my

Acknowledgements

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experiments; lab assistant Mr. Srinivas for always providing clean glasswares; the technical

staffs of CSIR-CCMB Fine Biochemicals, Mr. Kishore Joshi and Mr. M.C. Joseph for

providing reagents for my experiments.

This research would not have been complete without the active participation of the staff of

CSIR-CCMB Animal house. Dr. J. Mahesh Kumar performed all the implant surgeries with

great mastery. Mr. B Lalaiah and Mr. Ramappa always helped me in quality monitoring and

took care of my experimental animals with the utmost care and professionalism. Sincere

thanks to Dr. V. Radha for giving me the permission to carry out my cell culture experiments

and the technical officers, Mr. A. Mubarak Ali and Ch. Varalakshmi, for training me with the

basics of cell culture and for their cooperation. I offer my heartfelt gratitude to Ms. Richa

Khanna and Dr. Nitin Tupperwar for their valuable time and helping me with the basics of

RT-PCR, primer design and standardisation. Heartfelt gratitude to Dr. Pankaj Chaturvedi and

Dr. Poonam Sehgal, for helping me, on several occasions, with my work and presentations. I

thank the CSIR-CCMB transport facility for providing the vehicles, many times on a very

short notice, for transport of the experimental rabbits to the National Institute of Nutrition

(NIN), Hyderabad for imaging purpose. Dr. Nemani Harishankar and other staff members of

the Animal house and Imaging facilities of NIN have helped me in DXA imaging and

overnight monitoring of my experimental rabbits. I also extend my gratitude to Dr. Jane

Allardyce, Technical and Academic Editor of IFM, for editing my thesis at a short notice.

How can I not mention my friends from CSIR-CCMB who filled my life with fun and

laughter? Richa Khanna, Satya Brata Rout, Suraj Nongmaithem, Ishan Rastogi, Vidhya

Krishnamoorthy, Swati Singh, Komal Pawar, Dr. Asfarul, Sadeem — life wouldn’t have been

the same without them. I thank my neighbors and my constant companions in RTC Geelong

— Dr. Dipali Devi, Ms. Mayurakshi Bhatta and Ms. Shyama Ranade — for being with me

through the ups and downs. A special thanks to my family and friends in Australia for the

good times I had during my stay — Mr. Nova Basumatary, Mrs. Chandana Basumatary,

Antara aka Lulu Basumatary, Kulfi, Mark Almond, Dr. Sitarama Raju Kada, Shailendra

Pratap Sonkar, Khushboo Rakha, Leena Ditty Sebestian, Shivani Srivastav, Sheena Mathew,

Harshali Patil and Urvashi Bhati.

Acknowledgements

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Somebody who has always stood by me, encouraged me, lifted my spirit and without whom

this whole journey would have been very lonely — Dr. Vikash Agarwal — all I can say is

“Thank you for being there!” And, last but not the least, the most important people in my life

— my parents, my grandmum and the members of my extended family — I will never be

able to thank them enough for all the support and freedom they have given me to reach till

here.

My heartfelt thanks to all those, whose names I might have missed mentioning, who have

been associated with my thesis work and me directly or indirectly!

Dolly Mushahary

25th July 2014

Table of Contents

iv

Table of Contents

Acknowledgements ..................................................................................................................... i

Abbreviations ............................................................................................................................ ix

Chemical Formulae .................................................................................................................... x

Units .......................................................................................................................................... xi

Abstract .................................................................................................................................... xii

List of Publications ................................................................................................................. xiv

Awards ..................................................................................................................................... xv

Presentations ............................................................................................................................ xv

List of Figures ......................................................................................................................... xvi

List of Tables ........................................................................................................................... xx

Chapter 1. Introduction .............................................................................................................. 1

1.1 Hypothesis ........................................................................................................................ 2

1.2 Thesis Objectives ............................................................................................................. 3

1.3 Thesis Outline .................................................................................................................. 3

Chapter 2. Literature Review ..................................................................................................... 6

2.1 Introduction ...................................................................................................................... 6

2.2 Bone Structure and Morphology ...................................................................................... 6

2.3 Bone Composition ............................................................................................................ 8

2.4 Biology of Bone Cells ...................................................................................................... 9

2.4.1 Osteoclasts ................................................................................................................. 9

2.4.2 Osteoblasts ............................................................................................................... 12

2.4.3 Osteocytes ................................................................................................................ 14

2.5 Biomaterials and Biocompatibility- General Concepts .................................................. 17

2.6 Desired Properties of Biomaterials ................................................................................ 18

2.7 Applications of Biomaterials .......................................................................................... 19

2.8 The Necessity of Biomaterials in Bone Tissue Engineering .......................................... 20

2.8.1 Metallic implants in orthopaedics ............................................................................ 22

2.8.2 Mg and Mg alloys as biodegradable implants ......................................................... 25

2.9 Surface Modification of Magnesium Alloys .................................................................. 27

Table of Contents

v

2.9.1 Chemical modification ............................................................................................ 28

2.9.2 Physical modification .............................................................................................. 29

2.10 Summary ...................................................................................................................... 33

Chapter 3. Materials and Experimental Techniques ................................................................ 34

3.1 Sample Preparation ........................................................................................................ 34

3.2 Microstructure, Mechanical and Surface Properties ...................................................... 34

3.2.1 XRD and compressive test ...................................................................................... 34

3.2.2 Optical microstructure and surface morphology by SEM ....................................... 35

3.2.3 Surface roughness and surface free energy ............................................................. 35

3.2.4 Hydrogen gas evolution ........................................................................................... 35

3.3 Collagen Type-I Extraction, Characterisation and Coating ........................................... 36

3.3.1 Protein extraction, precipitation and estimation ...................................................... 36

3.3.2 SDS-PAGE and Western blotting............................................................................ 36

3.3.3 Dip coating .............................................................................................................. 37

3.3.4 Characterisation of protein and coated surface by FT-IR ........................................ 37

3.4 Cell Culture .................................................................................................................... 38

3.4.1 Cell line maintenance .............................................................................................. 38

3.4.2 MTT assay for cell viability .................................................................................... 38

3.5 Techniques for Gene Profiling ....................................................................................... 39

3.5.1 Cell culture for gene expression studies .................................................................. 39

3.5.2 Total RNA isolation................................................................................................. 40

3.5.3 RNA quality check and quantification .................................................................... 40

3.5.4 cDNA synthesis/reverse –transcription reaction ..................................................... 41

3.5.5 Quantitative real time polymerase chain reaction (RT-PCR) .................................. 41

3.5.6 Agarose gel electrophoresis ..................................................................................... 42

3.6 Animal Experiments ....................................................................................................... 43

3.6.1 Implant preparation and surgery .............................................................................. 43

3.6.2 Blood sample collection, blood cell count and blood biochemical assays .............. 43

3.6.3 Radiology and DXA ................................................................................................ 44

3.6.4 Anaesthesia, surgery and euthanasia ....................................................................... 44

3.7 Bone Processing and Sectioning .................................................................................... 45

3.7.1 MMA embedding .................................................................................................... 45

3.7.2 Bone sectioning ....................................................................................................... 46

Table of Contents

vi

3.7.3 Deplasticisation ....................................................................................................... 47

3.8 Histology ........................................................................................................................ 47

3.8.1 Hematoxylin and Eosin ........................................................................................... 47

3.8.2 Tartrate Resistant Acid Phosphatase (TRAP) assay ................................................ 47

3.8.3 Masson Trichrome ................................................................................................... 48

3.8.4 Sirius Red Staining .................................................................................................. 48

3.9 Immuno-Histochemistry................................................................................................. 49

3.9.1 Collagen type-I and Vimentin ................................................................................. 49

3.10 Microscopy ................................................................................................................... 50

Chapter 4. Effect of Zirconium, Calcium and Strontium Contents on Implant-Induced

Osseointegration ...................................................................................................................... 51

4.1 Introduction .................................................................................................................... 51

4.2 Results ............................................................................................................................ 53

4.2.1 Mechanical and surface properties .......................................................................... 53

4.2.2 In vivo bone formation ............................................................................................. 59

4.3 Discussion ...................................................................................................................... 68

4.3.1 Analysis of alloying and matrix properties .............................................................. 68

4.3.2 Role of individual alloying elements in implant stability ........................................ 68

4.3.3 Role of individual alloying elements in bone formation ......................................... 69

4.3.4 Analysis of cell and matrix properties in the new bone .......................................... 69

4.3.5 Role of surface energy in in vivo biological studies ................................................ 71

4.3.6 Recommended strategies for material evaluation and significance of Sr levels in

Mg-Zr implants ................................................................................................................. 72

4.4 Summary ........................................................................................................................ 73

Chapter 5. Collagen Type-I Leads to In Vivo Matrix Mineralisation and Secondary

Stabilisation.............................................................................................................................. 75

5.1 Introduction .................................................................................................................... 75

5.2 Results ............................................................................................................................ 77

5.2.1 Experimental animals .............................................................................................. 77

5.2.2 Alloy microstructure and mechanical properties ..................................................... 77

5.2.3 Collagen coating properties ..................................................................................... 79

5.2.4 Surface roughness and surface energy ..................................................................... 80

Table of Contents

vii

5.2.5 Surface morphology by SEM .................................................................................. 82

5.2.5 Cell viability ............................................................................................................ 82

5.2.7 Histology and immunohistochemistry ..................................................................... 82

5.2.8 BMC/BMD by DXA and blood cell count .............................................................. 88

5.3 Discussion ...................................................................................................................... 89

5.3.1 Col-I coating and secondary stabilisation ................................................................ 91

5.3.2 Col-I coating and bone remodeling ......................................................................... 93

5.4 Summary ........................................................................................................................ 94

Chapter 6. Strontium Content and Collagen-I Coating Influence Osteogenesis and Bone

Resorption ................................................................................................................................ 96

6.1 Introduction .................................................................................................................... 96

6.2 Results ............................................................................................................................ 97

6.2.1 Cell viability, cell adhesion and mineralisation....................................................... 97

6.2.2 In vitro expression of bone markers ........................................................................ 98

6.2.3 In vivo bone inducing activity ............................................................................... 100

6.2.4 Implant induced bone mineralisation .................................................................... 102

6.2.5 Bone density and mineral concentration ............................................................... 103

6.3 Discussion .................................................................................................................... 104

6.3.1 Sr concentration and Col-I coating enhance osteogenesis..................................... 105

6.3.2 Role of Sr concentration in bone formation and resorption .................................. 105

6.3.3 Role of Col-I in enhanced bone formation ............................................................ 106

6.3.4 BMC/BMD and bone mineralisation ..................................................................... 107

6.4 Summary ...................................................................................................................... 107

Chapter 7. General Discussion ............................................................................................... 109

7.1 Highlights of Research ................................................................................................. 109

7.1.1 Impact of Mg alloying elements in bone formation and osseointegration: the

requirement to alloy Mg- strength versus toxicity .......................................................... 109

7.1.2 Impact of Col-I coating on osseointegration and implant stability ....................... 110

7.2 Work Required to Connect the Gaps............................................................................ 111

7.2.1 Live cell imaging for studying the cellular dynamics to degrading alloys ............ 111

7.2.2 Measurement of in vivo corrosion or biodegradation rate ..................................... 111

7.2.3 Measurement of bone bonding strength ................................................................ 112

Table of Contents

viii

Chapter 8. Conclusions and Recommendations for Future Work ......................................... 113

8.1 Conclusions .................................................................................................................. 113

8.2 Recommendations for Future Work ............................................................................. 115

Bibliography .......................................................................................................................... 117

Abbreviations

ix

Abbreviations ALP Alkaline phosphatise

Al Aluminium

ANOVA Analysis of Variance

BMC Bone Mineral Content

BMD Bone Mineral Density

BSA Bovine Serum Albumin

Ca Calcium

Col-I Collagen Type-I

dia Diameter

DAB Di-amino Benzidine

DXA Dual X-Ray Absorptiometry

ECM Extracellular matrix

EDM Electrical discharge machining

EDTA Ethylene di-amine tetra-acetic acid

Fe Iron

FT-IR Fourier transform infrared spectroscopy

HCT Hematocrit

HE Hematoxylin & Eosin

H2 Hydrogen

HRP Horse Radish Peroxidase

ht Height

Ig Immunoglobulin

IPA Isopropyl alcohol

LYM Lymphocytes

MCV Mean Corpuscular Volume

MCHC Mean cell hemoglobin concentration

Mg Magnesium

Mg2+ Magnesium ion

MMA Methyl Methacrylate

Mn Manganese

MT Masson Trichrome

MTT 3- (4, 5-Di-methylthiazol-2-yl) -2, 5-diphenyltetrazolium bromide

Abbreviations

x

OD Optical Density

PBS Phosphate buffered saline

PLT Platelet

RGD Arginyl-Glycyl-Aspartic acid

SBF Simulated Body Fluid

SDS-PAGE Sodium Dodecyl Sulfate-Polyacrylamide Gel Electrophoresis

SEM Scanning Electron Microscopy

Si Silicon

SGOT Serum glutamic–oxaloacetic transaminase

SGPT Serum glutamic pyruvic transaminase

Sr Strontium

Sr2+ Strontium ion

TBS Tris-Buffered Saline

TRAP Tartrate Resistant Acid Phosphatase

WDXRF Wavelength Dispersive X-Ray Fluorescence Spectroscopy

Wt% Weight percent

XRD X-Ray diffraction

Zn Zinc

Zr Zirconia

α Alpha

β Beta

Chemical Formulae CH3 Methyl

CO2 Carbon dioxide

CrO Chromium oxide

H2O2 Hydrogen peroxide

MgO Magnesium oxide

NaCl Sodium Chloride

O Oxygen

OH- Hydroxyl ion

SiC Silicon carbide

Abbreviations

xi

Units C Celcius

cm Centimeter

h Hours

kg Kilogram

kV Kilo Volts

M Molarity

mA Mili Ampere

mg Miligram

min Minutes

mm Milimeter

MPa Megapascal

N Normality

Ra Surface roughness

s Seconds

Xg Gravity

μg Microgram

μJ Microjoule

μL Microlitre

Abstract

xii

Abstract

Development of new biodegradable implants and devices is necessary to meet the

increasing needs of regenerative orthopaedic procedures. An important consideration

while formulating new implant materials is that they should physico-chemically and

biologically mimic bone-like properties. In this thesis, I have characterised four new

Magnesium-Zirconia (Mg-Zr)-based biodegradable alloys, containing different quantities

of Strontium (Sr) or Calcium (Ca). The mechanical and physico-chemical

characterisation of these alloys was completed using X-ray diffraction (XRD),

optical/electron microscopy, compression test and surface energy measurements,

respectively and the corrosion behaviour was established by their hydrogen production

rate in simulated body fluid (SBF). The alloys were grouped into two categories: bare

alloys without surface modification and coated alloys with Collagen Type-I (Col-I)

coating. Col-I extracted from rat-tail, and characterised using Fourier Transform Infra-

Red (FT-IR) spectroscopy, was used for dip-coating the Mg-based alloys. Both types of

alloys were implanted into cavities made in the femur bones of male New Zealand White

rabbits; cavities with no implant and observed for the same periods were kept as controls.

The bone induction, osseointegration and implant stabilisation was assessed by

histological, immunohistochemical and radiological methods. In vivo bone formation and

mineral deposition around the implants was quantified by measuring the Bone Mineral

Content/Density (BMC/BMD) using Dual-Energy X-Ray Absorptiometry (DXA).

MC3T3-E1 osteoblasts were used to evaluate the in vitro cytocompatibility of both the

bare and Col-I coated alloys. Quantitative RT-PCR was done to analyse the induction of

bone marker genes in MC3T3-E1 cells by the alloys.

Our results showed that Mg-Zr alloys containing appropriate quantities of Sr (2wt%)

were more efficient in inducing implant osseointegration than other alloys. The Col-I

Abstract

xiii

coated Mg-Zr-Sr alloy implants demonstrated increased mineral deposition on the peri-

implant surface in shorter periods of implantation with a dramatic decrease of osteoclast

activity, as compared to the uncoated alloys, indicating the role of Col-I and Sr ion

concentration in bone resorption and remodelling. The Col-I coated Mg-Zr-Ca alloys

showed a tendency to form superior trabecular bone structure with better osteoinduction

around the implants compared to the control and uncoated ones, indicating the role of Ca

and Col-I coating in osseointegration and secondary stabilisation through the

phenomenon of contact osteogenesis.

These experiments provide strong evidence that implants made with Col-I coated Mg-Zr

alloys could induce early onset of bone remodelling in the implant area, which results in

superior osseointegration and quality of bone mineralisation as compared with uncoated

Mg-Zr alloys. These results are discussed in the context of the physicochemical and

biological properties of the alloys, which could be useful in determining the constitution

of future generations of biodegradable orthopaedic implants. The Col-I coating induced

early and enhanced bone remodelling could also reduce the risk of infections during the

early stages of bone healing and wound recovery.

List of Publications

xiv

List of Publications 1.“Strontium content and collagen-I coating of Magnesium-Zirconia-Strontium implants

influence osteogenesis and bone resorption”

Mushahary D, Wen C, Kumar MJ, Harishankar N, Pande G, Hodgson P, Li Y.

Clin Oral Implants Res 2014; DOI: 10.1111/clr.12511 [Epub ahead of print].

2. “Collagen type-I leads to in vivo matrix mineralization and secondary stabilization of Mg-

Zr-Ca alloy implants”

Mushahary D, Wen C, Kumar MJ, Lin J, Harishankar N, Hodgson P, Pande G, Li Y.

Colloids Surf B Biointerfaces 2014;122:719-728.

3. “Effect of Collagen type-I on the rate of osseointegration of Ca-containing biodegradable

Mg-Zr based alloys [Abstract]” In: Proceedings of the 4th International Symposium on

Surfaces and Interfaces for Biomaterials, 24-28th September 2013, Rome, Italy.

Mushahary D, Wen C, Kumar MJ, Harishankar N, Pande G, Hodgson P, Li Y.

Eur Cell Mater 2013;26(S6):p22.

4. “Zirconium, calcium, and strontium contents in magnesium based biodegradable alloys

modulate the efficiency of implant-induced osseointegration”

Mushahary D, Sravanthi R, Li Y, Kumar MJ, Harishankar N, Hodgson P, Wen C, Pande G.

Int J Nanomedicine 2013; 8:2887-902, DOI: 10.2147/IJN.S47378

5. “Histological analysis of cells and matrix mineralization of new bone tissue induced in

rabbit femur bones by Mg-Zr based biodegradable implants”

Ragamouni S, Kumar JM, Mushahary D, Nemani H, Pande G.

Acta Histochem 2013;115(7):748-56, DOI: 10.1016/j.acthis.2013.03.004

6. “Mg-Zr-Sr alloys as biodegradable implant materials”

Li Y, Wen C, Mushahary D, Sravanthi R, Harishankar N, Pande G, Hodgson P.

Acta Biomater 2012;8(8):3177-88, DOI: 10.1016/j.actbio.2012.04.028

Awards and Presentations

xv

Awards 1. “Best Poster Award” at the 11th NJ Symposium on Biomaterials Science. “Bioactive

scaffolds – from the nano to the macro scale” held in New Brunswick, New Jersey, USA,

October 2012.

2. “Best Poster Award” at the 2nd DIRI Symposium 2011 – Frontiers in Science in TERI,

Gurgaon, India, November 2011.

3. Deakin University travel fund support for oral presentation in the 4th International

Symposium on Surface and Interface of Biomaterials ISSIB-IV, Roma, Italy, October 2013.

Presentations 1. 4th International Symposium on Surface and Interface of Biomaterials ISSIB-IV,

September 2013, Roma, Italy

2. ASMR student Research Symposium, June 2013, Melbourne, Victoria, Australia.

3. DIRI Symposium 2012 – Frontiers of Science – 2nd Edition, November 2012, Gurgaon,

New Delhi, India.

4. 11th NJ Symposium on Biomaterials Science. “Bioactive scaffolds – from the nano to the

macro scale”, October 2012, New Brunswick, New Jersey, USA.

5. 2nd DIRI Symposium “Frontiers in Science”, November 2011, Gurgaon, New Delhi, India.

List of Figures

xvi

List of Figures

Figure 2.1. The structure of long bones. .................................................................................... 7

Figure 2.2. Composition of bone matrix. . ................................................................................. 9

Figure 2.3. The generation, maturation and death of osteoclasts. ........................................... 10

Figure 2.4. Representative image of morphology of a bone-resorbing osteoclast................... 11

Figure 2.5. The phenomenon of osteoblastogenesis, describing the osteoblast generation,

maturation and death. ............................................................................................................... 12

Figure 2.6. An overview of the organisation of bone cells. ..................................................... 13

Figure 2.7. Schematic representation of osteocyte differentiation from osteoblasts.. ............. 15

Figure 2.8. Model depicting the bone remodelling compartment (BRC). .............................. 16

Figure 2.9. Figure illustrating the steps of Col-I fibril formation [134]. ................................ 32

Figure 4.1. X-ray diffraction spectra of four Mg alloys. ......................................................... 54

Figure 4.2. Scanning electron micrographs of Mg alloys before incubation in the culture

media. ....................................................................................................................................... 55

Figure 4.3. Scanning electron micrographs of Mg alloys after incubation in the culture media

for 7 days.................................................................................................................................. 56

Figure 4.4. Representative figures of contact angles of the four Mg alloys used for calculating

surface energy. ......................................................................................................................... 58

List of Figures

xvii


surface energy.. ........................................................................................................................ 58

Figure 4.6. Hematoxylin and eosin stained bone sections (100×) of Mg alloys 3 months post-

implantation.. ........................................................................................................................... 59

Figure 4.7. Tartrate-resistant acid phosphatase staining of bone sections (100×) 3 months

post-implantation.. ................................................................................................................... 60

Figure 4.8. Immunostaining of implanted bone sections (100×) by col-I, 3 months post-

implantation.. ........................................................................................................................... 61

Figure 4.9. Masson’s Trichrome stained sections (100×) of magnesium alloys 3 months post-

implantation.. ........................................................................................................................... 62

Figure 4.10. Physiological, physicochemical, and anatomical effects of the four Mg alloys

used for implantation.. ............................................................................................................. 64

Figure 5.1. Panels A and B show the phases of Mg-5Zr and Mg-5Zr-Ca alloys, respectively,

by X-Ray Diffraction.. ............................................................................................................. 78

Figure 5.2. Coomassie stained SDS-PAGE of Collagen type-I extracted from rat tail on 8%

resolving polyacrylamide gel ................................................................................................... 79

Figure 5.3. Thickness of the coated collagen on the alloy surface visualised by SEM. .......... 80

Figure 5.4. Shapes of water and glycerol droplets used to estimate the contact angles of

uncoated and Col-I coated Mg-based alloys. ........................................................................... 81

Figure 5.5. Scanning electron micrographs of the uncoated and Col-I coated Mg-based alloys.

.................................................................................................................................................. 82

Figure 5.6. Images of bone sections stained with HE and TRAP after 1 month and 3 months

implantation with coated Mg-5Zr and coated Mg-5Zr-Ca alloys.. .......................................... 84

List of Figures

xviii

Figure 5.7. Mineralisation of the newly formed bone shown by Masson’s Trichrome staining

of bone sections after 1 month and 3 months implantation, and Collagen-I immunostained

images of bone sections after 1 month and 3 months implantation. ........................................ 85

Figure 5.8. Vimentin immunostained images of coated Mg-5Zr and coated Mg-5Zr-Ca alloys

implanted bone sections after 1 month and 3 months implantation......................................... 86

Figure 5.9. Histology images of uncoated Mg-5Zr and Mg-5Zr-Ca implanted bone sections

after 3 months implantation. .................................................................................................... 87

Figure 5.10. DXA radiography detailing the BMC and BMD of the uncoated and Col-I coated

Mg-based alloys after 1 month and 3 months implantation. .................................................... 88

Figure 6.1. Cell viability ratio of MC3T3-E1 osteoblasts cultured with alloy extracts of bare

and Col-I coated Mg-Zr-Sr alloys and incubated for 24 h. ...................................................... 97

Figure 6.2. Morphology of MC3T3-E1 osteoblasts on bare and coated Mg-Zr- based alloy

surfaces after 24 h of incubation. ............................................................................................. 98

Figure 6.3. The expression of genes related to bone specific markers, proteins related to bone

ECM and proteins of non-collagenous bone matrix after 24 h (1d) and 7 days (7d) of

incubation in the respective coated-alloy extracts.. ................................................................. 99

Figure 6.4. Hematoxylin and Eosin staining of bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys after 1 month and 3 months post-implantation. ..................... 100

Figure 6.5. TRAP staining of osteoclasts for bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys after 1 month and 3 months implantation. ............................. 101

Figure 6.6. Masson’s Trichrome staining of bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys 1 month and 3 months post-implantation. ............................. 102

Figure 6.7. Alizarin Red staining of bone sections implanted with Col-I coated Mg-Zr-based

alloys after 1 month and 3 months implantation. ................................................................... 103

List of Figures

xix

Figure 6.8. Quantitative in vivo mineralisation of the peri-implant area by evaluation of femur

BMC and BMD of experimental animals implanted with Col-I coated- Mg-Zr-based alloys

after 1 month and 3 months implantation. ............................................................................. 104

List of Tables

xx

List of Tables Table 2.1. Commonly used biomaterials for implant applications .......................................... 18

Table 2.2. Properties of a desirable metallic implant, along with their effects from inadequate

functionality (67). .................................................................................................................... 23

Table 2.3. Summary of the physical and mechanical properties of various implant materials in

comparison to natural bone (8). ............................................................................................... 25

Table 3.1. Forward and reverse primers of genes used for quantitative RT-PCR experiments.

.................................................................................................................................................. 42

Table 4.1. Nominal composition of Mg-based alloys ............................................................. 53

Table 4.2. Mean left and right contact angle (CA) values and the respective surface energies

(SE), with statistical standard deviations, of all the four Mg-based alloys. ............................. 57

Table 4.3. Blood cell count of control and experimental animals 3 months after implantation.

.................................................................................................................................................. 66

Table 4.4. Values of serum biochemical parameters of control and experimental animals after

3 months implantation.............................................................................................................. 67

Table 5.1. Measured body weights of experimental animals for Col-I coated Mg-5Zr and Mg-

5Zr-Ca alloy implants after 1 month and 3 months implantation. ........................................... 77

Table 5.2. Surface roughness of the bare and Col-I coated Mg-based alloy discs. ................. 80

Table 5.3. Individual surface energy of Col-I coated Mg-based alloys calculated from the

respective left and right contact angles measured using water and glycerol. .......................... 81

Table 5.4. Blood cell count of the control and experimental animals implanted with Col-I

coated Mg-5Zr and Mg-5Zr-Ca alloy implants. ...................................................................... 90

Chapter 1. Introduction

1

Chapter 1

Introduction

Magnesium (Mg) alloys are being studied as suitable orthopaedic biodegradable implants

owing to their degradability and resemblance to human bone. Mg is the eighth most abundant

element and is present in the Earth’s crust in about 2% [1]. It is the third most plentiful

element dissolved in a concentration of about 0.13% in seawater and its compounds are

produced from seawater, well and lake brines and bitterns, as well as from various minerals

like dolomite, carnallite, magnesite and brucite [1].

Mg is naturally present in biological systems as Mg2+ ions and has a stimulatory role in bone

formation [2-5]. Because of its low weight and natural biocompatibility, it has been suggested

as a good implant material. In addition, Mg has some chemical properties that make it a very

close mimic to natural bone. It has a density of 1.74 g/cm3, its elastic modulus is about 45

GPa and it has a compressive strength close to that of the cortical bone [6-10]. Most

importantly, due to their biodegradable property, Mg alloy implants are converted to a

soluble, non-toxic oxide that is safely excreted in the urine [10, 11]. Hence, a second surgery

for the removal of Mg implant is not required.

Despite all the above advantageous properties, Mg has certain constraints as a bio-implant.

Upon implantation, Mg alloy degrades very rapidly in vivo due to its low corrosion

resistance. The degradation takes place according to the following reaction:

Mg + 2H2O→ Mg(OH)2 + H2↑

This leads to the production of a large volume of hydrogen gas and results in a loss of

mechanical integrity even before the tissue has sufficiently healed and new bone tissue has

adequately regenerated [12-16]. Hydrogen gas production leads to increased alkalinity and


2

creates health risks for the patients [17]. This becomes the most challenging disadvantage of

using pure Mg as bio-implants. Its poor mechanical strength and elastic modulus also do not

meet the requirements of the patients to sustain the rigors of daily activity [6, 8, 14, 16].

Alloying Mg with other metals can enhance its performance because of the improved

corrosion resistance and mechanical properties. Thirdly, the biocompatibility of existing Mg

alloys is inadequate for biomedical applications because traditionally, they have been used

for non-biological applications, including pyrotechnic and metallurgical, chemical,

electrochemical, and structural [18]. Enhancement of the biocompatibility of Mg-based alloys

is therefore a major challenge, and achieving optimal level of biocompatibility has become a

critical criterion for selecting any biodegradable implant material.

1.1 Hypothesis

Although alloying of Mg can increase the corrosion resistance of these alloys to some extent,

it cannot significantly improve the bone response upon implantation. The surrounding tissue

interacts only with the surface of the implant. Osseointegration depends on the bio-

mineralisation of the implant into the surrounding tissue. As this interaction between the

implant and the cells/tissues at the tissue–implant interface is a surface phenomenon, surface

properties play a major role in determining both the biological response to implants and the

material response to the physiological condition. Therefore, modification of the alloy surface

is necessary to form a hard and yet biodegradable, biocompatible and corrosion resistant

layer, while still preserving the bulk properties of the alloy. With this background, this thesis

hypothesises the following:

1. Zr is a biocompatible metal and has excellent positive impacts on the mechanical and

corrosion properties of Mg alloys, when it is used at ≤ 5wt%.

2. Sr has a positive impact on both the mechanical and biological properties of Mg-Zr-Sr


3

alloys and its concentration also should be ≤5wt%.

3. The surface topography of Mg-Zr-Sr alloy samples is compatible as in vivo implants and

these materials can generate new bone when used as bio-implants.

4. Surface coating of Mg-Zr based alloys with Collagen Type-I (Col-I) will enhance the in

vivo biocompatibility and new bone forming ability of these alloys.

1.2 Thesis Objectives

The objective of this thesis was to evaluate the in vivo biocompatibility, osteogenicity and

mineral deposition by four Mg-Zr based alloys namely: Mg-5Zr, Mg-Zr-2Sr, Mg-2Zr-5Sr and

Mg-5Zr-Ca. These objectives were achieved by investigation of the following aspects of the

alloy implants:

Demonstrating the role of alloying elements - Zirconia (Zr), Calcium (Ca) and

Strontium (Sr) in Mg-based alloys in modulating osseointegration, by looking into the

physico-chemical properties of these alloys.

Surface modification of Mg-Zr-Ca alloys by Col-I coating and looking into its matrix

mineralisation and secondary stabilisation.

Evaluation of the effect of two different concentrations of Sr on bone growth and

mineral deposition, and the bone resorption capacity of Mg-Zr based alloys, by Col-I

coating.

1.3 Thesis Outline

Chapter two discusses the literature review concerning the thesis, with emphasis on bone

biology and types of bone cells and bone matrix. Also discussed is the role of biomaterials in

orthopaedics with emphasis on biodegradable Mg-based alloys. The various surface


4

modification techniques developed to date, to improve the material performance as implants

and for enhancing their bioactivity, are also discussed in detail.

Chapter three lists the materials and methodology used in the experiments. However,

methods specific to each chapter are discussed in detail in the individual chapters.

Chapter four investigates the constituents and the role of individual alloying elements in

controlling the in vivo implant osseointegration. The effect of mechanical and physico-

chemical properties on the in vivo peri-implant cellularity and bone formation has been

discussed in this chapter, with highlights on proximity and the extent of bone formation by

individual implants.

Chapter five examines the necessity and importance of surface modification by Col-I coating

to Mg-Zr-Ca alloys and discusses the influence of coating on implant secondary stabilisation.

The matrix mineralisation and bone remodelling has been discussed in the context of Col-I

coating.

Chapter six describes the role of Sr-wt% and Col-I coating on the quality of new bone

formation by looking into the qualitative trabecular bone mineralisation and bone resorption.

Also discussed in this chapter is the significance of Sr concentration in the in vitro gene

expression of bone formation markers.

Chapter seven discusses the summary of the findings and highlights from this study and the

work required to connect the gaps between the results.


5

Chapter eight provides the conclusion of the thesis and discusses the directions for future

research.

Chapter 2.Literature Review

6

Chapter 2

Literature Review

2.1 Introduction

Orthopaedic implants, including modern dental implants, have achieved encouraging and

successful results in the field of health care. The available orthopaedic materials have helped

to restore structural integrity and functions in many cases including bone fractures [12] and

total joint replacement [19], thereby improving quality of life. However, before efficient

substitution of damaged bone tissue is achieved, it is important to understand the intricacies

involved in the organisation of bone tissue as a whole, as the microenvironment of the

scaffolds play an important role in stimulating tissue regeneration. This review summarises

the architecture of bone tissue with a major focus on different cell types present in the bone

and their functions. The orthopaedic biomaterials and materials related to Mg and its alloys,

developed till date as implant materials for bone tissue engineering applications are

discussed. The surface properties and surface modifications of these materials and their

physico-chemical properties in relation to their responses to the cell biological processes are

highlighted.

2.2 Bone Structure and Morphology

Bone consists of two types of osseous tissue namely the compact (cortical) bone and the

spongy (trabecular) bone. Compact/cortical bone is very dense and solid and is found on the

surface of the bone, whereas spongy/trabecular bone is relatively porous and is found in the

interior of bone. The relative amounts of compact and spongy bone are related to function.

The compact and trabecular bone are part of the long bones.


7

The surfaces of bones are wrapped with connective tissue called the periosteum, consisting of

a fibrous outer layer and a cellular inner layer. It functions to separate the bone from the

surrounding tissues and to provide a pathway for circulatory and nervous supply, along with

assistance in bone growth and repair. It is also essential in forming joints because of its

adjacent placement to tendons and ligaments. Beneath the periosteum lies endosteum, that

lines the marrow cavity, the trabeculae of the spongy bone, and the inner surfaces of the

central canals. A representative figure of a long bone, the humerus, is given in Fig 2.1.

The long bone consists of a shaft, the diaphysis, and swollen ends on both sides, the

epiphysis, in between two metaphysic regions. A cavity-like nutrient foramen allows nerve

fibres and blood vessels to enter the diaphysis. Inside the diaphysis is a space called the

medullary (or marrow) cavity that contains bone marrow. The process of bone remodelling

occurs in this region, between the outer periosteum and an inner endosteal layer. The cortical

Figure 2.1 The structure of long bones


8

bone contains a series of numerous concentric circular osteons, surrounded by interstitial

lamellae, which in turn are connected by haversian canals that house the blood capillaries,

veins and nerves of the bone tissue.

2.3 Bone Composition

Almost all the bone mass is made up of matrix. Approximately 65% of bone is composed of

inorganic component and mineral (primarily carbonated hydroxyapatite) that provides

strength and resistance to compression. The organic component of bone, mostly type-I

Collagen, contributes about 20-25% to its composition (Fig 2.2). The remaining almost 10%

is made up of water which remains bound to the collagen-mineral composite. The organic

component of the bone mass is called the osteoid. Collagen fibres of type-I are the major

component of the osteoid, around 90% of the organic portion, with smaller amounts of type

III and V that are also found in the surrounding zone of the bone cells. The remaining 10% of

the organic component makes up the non-collagenous proteins (NCPs), which play important

roles in regulating collagen formation, cell attachment and mineralisation. Of this small

amount of NCPs, almost 85% constitutes the extracellular proteins and the remainder is the

cellular proteins of the bone cells.


9

2.4 Biology of Bone Cells

The bone mass of a living individual is maintained by the coordinated activities of bone cells.

The major types of cells that make up the bone can be typically divided into three types,

namely osteoclasts, ostoblasts and osteocytes.

2.4.1 Osteoclasts

Osteoclasts are the cells responsible for bone resorption by their ability to dissolve mineral

and degrading organic bone matrix in collagen fibres. Osteoclasts are derived from

hematopoietic stem cells, or hematopoietic monocyte-macrophage lineage [20], and so the

earliest precursor of the osteoclasts is found within the granulocyte-macrophage colony-

stimulating unit (CFU-GM) in bone marrow. These mononuclear precursors fuse to form the

mature multinucleated osteoclasts and differentiate through the same pathway as that of

macrophages and dendritic cells. Hence, a pro-myeloid precursor can differentiate into:

Figure 2.2. Composition of bone matrix. The non-collagenous proteins (NCPs) of the organic

matrix contribute to cell adhesion and mineralisation with the cellular and extracellular

proteins.


10

osteoclasts, when exposed to receptor activator of NF-kB ligand [RANKL/tumor necrosis

factor related activation induced cytokine (TRANCE), osteoprotegerin ligand (OPGL) or

osteoclast differentiation factor (ODF)], a macrophage when exposed to macrophage colony-

stimulating factor (M-CSF), or a dendritic cell when exposed to granulocyte-macrophage

colony-stimulating factor (GM-CSF) [21-24].

The bone resorption by osteoclasts involves the following stepwise cellular activities:

1. Migration to the resorption site and its attachment to the bone

2. Polarisation and formation of new membrane domains

3. Dissolution of hydroxyapatite

4. Degradation of organic matrix

5. Removal of degradation products from the resorption cavity/lacuna

6. Apoptosis of the osteoclasts or their return to the non-resorbing state

Figure 2.3. The generation, maturation and death of osteoclasts. Osteoclast precursors

proliferate in response to cell specific receptor-activated molecules and growth factors and

form committed osteoclast precursors which then differentiate to form multinucleated

osteoclasts. Multinucleation is a key morphological feature of osteoclasts that distinguishes

mature osteoclasts from immature cells.


11

After migration of the osteoclast to the resorption site, the resorbing osteoclasts become

highly polarised cells and secrete numerous specialised membrane domains including the

sealing zone, a ruffled border, a functional secretory domain and a basolateral membrane (Fig

2.4). The process of bone resorption occurs in resorption bays called Howship’s lacuna. The

osteoclasts attach to the bone membrane through the sealing zone and form a resorbing

organelle by fusion of intracellular acidic vesicles with the ruffled border, the region of

plasma membrane facing the bone [25, 26].

Figure 2.4. Representative image of morphology of a bone-resorbing osteoclast. FSD-

functional secretory domain, SZ- sealing zone. Mature osteoclasts become polarised with the

apical domain towards the bone surface. Bone dissolution occurs through the action of matrix

metalloproteases (MMPs), Cathepsin K (Cat K) and Tartrate-Resistant Acid Phosphatase

(TRAP). The degradation debris are internalised into the cell and degraded by the lysosomes

or via transcytosis.


12

During resorption, the plasma membrane attaches tightly to the bone matrix and seals the

resorption site from its surroundings. The osteoclasts then resorb the mineralised bone matrix

by dissolution of crystalline hydroxyapatite and proteolytic cleavage of the collagen-rich

organic matrix. The tightly packed hydroxyapatite crystals are dissolved by targeted secretion

of HCl through the ruffled border into the resorption lacuna [25, 26]. The proteolytic

enzymes- lysosomal cysteine proteinases, matrix metalloproteinases (MMPs) (MMP-9 -

gelatinase B/Collagenase Type-IV) and cathepsin K play significant roles in the resorption

process [27-29]. After it has fulfilled its resorption task, the osteoclast undergoes fission into

mononuclear cells or it dies by apoptosis.

2.4.2 Osteoblasts

Osteoblasts are the bone forming cells that originate from mesenchymal progenitor cells. The

functions of osteoblasts include bone matrix protein secretion and bone mineralisation. After

completion of bone matrix formation, some mature osteoblasts remain entrapped in the bone

matrix and become osteocytes, some flatten to cover the quiescent bone surface as bone

lining cells and the rest of them die by apoptosis (Fig 2.5).

Figure 2.5. The phenomenon of osteoblastogenesis describing the osteoblast generation,

maturation and death. Mature osteoblasts are surrounded by the bone matrix and are

transformed into osteocytes, or are flattened to cover the quiescent bone surface as lining

cells, or undergo cell death by apoptosis.


13

Osteoblast cells are cuboidal with a diameter of 15-30 μm. They contain a large nuclei

located close to the basal membrane of the cell, an enlarged Golgi apparatus on the apical

surface of the nuclei and extensive endoplasmic reticulum. They aid in bone formation by

generating, depositing and mineralising bone matrix. They are present in the periostium

region of the bone; however, a major reservoir of osteoblast and osteoprogenitor cells is the

endosteum surface of the bone [30]. Osteoblasts express high levels of alkaline phosphatase

(ALP) and osteocalcin, and the circulating concentrations of these proteins reflect the rate of

bone formation. Osteoblasts also secrete type-I Collagen abundantly along with other

specialised matrix proteins to form osteoid, the organic phase of bone that serves as a

template for subsequent deposition of mineral in the form of hydroxyapatite. The osteoblasts

deposit the extracellular matrix beneath the periosteum, entrapping the blood vessels by the

formation of ridges and grooves, which eventually fuse together with the deposited matrix

Figure 2.6. An overview of the organization of bone cells at the periosteum and the

endosteum, and the formation of osteoid by the osteoblasts.


14

thereby, forming the osteon (Fig 2.6) [31].

The proliferating osteoblast progenitors express transcription factors of the helix-loop-helix

family (eg, DNA-binding protein inhibitor ID) and the osteoprogenitor population is

maintained by inhibiting the expression of genes that characterise the mature osteoblast

phenotype. Two major proteins essential for establishing the osteoblast phenotypes are

RUNX2 and transcription factor Sp7 or Osterix and the two major signalling pathways that

promote osteoblast differentiation are those activated by BMPs and Wnts.

Wnt proteins play key roles in the development and homeostasis of bone through their

involvement in cell differentiation, proliferation and apoptosis. They stimulate the

differentiation of preosteoblasts and the undifferentiated mesenchymal cells toward the

osteoblastic lineage. The population of osteoblasts can increase either by an increase in their

rate of production from progenitors, a decrease in the rate of their death by apoptosis, or a

combination of the two. After completing bone matrix synthesis, some osteoblasts become

flattened and cover the inactive bone surface as bone-lining cells.

2.4.3 Osteocytes

Osteocytes are formed from osteoblasts that become embedded in the bone matrix during the

process of bone mineralisation. Osteocytes are the most abundant cells in bone comprising

more than 90% of cells in the bone matrix. Mature osteocytes are stellate shaped enclosed

within the lacuno-canalicular network of bone. They radiate slender cytoplasmic processes in

all directions called canaliculi. These radiating canaliculae act as means of communication

between the neighbouring osteocytes and the bone surface through gap junctions. It is

through these connections that osteocytes facilitate the access of oxygen and nutrients (Fig

2.7).


15

Osteogenic cells arise from mesenchymal stem cells (Fig 2.5). The mechanism of osteocyte

formation and differentiation from osteoblasts is however unknown. Stimulation of

osteoblasts to extend cytoplasmic projections and to make contact with embedded cells may

result in their differentiation into osteocytes [32]. Hence, the stimuli that alter osteoblast fate

would also affect osteocyte formation. Nevertheless, osteocytes are responsible for the

biomechanical regulation of bone mass and structure through the formation of the bone

remodelling compartment (BRC) (Fig 2.8) [33], as osteocytes detect fatigue-induced

replacement of damaged bone through remodelling. The osteocytes sense the need for bone

resorption and send signals to lining cells, which retract from the bone surface to form a

BRC. Osteocytes in turn release molecules, including sclerostin and Dkk-1, which induces

osteoblasts to differentiate and become functional. Osteocytes also respond to changes in

mechanical load by inducing local changes in bone mass through modelling. Furthermore,

osteocytes detect changes in the circulating hormonal levels and respond to it by altering the

rate of bone formation and resorption.

Figure 2.7. Schematic representation of osteocyte differentiation from osteoblasts. The matrix-

producing osteoblasts either become an osteocyte, a lining cell, or undergo apoptosis.


16

Osteocytes produce high levels of sclerostin that can be detected in the canaliculi. Sclerostin

acts as an antagonist to the BMP family of proteins and the Wnt signalling pathway. Wnt are

essential for osteoblastogenesis and provide the stimulus for differentiation of mesenchymal

progenitors to the osteoblast lineage, along with regulation of osteoblast activity. Osteocytes

regulate the function of osteoblasts and osteoclasts. In response to mechanical and hormonal

signals, osteocytes secrete factors (such as osteoprotegerin OPG, RANKL, and sclerostin)

that affect other bone cells by paracrine or autocrine mechanisms, and hormones that affect

other tissues by endocrine mechanisms.

The life span of osteocytes is determined by bone turnover, in which the osteoclasts resorb

bone and destroy osteocytes [34] along with the bone itself. Osteocytes may live for decades

if the particular bone in which they reside has a slow turnover rate. Osteocytes die by

apoptosis similar to osteoblasts and osteoclasts and decreased osteocyte viability indicates

Figure 2.8. Model depicting the bone remodelling compartment (BRC). Osteoclast precursors

are transported to the BRC by marrow capillaries where they differentiate to mature

osteoblasts under the effect of differentiating factors such as RANKL and OPG, which are

derived from pre-existing osteocytes and initiate bone remodelling.


17

bone fragility. The rest are re-embedded during the process of new bone formation which in

turn is resorbed later by osteoclasts. The level of apoptotic osteocytes is higher in unloaded

bones and in bones exposed to high levels of mechanical strain [35]. In either case, increased

apoptosis in osteocytes can be seen before any osteoclast resorption, and the apoptotic

osteocytes accumulate in areas that are subsequently removed by osteoclasts.

2.5 Biomaterials and Biocompatibility- General Concepts

A biomaterial or a bioactive material refers to any non-living material that is designed to

mimic a specific pre-existing micro niche and can interact, when integrated, with the

biological system of the animal/human body. Biomaterials can induce the host cells to

differentiate and regenerate to the desired tissue according to physiological pathways. Hence,

a material that is inserted into a host tissue and is destined to induce a desired cellular

behaviour belongs to a biomaterial family.

Based on the material-host tissue reaction, biological materials are classified into three

distinct categories:

1. Biotolerant materials are separated from bone tissue by a layer of fibrous tissue.

2. Bioactive materials have the property of establishing chemical bonds with bone tissue,

known as osseointegration.

3. Bioinert materials have the possibility of direct contact with the adjacent bone tissue.

However, there is no chemical reaction between the implant and the tissue.

Examples of some of the accepted biomaterials are given below:


18

Types of biomaterials Materials used

Metals Stainless steel, Ti-based alloys, Co-Cr based alloys, Au alloys, Ag products, platinum electrodes

Ceramics Alumina, zirconia, bioactive glasses, hydroxyapetite, calcium phosphates, porcelains, carbons

Polymers Polyethylenes, polypropelenes, polyamides, polyesters, polyurethanes, poly (methyl-methacrylate) (PMMA),

hydrogels

Composites Bisphenol-A-Glycidyl-Methacrylate (Bis-GMA)-quartz/silica fillers, PMMA-glass fillers

Natural biomaterials Silk fibroins, collagen, coral, chitins, keratin, cellulose

Table 2.1. Commonly used biomaterials for implant applications

The recognition of specific types of interface between a biomaterial and the host tissue

brought the concept of biocompatibility. Also known as tissue compatibility, according to

David F. Williams, biocompatibility refers to the ability of the substrate to support cellular

activity, including the facilitation of molecular and mechanical signalling systems, in order to

optimise tissue regeneration, without eliciting undesirable local or systemic responses in the

eventual host” [36]. Hence, biocompatibility is the degree of reaction caused by a foreign

material in the body. The biocompatibility of the material with the host tissue is a mutual

interaction eliciting a response from both the material and the host tissue. These changes in

the material and biological tissue may further lead to several other changes.

2.6 Desired Properties of Biomaterials

The most important property of a biomaterial is that it should have a biocompatible

chemical composition so as not to elicit an adverse tissue reaction when used for its

purpose.

It should have a very good resistance to degradation and corrosion.


19

It should have enough strength to sustain the mechanical load of the host joint during

and after the surgery.

A low modulus (elasticity) to minimise bone resorption.

A high wear resistance to minimise wear and tear during weight loading.

Apart from the above clinical requirements, the manufacturing and economic

requirements of materials should also be considered. Manufacturing requirements

prioritise the materials to have properties that permit its fabrication in the optimum

design configuration. Although the economic requirements are not a critical issue in

human applications, cost effective development of materials must be considered for

veterinary applications.

2.7 Applications of Biomaterials

Oral/dental applications: Replacement of tooth and supporting gum tissues due to

cavities and diseases can be done by a variety of materials. Metallic biomaterials like

Ti and Co-Cr based metals have been recently used for tooth anchorage applications

[37, 38]. Other tooth implants consist of ceramics including alumina [39] and dental

porcelains [40, 41].

Orthopaedic applications: Osteoarthritis and rheumatoid arthritis can lead to

disturbances in the movement of joints such as hips, knees and elbows along with

considerable amounts of pain. Normally used for load bearing applications, metallic,

ceramic and polymeric biomaterials have been used for replacements of such load-

bearing joints. Recently developed orthopaedic biomaterials include Zr, Ti and other

ceramics used for joint replacements [42], porous poly (para-phenylene) scaffolds

[43], whereas hydroxyapatite assisted implants have been used for bone bonding

applications [44, 45].


20

Cardiovascular applications: The most important criteria for the cardiovascular

implants is to have a unique blood biocompatibility so that the host body does not

reject the implant and to avoid the occurrence of an adverse thrombogenic (clotting)

or hemodynamic (blood flow) response. Cardiovascular applications include materials

for heart valves, artificial heart implants, vascular grafts and stents, and ventricular

assisting devices. Materials used for fabrication of cardio-vascular implants include:

polyhydroxyalkanoates [46, 47], synthetic degradable polymers such as Polyglycolic

acid (PGA), Polylactic acid (PLA) and Poly-E-caprolactone (PCL) [48-50], stainless

steel, Co-Cr alloys and titanium.

Drug-delivery systems: Many attempts have been made to incorporate drug

reservoirs into implantable devices for a sustained and preferably controlled release.

Recent evolution of implantable drug delivery systems includes Poly (lactic-co-

glycolic acid) (PLGA) based biodegradable polymers [51], silk fibroin [52] and

chitosan [53].

Ophthalmics: The tissues of the eye can suffer from several diseases leading to

reduced vision and cause blindness. Researchers are continuing into the replacement

of the damaged eye tissue with an intraocular lens made of natural or synthetic

materials like silk fibroin and chitosan [54-56].

2.8 The Necessity of Biomaterials in Bone Tissue Engineering

Bone is a highly vascularised tissue that can remodel itself throughout its life. Apart from

giving structural functions to the body, it is also involved in maintaining homeostasis of the

body through its storage of Ca and P ions and by regulating the electrolyte concentrations in

the body. Due to its high regenerative capacity, majority fractures heal well without the need

for external intervention. In spite of this, bone is often prone to injuries due to trauma, aging


21

or diseases and requires surgical intervention. The US data on bone therapy market report

estimated 6.2 million bone fractures every year, of which 10% cases fail to heal properly

[57]. It is estimated that around 10 million people are affected by osteoporosis and the

number is likely to increase to 14 million by 2020 [58, 59]. Recent statistics from the

National Osteoporosis Society (NOS) report an estimated expenditure of £1.73 billion for

treating hip fractures in the United States and massive amounts spent by the National Health

Service (NHS) for treating osteoporosis related fractures [60]. Furthermore, it is estimated

that approximately 2.2 billion bone graft procedures are being performed worldwide,

annually to repair bone defects [61]. Currently, the gold standard treatment involves the use

of autografts obtained by harvesting the donor bone from a non-load-bearing site in the

patient and transplanting into the defective site. There has been a growing need for

considerable autologous bone grafting in spinal fusion, which was reported to be the 19th

most common in-patient procedure in the United States in 2003 [62]. Autologous bone

transplantation gives the best clinical outcome, as there remain few complications related to

immunity and diseases, unlike the allogenic bone grafts from cadavers. Unfortunately, there

is a very short supply of autografts and considerable donor site morbidity associated with its

harvest [63, 64]. Therefore, it is essential to find ideal substitutes and materials for restoring

the normal functioning of the damaged bone tissue.

The success of orthopaedic biomaterials in healing bone defects in the past few decades has

led to further developments in this field. While materials proposed for implantation in the

past were designed to be ‘bio-inert’, materials scientists have now moved towards the design

of ‘bioactive’ materials that establish bonding with biological molecules or cells and

regenerate tissues [65, 66]. Applications of these bioactive materials are many, including

craniofacial surgery, dental and fracture fixation and total joint arthroplasty. For orthopaedic

applications, biomaterials should have the properties of biocompatibility, possess adequate


22

mechanical properties with high corrosion and wear/tear resistance and provide bioactive

surfaces for osseointegration upon implantation. For implantation in bone applications,

materials should preferably be osteoinductive (capable of promoting the differentiation of

progenitor cells down an osteoblastic lineage), osteoconductive (support bone growth and

encourage the ingrowth of surrounding bone), and capable of osseointegration (integrate into

surrounding bone). The ideal bioactive material is one which gets resorbed and replaced over

time by the body's own newly regenerated biological tissue [65]. The materials used for

orthopaedic surgery include bioactive inorganic materials, polymers, hydrogels, fabrics,

ceramics, composite materials, metals and natural materials [67]. The major requirement of

prosthetics is their use in load bearing applications and the criteria of long-term performance.

However, this review focuses in the directed use of biomaterials as orthopaedic implants for

load bearing applications.

2.8.1 Metallic implants in orthopaedics

The increasing trend to replace defective tissue in orthopaedics has come from the use of

metals, especially in total joint replacements such as hip and knee. Unfortunately, the current

available metallic materials have been found to fail in long-term usage, as the load bearing

implants must satisfy specific properties to be used for long-term applications. The desired

metallic implants need to fulfil the mechanical and non-mechanical criteria to be used as

devices for bone repair and healing. Table 2.2 describes the properties that make the metal a

suitable implant material [68]. The currently used major groups of load bearing metal

implants for applications in the biomedical field include, stainless steel, cobalt-chromium

alloys and titanium and its alloys.


23

2.8.1.1 Stainless steel

Stainless steel is widely used in orthopaedic applications due to its low cost, superior

fabrication properties, high corrosion resistance, and its availability [76]. Stainless steel is the

generic name for a number of steels used primarily because of their resistance to a wide range

of corrosive agents. This resistance is due to their high Cr content (17-20 wt%) that forms a

chromium oxide coating (Cr2O3). The Cr in the stainless steel has a great affinity for oxygen

and hence leads to formation of a passive, highly adhesive and self-healing chromium oxide

layer on the surface of the steel [74]. The main stainless steel used for manufacturing

implants is an austenitic type called 316L [77], which is widely used in trauma-related

temporary devices such as fracture plates, screws and hip nails. However, austenitic stainless

Requirements Implications

Mechanical properties

1. Long fatigue life [68] Implant mechanical failure and revision surgery

2. Adequate strength [69] Implant failure, pain to patient and revision surgery

3. Modulus equivalent to that of bone [69, 70]

Stress shielding effect, loosening failure, revision surgery

4. High wear resistance [71] Implant loosening, severe inflammatory response, destruction of the healthy bone producing debris which can go to blood

Non-mechanical properties

5. High corrosion resistance [72] Releasing non-compatible metallic ions and allergic reactions

6. Biocompatibility [73, 74] Body reaction and adverse effects in the organic system

7. Osseointegration [69, 75] Fibrous tissue between the bone and the implant, not well integration of the bone and implant and finally implant loosening.

Table 2.2. Properties of a desirable metallic implant, along with their effects from inadequate

functionality (67).


24

steel has a very poor wear resistance with a high modulus of 200 GPa compared to the natural

bone, resulting in rapid loosening of the implant due to the debris generated that creates risks

of allergic reactions [78], which limits its use for medical applications.

2.8.1.2 Cobalt-Chromium alloys

After stainless steel, Co-based alloys are the most commonly used implants in orthopaedic

applications. They can form a passive oxide layer in a chloride-containing environment and

hence are highly resistant to corrosion within the human body [79, 80]. There are two types

of Co-Cr alloys commonly used in dentistry and artificial joints, and in making the stems of

prosthesis respectively, namely Co-Cr-Mo and Co-Ni-Cr-Mo. These materials have better

mechanical properties compared with stainless steel [74, 79], such as high resistance to

fatigue and a good wear and corrosion resistance. However, these alloys may release

elements like Cr and Co during the process of corrosion [73], which are even more toxic than

the corrosion products of stainless steel. Moreover, their high elastic modulus of 220-230

GPa, which is almost similar to that of stainless steel (approximately 200 GPa), makes them

unsuitable for implant purposes.

2.8.1.3 Ti and its alloys

Ti and its alloys contain many desirable characteristics of an implant material, which make

them excellent metals for orthopaedic applications over stainless steel and Co-Cr alloys [70,

81]. The high strength, good resistance to corrosion, complete inertness to body environment,

enhanced biocompatibility, moderate elastic modulus of approximately 110 GPa displayed by

Ti are suitable choice for implantation. Ti and its alloys have extensive applications in

biomedical field, especially in load bearing applications such as artificial hip joints and

artificial knee joints, and as stents, dental implants, bone plates and screws. The


25

commercially available pure Ti (CP Ti) and Ti–6Al–4V ELI (Ti64) are the most commonly

used titanium materials for implant applications. But the long-term use of these materials may

release of certain unwanted ions like aluminium and vanadium which are associated with

diseases like Alzheimer and neuropathy [70, 74]. Further, Ti alloys suffer from severe wear

in the presence of friction from other metals or from itself [82]. Therefore their application

becomes limited only to the locations where wear resistance is of little or no importance.

2.8.2 Mg and Mg alloys as biodegradable implants

For many years, stainless steel, cobalt-chromium, and titanium alloys have been the primary

biomaterials used for load-bearing applications. However, such conventional metallic

materials are biocompatible but not biodegradable, and they provide only short-term

structural support. Biomaterials that can be reabsorbed into the body after healing, like Mg-

based alloys, provide a number of benefits over traditional permanent implants. The

disadvantages of using conventional surgical alloys, like stainless steel, cobalt, chromium and

nickel-based alloys have been already discussed in the previous section. Conversely, Mg has

many attractive physical characteristics very close to that of the human bone (Table 2.3).

Mg has high specific strength that increases its load bearing capacity. Its low elastic modulus

improves stress-shielding effects between the bone and the implant material. Mg is essential

for human metabolism and is associated to various pathological conditions [5]. Mg is the

second most common intracellular ion in our body and serves as a cofactor for many

enzymatic reactions ranging from muscle contraction to neuronal control [83]. Almost 53%

of Mg in the body is stored in bone in the form of hydroxyapatite in the inorganic matrix

[84]. Magnesium-based materials were first introduced by Lambotte as a pure Mg plate,

along with gold-plated steel nails, to secure lower leg bone fractures [8, 85]. However, the

high rate of in vivo corrosion of the implant made it degrade entirely in as early as 8 days.


26

Since then, numerous attempts have been made to increase the corrosion resistance of Mg

implants by studying its corrosion properties both in vitro [86] and in vivo [6]. Eventually

these studies highlighted other beneficial in vivo properties of Mg and its alloys like its ability

to stimulate bone growth and healing [6]. Furthermore, the corrosion and degradation

products of Mg are harmless and are excreted through urine [87].

Properties

Natural bone

Mg Ti alloy Co-Cr alloy

Stainless steel

Synthetic HAP

Density (g/cm3) 1.8-2.1 1.74-2.0 4.4-4.5 8.3-9.2 7.9-8.1 3.1

Elastic modulus (GPa)

3-20 41-45 110-117 230 189-205 73-117

Compressive yield strength (MPa)

130-180 65-100 758- 1117

450-1000 170-310 600

Fracture toughness (MPam1/2)

3-6 15-40 55-115 N/A 50-200 0.7

Table 2.3. Summary of the physical and mechanical properties of various implant materials in

comparison to natural bone (8).

However, the major drawback of using Mg as implant material is its low corrosion resistance

which leads to rapid release of degradation products and causes a reduction in the mechanical

integrity of the implant, even before the defective bone is healed properly [8]. The high

corrosion rate also leads to rapid production of hydrogen gas leading to formation of gas

bubbles that can accumulate around and cause implant loosening. The localised formation of

hydrogen gas also increases the pH around the implant area. To successfully use Mg alloys as

implant material, their properties should be such that the cells can synthesise and deposit

extracellular matrix (ECM) for their support and function before the structural integrity of the

implant is compromised. To date, many elements have been explored for alloying pure Mg to

improve its in vitro and in vivo performance by studying the degradation and biological

behaviour of its alloys like AZ31 (Mg–3Al–1Zn), AZ91/AZ91D (Mg–9Al–1Zn), WE43

(Mg–4Y–3RE) and LAE442 (Mg–4Li–4Al–2RE) [12, 88, 89]. However, many of these


27

alloys contribute to the cell cytotoxicity because of its elemental constitutions [90-93].

Hence, development of new biodegradable Mg alloys with the combination of completely

biocompatible alloying elements is necessary. Further, several surface modification methods

have also been established to improve the corrosion performance of Mg alloys and to enhance

the cellular properties upon its implantation.

2.9 Surface Modification of Magnesium Alloys

As discussed above, Mg-based implants are becoming promising candidates for fracture

fixation because of their stress shielding capacity during fracture healing and bone

remodelling. Mg-based biodegradable implants are now being used for cardiovascular stents

and have also been successfully tested in animal models [94, 95] and has progressed with

clinical trials in patients with peripheral arterial obstructions and coronary artery disease [96,

97]. However, much effort is required to avoid its fast degradation in vivo for orthopaedic

applications in fracture fixation. The following are the main concerns regarding their

orthopaedic applications [87, 98].

1. Rapid degradation before the time required for fracture healing, leading to insufficient

mechanical support for bone fracture fixation;

2. High degradation resulting in local accumulation of alkaline ions that induces local

inflammation after implantation;

3. Rapid release of Mg ions due to high degradation rate causing adverse effects, such as

diarrhoea and abdominal cramps, leading to renal failure.

To overcome these drawbacks, surface modification methods such as coatings on Mg-based

materials are proven to be efficient to control the degradation and corrosion rate [17, 99,

100]. The surface modification methods developed in Mg alloys to date can be divided into

two major groups - chemical modification and physical modification.


28

2.9.1 Chemical modification

Chemical modification involving the use of chemical or electrochemical treatments to cover

the surface of Mg alloys is basically used to remove the native oxide layer that is formed due

to high reactivity of Mg matrix leading to surface corrosion [101, 102]. The surface

morphology can be easily controlled at as low as nanometre levels. Moreover, chemical

methods can be manipulated easily and are generally cost effective without the requirement

for special facilities. The following are several methods of chemical modification.

2.9.1.1 Acid etching

Acid etching removes the native and porous oxide films through the creation of uniform and

compact layers, leading to the minimisation of local degradation behaviour. Acids such as

phosphoric acid (H3PO4) [17] and sulphuric acid (H2SO4) [103] have been used so far for this

purpose. A dense and homogenous layer of Mg3(PO4)2 coating was found on the surface of

the treated samples that showed much lower degradation rate as compared to the control [17,

103].

2.9.1.2 Alkaline treatment

Alkaline treatment of Mg substrates leads to formation of a thin layer of passive film on the

surface of the alloy during their immersion in cell culture medium, which inhibits the

increase of pH [104]. In addition to alkali, other alkaline solutions such as Na2HPO4, Na2CO3

and NaHCO3 [105] have been used to modify the Mg alloy surfaces. The treatment of Mg

alloys with alkali and alkaline solutions leads to the formation of a passive layer in the form

of Mg(OH)2, MgCO3, and MgO. The resistance to degradation depends on the thickness of

the film formed, which further depends on the surface roughness of the alloy. Hence, surface


29

topography of the alloy needs to be controlled during alkaline treatments for desirable

protective films.

2.9.1.3 Fluoride treatment

Fluorine is one of the most important components of skeleton and teeth in humans [106] and

hence its use can be beneficial to the skeletal system. Fluoride treatment leads to replacement

of the original oxide layer in Mg alloys, with a thin and homogeneous layer of magnesium

fluoride (MgF2) that shows higher corrosion resistance. This layer is found to have high

density, of lower solubility in aqueous solution and harmless to the host organism [107]. Also

used in Mg alloy coatings are MgF2, ammonium fluoride (NH4F) solution [103] and

hydrogen fluoride (HF) [108] to improve the corrosion resistance of pure Mg.

Apart from the chemical techniques, many other electrochemical methods were developed to

improve the corrosion property of Mg alloy such as anodising coating [109], electron beam

treatments [110] and ion implantation [111-113] that uses different applied high power

voltages to reduce the degradation rate of Mg in immersed state.

2.9.2 Physical modification

In the physical approach of surface modification, there is no chemical bond formation

between the outer coating and the surface of the substrate. However, it induces structure

modification of the entire scaffold by introducing new layers or phase such as polymers and

inorganic compounds [114-116] leading to better performance in the corrosion resistance of

the Mg scaffolds.


30

2.9.2.1 Apatite coating

Apatite is one of the main component of natural bone matrix (Section 2.3), which is used

mostly in the form of hydroxyapatite [HA, Ca10(PO4)6(OH)2]. Apatite can make a remarkable

impact on the growth of fractured bone due to its excellent mechanical and biological

properties like relative low solubility with high thermal stability [114] and cell adhesion with

osteoinduction [117] respectively. HA is highly bioactive and thus can enhance bone bonding

to the implant surface and prevent formation of fibrous tissue in and around the implant area

[118]. To date, HA has been used by itself or in combination with other chemicals like

fluorine doped HA [119, 120] to enhance the bioactivity of Mg alloys. However, HA has a

relative high dissolution rate and low adhesion, which make it unfavourable for long-term

stability of the implants.

2.9.2.2 Organic polymer coatings

Many polymers have been used as coatings for biomedical applications as degradable

polymers can insulate the solution and substrates more effectively compared with other

surface coatings. Polymers have been used in the form of PGLA [poly(lactide-co-glycolide)]

[115] and PCL (polycaprolactone) [99] coatings on Mg alloy implants to control their

bioactivity. However, pre-treatment steps are required to use these polymers as coatings so

that it can provide the ability of tight cell adhesion. For this reason, it is crucial to seek more

suitable polymers as protective layers for developing medical implants.

2.9.2.3 Metal oxide coatings

Metal oxides are good choices for modifying the degradation properties of Mg alloys as the

oxidative forms of some non-toxic metal elements can form a protective deposition layer on

the alloy surface. Metallic coatings like Titanium oxide (TiO2) films [121] and Cerium oxide


31

(CeO2) coatings [122, 123] have been reported to improve the biodegradation properties of

Mg alloys. The effects of coating showed a decrease in the corrosion results by about 30

times for the composite-coated samples when compared to the uncoated ones.

2.9.2.4 Morphological surface modifications

Morphological surface modifications are done to create three-dimensional features in the

form of pores, gratings, columns, dots, pits and random surface roughness [124, 125].

Morphological modification of Mg alloys can be done by mechanical and heat treatment like

deep rolling, shot peening and roller burnishing [126] and involves physical treatment,

shaping or removal of the material’s surface. The main aim of these kinds of mechanical

modification is to attain specific surface topographies and roughness, remove surface

contaminants and to improve cell adhesion.

2.9.2.5 Bioactive surface modifications

Bioactive and functional surface modification of Mg alloys is done to make them suitable to

react with ions and molecules and to activate/promote osteoconduction and osseointegration.

In addition to the above chemical and physical surface coating methods, several other coating

procedures also include immobilisation techniques that can reduce cytotoxicity of the alloy

and enhance their cytocompatibility. Many authors have reported several coating agents in

the form of proteins and peptides that are bioactive and can influence osseointegration.

Among them, Collagen type I (Col-I) [127, 128], Bone Morphogenetic Protein-2 (BMP-2)

[129, 130] and calcium phosphate [131, 132] are the most used agents for coating. Col-I,

extracted from rat-tail, has been used as coating agent in my work. The description of this

protein is given below.


32

Col-I is the predominant component of connective tissue matrix. It is one of the major

constituents of bone ECM comprising 90% of the protein matrix of adult bone, and serves as

a basis for mineral scaffold [133]. Col-I has three polypeptide chains (α chains) that form a

unique triple-helical structure. It is a heterotrimer of two α1 and one α2 chains, α1(I) is more

conserved than the α(I) chain. It contains an uninterrupted triple helix of approximately 300

nm in length and 1.5 nm in diameter flanked by short non-helical telopeptides. Each chain

has a repeating Gly-Xaa-Yaa sequence, in which Xaa and Yaa can be any amino acid but are

frequently the amino acids proline and hydroxyproline. It is found predominantly in skin,

bone, tendon, ligaments and cornea.

Figure 2.9. Figure illustrating the steps of Col-I fibril formation [134]. Col-I affects the

attachment and differentiation of osteoblast cells by binding to integrins [135-138]. Earlier

studies illustrated that bone cells adhere to Col-I because it contains RGD sequence that are

recognised by integrins. However more recent studies on this protein showed that its α1 chain

contains DGEA motif which is recognised by integrin α2β1 [136, 139].


33

The telopeptide does not contain the Gly-Xaa-Yaa conserved sequence and does not take the

form of triple helix although it is very important in fibril formation. Col-I self-assemble into

fibrils [134] and the process is driven by loss of solvent molecules from the protein surface.

This assembly of monomers into fibrils minimises the surface area/volume ratio. The fibril

formation occurs after the soluble procollagens are converted into collagens by specific

enzymatic cleavage of terminal propeptides by the procollagen metalloproteinases (Figure

2.9).

2.10 Summary

The degradation property of Mg-based alloy has become the most important aspect with

regard to its success, because of the many advantages specified in this review. Ideally the

coating itself should degrade gradually, helping to control the overall corrosion process,

while leaving no harmful traces of the corroded products. Further, no clear evidence exists

regarding the osseointegration and enhanced cellular response to surface modified alloys of

Mg containing different elemental constitutions of Zr, Ca and Sr. It is important to understand

the specific role of elemental constitutions of different concentrations of alloyed elements in

Mg alloys and their cellular response when coated with Col-I. This forms the basis of the

objective of this study.

Chapter 3. Materials and Experimental Techniques

34

Chapter 3

Materials and Experimental Techniques

3.1 Sample Preparation

The Mg-Zr-Sr/Ca alloys were prepared by casting from melts of pure Mg (99.9%), Ca

(99.9%) and master Mg-33Zr and Mg-30Sr alloys. The contamination of molten Mg by Fe

was avoided by using a coated steel crucible and the Mg alloys were melted under an

atmosphere of high purity argon. The melt was maintained by stirring for 30 min and was

casted into cylindrical steel dies at 700 °C after preheating to 250 °C, with an inner diameter

of 22 mm. This Mg alloy ingots were machined to remove the outer layer to obtain

cylindrical bars with diameters of 9 mm and 5 mm by electrical discharge machining (EDM).

Disc samples with a diameter of 9 mm and thickness of 2 mm were machined along the long

axis of the cylindrical bars for microstructure characterisation and in vitro cell culture study.

Cylindrical samples with a diameter of 5 mm and length of 10 mm were machined in the

same way for the compressive test. Similarly, cylindrical samples with a diameter of 2 mm

and length of 4 mm were cut along the long axis for their in vivo assessment. The chemical

compositions of the alloys were determined by wavelength dispersive X-Ray fluorescence

(WDXRF) spectroscopy (S4 Pioneer, Bruker, Germany).

3.2 Microstructure, Mechanical and Surface Properties

3.2.1 XRD and compressive test

X-Ray diffraction (XRD) analysis was carried out to characterise the microstructural phases

of the Mg alloys. To characterise the phase constituents of the alloys, XRD analysis was

performed at 40 kV and 30 mA with a scanning rate of 0.05 degree 2 theta per min. The scan


35

range was from 20 to 80 degrees. To evaluate the mechanical properties, compressive tests

were carried out at an initial strain rate of 10-3 s-1 using an Instron universal tester equipped

with a video extensometer (Instron 5567, USA).

3.2.2 Optical microstructure and surface morphology by SEM

All the Mg alloy samples for optical microscopic observation were etched with picric acid (a

solution of 1.5 g picric acid, 25 ml ethanol, 5 ml acetic acid and 110 ml distilled water). The

surface morphology of the Mg-based alloys was investigated using SEM (3400N Hitachi,

Japan) after gold coating in an ion sputter coater SC7620 (Polaron Range, UK).

3.2.3 Surface roughness and surface free energy

Surface roughness (Ra) measurement [140] of the bare and coated Mg-based alloys was done

using stylus profilometry (Taylor Hobson, London, UK). Five measurements were performed

for each group of samples to evaluate the average Ra values. The free surface energy of the

bare and coated alloys was estimated by measuring the static contact angle ( ) using 2 μL of

water and glycerol with a sessile drop method in a video-camera based OCA 40 micro,

DataPhysics (Germany). Average values of for left and right side of the drop were

measured by Wu approach [141] for each solvent at three different positions on each alloy

surface. This value was used for calculation of the free surface energy. Surface energies of

the substrates were calculated using SCA20_U software (provided by the manufacturers)

applying the Wu and Owens- Wendt method [141]. Average results from three independent

experiments were used for calculating the mean surface energy of each alloy.

3.2.4 Hydrogen gas evolution

The hydrogen gas evolution by Mg alloys was measured by following the protocol described

by Aung and Zhou (2010) [142]. The Mg alloys were immersed in simulated body fluid


36

(SBF) and the volume of hydrogen gas generated was measured as a function of immersion

time. Disc samples of 9 mm diameter and 2 mm thickness were used for the hydrogen

evolution tests. The Mg alloy samples were mounted in epoxy resin with an exposed area of

0.64 cm2 (9 mm diameter). The mounted sample with an exposed area upwards was

immersed into SBF and incubated at 37 °C. A funnel and burette were used to cover the

sample and collect the hydrogen gas produced [142]. The volume of hydrogen gas was

measured using the burette.

3.3 Collagen Type-I Extraction, Characterisation and Coating

3.3.1 Protein extraction, precipitation and estimation

The Col-I was extracted according to the protocol described by Rajan et al. [143]. In brief,

Col-I was extracted from rat-tail tendons in acetic acid solution to obtain a crude collagen. It

was then dissolved in 0.25 N acetic acid and then precipitated in 2.4 M NaCl to obtain a type-

I collagen. The dissolved collagen was then dialysed against 0.25 N acetic acid for 48 h at 4

°C, followed by lyophilisation. The samples were stored in -80 °C for further use. For

experimental use, samples were reconstituted in 0.25 N acetic acid and estimated by Sirius

Red method [144].

3.3.2 SDS-PAGE and Western blotting

The reconstituted Col-I was subjected to Sodium Dodecyl Sulfate Polyacrylamide Gel

Electrophoresis (SDS-PAGE) on 8% gel to visualise the Col-I bands. Standard Col-I was

obtained from CSIRO (Monash University campus, Australia). The polyacrylamide gel

contained 1.5M (pH 8.8) and 1M (pH 6.8) Urea (Bio-Rad, Hercules, California, US) for

resolving and stacking gel, respectively, with 30% Acrylamide (Bio-Rad), 10% Ammonium

Persulfate (APS) (Bio-Rad) and 0.2% TEMED (Sigma Chemicals, St. Louis, Missouri, US).


37

Proteins were visualised by Coomassie Brilliant Blue R-250 staining (Bio-Rad). Images were

acquired using Fluor- S Multi Imager (Bio-Rad). The polyacrylamide gel was electroblotted

onto PVDF membrane and blocked overnight in 5% Bovine Serum Albumin (BSA) (Sigma

Aldrich, St. Louis, MO, USA) in Tris-buffered saline (TBS) containing 0.1% Tween- 20.

Membranes were incubated with primary antibody (monoclonal IgG1 anti-bovine Col

antibody made in mice) (Sigma Aldrich, St. Louis, MO, USA) for 2 h, followed by

horseradish peroxidase (HRP) tagged secondary antibody (rabbit anti-mouse Ig) (Sigma

Aldrich, St. Louis, MO, USA) in TBS containing 0.1% Tween-20 for 1 h. Bound antibodies

were visualised using a chemi-luminescence kit (Roche Applied Science Inc., Penzberg,

Germany).

3.3.3 Dip coating

The Mg alloys were dip coated according to the procedure described by Geißler et al. [127].

The lyophilised Col-I of concentration 5 μg/μL was dissolved in 0.25 N acetic acid with

equal volume of 133 mM phosphate buffer (pH 6.5). The solution was then centrifuged at

5000 X g at 4 °C for 5 min. The alloys were then coated with Col-I solution using a Single

Dip Coater (Model No SDC 2007C, Apex Instruments Co., Kolkata, India) using SDC

Software provided by the instrument suppliers. The alloys were dipped with dipping/lifting

speeds of 60 mm/min for a waiting time of 5 min. Samples were then dried at room

temperature.

3.3.4 Characterisation of protein and coated surface by FT-IR

The standard Col-I and Col-I coated alloys were examined for transmission with an FT-IR

spectrophotometer (Vexter 70, Bruker, Germany) using the attenuated total reflectance

(ATR) method. For spectral acquisition, the system was used in point mode with a 4.0 cm-1


38

resolution at 32 scans in ATR spectrum mode. Spectra were analysed using the OPUS 5.5

software provided by the instrument suppliers.

3.4 Cell Culture

3.4.1 Cell line maintenance

Mouse osteoblast MC3T3-E1 cell line of subclone 4 from ATCC (American Type Culture

Cultivation, USA) was used for the culture experiments. The cells were maintained in α-

MEM (Invitrogen, Gibco, USA) supplemented with 10% fetal calf serum, 100 U mL-1

penicillin and 100 μg mL-1 streptomycin at 37 °C in a humidified atmosphere of 95% air and

5% CO2. The cells were used within 4 to 5 passages after revival from frozen condition

where they were maintained in original condition.

3.4.2 MTT assay for cell viability

The biocompatibility of bare and coated Mg alloys was evaluated using the indirect contact

method [145] according to ISO 10993-5 [146].

3.4.2.1 Preparation of alloy extract

Alloy extracts were prepared in alpha-MEM complete medium supplemented with 10% fetal

calf serum (FCS), 100 U mL-1 penicillin and 100 μg mL-1 streptomycin at 37 °C in a

humidified atmosphere of 95% air and 5% CO2 for 24 h with surface area of alloy to media

ratio of 0.6 cm2 mL-1. The supernatant media extract from each sample was collected and

filter-sterilised with a 0.22 μ filter (Merck Millipore, Darmstadt, Germany). Alpha-MEM

complete medium was used as negative control while alpha-MEM complete medium with

0.64% phenol was used as a positive control by incubating them under similar conditions.


39

3.4.2.2 Cell culture

Cell culture was done by suspending 5 X 103 cells in 100 μL complete media on a 96-welled

tissue culture plate (TPP, Switzerland). Cells were incubated for 30 min at 37 °C in a

humidified atmosphere of 95% air and 5% CO2. After which additional 2 mL media was

added to each well and incubated for 4 h. The media was then removed and the discs were

washed twice with PBS to remove the non-adherent cells. The percentage of viable cells on

the culture plate was calculated by performing MTT assay.

3.4.2.3 MTT assay

After incubation for 4 h, the culture media was replaced with 100 μL of respective alloy and

control extracts. The cells were further incubated for 12 h after which MTT (3- (4, 5-Di-

methylthiazol-2-yl) -2, 5-diphenyltetrazolium bromide) (Sigma) assay [147] was performed.

The experiments were carried out in triplicates. A standard curve was generated using

different cell numbers. The optical density (OD) of all the samples and the control were

converted to respective cell numbers and the relative cell viability for each sample was

calculated as:

Relative Cell Viability = (ODtest/ODnegative control) X 100%

3.5 Techniques for Gene Profiling

3.5.1 Cell culture for gene expression studies

Due to early degradable nature of Mg alloys under in vitro conditions, the gene expression

studies were performed by indirect method of cell culture. Cells were plated using complete

media as described above in a tissue culture T25 flask. After allowing initial adhesion of 4 h,

the complete media was replaced very gently by alloy extracts without disturbing the cells.

They were then incubated in standard conditions for 24 h (1d) and 7 days for gene expression


40

studies. The alloy extract was supplemented with conditioned media at an interval of 2 days

for 7 days culture to maintain the media volume.

3.5.2 Total RNA isolation

High quality total RNA was extracted and purified from the cells cultured with each alloy

extract different time points of 24 h and 7 days using the Qiagen RNA plus kit (QIAGEN,

USA) as per the manufacturer’s protocol. In brief, culture flasks were washed with PBS with

subsequent addition of 600 μL of Buffer RLT Plus (with β-mercapto ethanol). The culture

flasks were scrapped with a rubber policeman and the lysate was passed 5 times through a 20-

gauge needle. The lysate was then passed through a gDNA Eliminator spin column. This

column, in combination with the optimised high-salt buffer, allows efficient removal of

genomic DNA. Ethanol was added to the flow-through to provide appropriate binding

conditions for RNA, and the sample was then applied to an RNeasy spin column where total

RNA binds to the membrane and the contaminants are efficiently washed away with Buffer

RW1 and RPE. High quality RNA was then diluted in 50 μL of diethyl pyrocarbonate

(DEPC) treated ddH2O. RNA from two separate experiments was isolated for preparing

cDNA.

3.5.3 RNA quality check and quantification

The nucleic acid concentration of the RNA preparation was determined by measuring the

absorbance at 260 nm using a Nanodrop spectrophotometer ND1000 (Thermo Scientific,

USA). 50 μg of double stranded DNA, 33 μg of single stranded DNA (primers for PCR) and

40 μg of single stranded RNA was taken to be equal to 1.0 OD260. Purity was checked by

taking the ratio of the absorbance at 260 nm to 280 nm. The quantified RNA was checked on

0.8% agarose gel electrophoresis for its purity.


41

3.5.4 cDNA synthesis/reverse –transcription reaction

cDNA synthesis was done using superscript IIITM reverse transcriptase (RT) enzyme

(Invitrogen Life Sciences, USA). 2 μg RNA for each alloy type and for each time point was

added to a reaction mixture containing 10 mM deoxyribonucleotide triphosphate mix (dNTP)

and 50 μM oligodT (Oligo dinucleotide) for first strand synthesis at 65 °C for 5 min after

which it was instantly placed on ice for approximately 5 min. A master mix consisting 5X

reaction buffer, 0.1 M DTT (Di-thio-threitol), RNase inhibitor and superscript IIITM RT

enzyme was prepared and added to the first strand synthesised mixture by incubating at 50 °C

for 1 h followed by inactivation of RT at 70 °C for 15 min. cDNA synthesis was done for

both the RNA preparations and used for real time PCR experiments. DEPC-treated water was

used for the experiments.

3.5.5 Quantitative real time polymerase chain reaction (RT-PCR)

A 10 uL reaction mixture each consisting of triplicate samples of cDNA, specific primer mix

and Power SYBR green Master mix (ABI systems, USA) was setup in each well of a 384-

well reaction plate (ABI Systems, USA). Forward and reverse primers, specific for genes,

were adapted from Sista et al. (2013) [148]. cDNA for GAPDH was used as endogenous

control for calculating fold differences in RNA levels of cells cultured in the alloy extracts by

the 2- CT method. Table 3.1 shows the list of forward and reverse primers for genes used.

After sealing the reaction with optical adhesive cover (ABI Systems, USA) the PCR plate

was placed in Applied Biosystems 7900HT Fast real time PCR system machine (Applied

Biosystems, USA). The cycle conditions were set up as detailed: 50 °C for initial heating of 2

min, 95 °C for 1 min, 40 cycles of 95 °C for 30 s followed by 60 °C for 30 s and 72 °C

elongation for 30 s each. A dissolution curve analysis was done at the end of every reaction to


42

Table 3.1. Forward and reverse primers of genes used for quantitative RT-PCR experiments.

check for the specificity of the final product formed. The entire quantitative real time PCR

was done for duplicate RNA samples for calculating the statistical significance of the data.

3.5.6 Agarose gel electrophoresis

Agarose gel electrophoresis is used for checking the purity of isolated nucleic acids. In this

study, the purity of the isolated RNA was checked using 0.8% agarose gel prepared in TBE

Gene Primers (5’-3’) Product length (bp)

GAPDH Fwd- AGC GAG ACC CCA CTA ACA TCA 21

Rev– CTT TTG GCT CCA CCC TTC AAG T 22

ALP Fwd– ACC CGG CTG GAG ATG GAC AAA T 22

Rev– TTC ACG CCA CAC AAG TAG GCA 21

Collagen type-I Fwd– CTC CTG ACG CAT GGC CAA GAA 21

Rev– TCA AGC ATA CCT CGG GTT TCC A 22

Osteopontin Fwd– TGA TTC TGG CAG CTC AGA GGA 21

Rev– CAT TCT GTG GCG CAA GGA GAT T 22

Osteonectin Fwd– ATG TCC TGG TCA CCT TGT ACG A 22

Rev– TCC AGG CGC TTC TCA TTC TCA T 22

Bone sialoprotein Fwd– ACC GGC CAC GCT ACT TTC TTT A 22

Rev– GGA ACT ATC GCC GTC TCC ATT T 22

Osteocalcin Fwd– AGC AGG AGG GCA ATA AGG TAG T 22

Rev– TCG TCA CAA GCA GGG TTA AGC 21

BMP2 Fwd– TGG AAT GAC TGG ATC GTG GCA C 22

Rev– ACA TGG AGA TTG CGC TGA GCT C 22

Fibronectin Fwd– TGC AGT GGC TGA AGT CGC AAG G 22

Rev– GGG CTC CCC GTT TGA ATT GCC A 22


43

buffer (Tris boric acid). Ethidium bromide (EtBr) was added just before solidification of the

gel after boiling the agarose solution for ~10 mins. The solution was then poured into a

casting tray with multi-welled comb and allowed to solidify at room temperature. The

samples were then loaded into each well mixed with 5X loading dye consisting of

bromophenol blue (BPP), xylene cyanol and β-mercaptoethanol. EtBr is a chelating agent,

which chelates between the bases present in the double stranded helix of DNA and RNA. The

gel was run at a constant voltage of 100V for 1 h. The RNA bands were visualised by

exposing the gel to UV lamp. Similar procedure was used for checking the specificity of the

final product formed after RT-PCR where a 1% agarose gel was prepared.

3.6 Animal Experiments

3.6.1 Implant preparation and surgery

Prior to surgery all implants were sonicated in 2% chromium oxide for 20-30 min followed

by 10 min wash each with acetone, ethanol and distilled water respectively, air-dried and UV

sterilized. Male New Zealand White rabbits of approximate age 4–6 months were used for in

vivo studies. Three animals were used for each alloy and as controls in each experiment.

Biologically duplicate experiments were carried out for statistical analysis. Alloys were

implanted in the hole drilled in the lower end of the femur bone of rabbits as per the surgical

procedures cleared by the institutional animal ethics committee of Deakin University and

CSIR-CCMB.

3.6.2 Blood sample collection, blood cell count and blood biochemical assays

Blood samples were collected from all the control and experimental animals, by retro-orbital

puncture, before surgery, 1 month and 3 months post implantation. For blood cell count,

blood samples were collected in K3-EDTA tubes and routine analysis of blood samples was


44

done in a Tulip CounCell23 automatic blood analyser (The Tulip Group, India). Serum

samples were obtained by centrifuging the collected blood at 3,000 RPM for 10 mins at 4ºC

and stored in -80ºC until use. Serum biochemical tests were done in Tulip CORALAB-3000

serum analyser (The Tulip Group, India).

3.6.3 Radiology and DXA

X-Ray Radiography was performed to measure the Bone Mineral Content (BMC) and Bone

Mineral Density (BMD), around the implant area, of live experimental and control animals,

1, 2 and 3 months after implantation by Dual Energy X-Ray Absorbtiometry (DXA) in a

Hologic Discovery QDR® Series (Bedford, MA, USA) to follow up the healing process after

the surgery and before removal of the implanted femur. The mineral concentration and

density of the region was calculated using the QDR software provided by the instrument

suppliers.

3.6.4 Anaesthesia, surgery and euthanasia

Animal surgery was performed in an operating room. Rabbits were anesthetised by

intramuscular local administration of a combination of ketamine hydrochloride (150 mg/kg

body weight) and xylazine hydrochloride (30 mg/kg body weight) (Troy Laboratories,

Glendenning, NSW, Australia) in a volume of 0.5 ml/kg body weight, prior to surgery. The

hair in the thigh region was removed and the surgery area, located at the tibial end of the left

femur, was disinfected with Povidone iodine solution. After exposing the dorsal side at the

distal end of the left femur bone (about 1 cm above the knee joint), a small cavity of

approximately 2.5 mm diameter and 4.2 mm depth was drilled through the cortical region

using a paediatric bone driller (Asian Surgicals, Hyderabad, India). A sterile implant was

placed in the cavity and the implant was covered with the adjoining muscle tissue, which was

sutured using absorbable suture material (Ethicon, Somerville, NJ, USA), and the skin layer


45

surrounding the muscular region was sutured by interrupted suturing with non-absorbable silk

thread and a 3–0 needle (Ethicon, Somerville, NJ, USA). In the control group, no implant

piece was placed into the drilled cavity and the drilled hole was left bare. After surgery the

surgical area was applied with Betadine ointment (5%, w/w Providone iodine, Win-Medicare,

New Delhi, India) and 0.5 ml Meloxicam (Intas Pharma, Ahmedabad, India) was injected

subcutaneously to prevent infection and inflammation in the surgical area; Lorexane ointment

was applied locally until the wound healed. All animals were kept under close observation

until the end of the experimental period during which ad libitum access to food and water

(sterile) without antibiotics was available to the animals. They were also allowed to move

freely within the cage without any support. All the animals were kept under close observation

for 3 months, during which time no mortality or morbidity in the animals was observed. After

the period of observation for 1 and 3 months, the animals were euthanised using carbon

dioxide inhalation for biochemical, radiological, and histological analyses.

3.7 Bone Processing and Sectioning

Immediately after excision each bone specimen (approximately 1 cm3) was fixed in 35 ml of

10% neutral buffered formalin (NBF) prepared in phosphate buffered saline (PBS) (pH 7.2),

for a period of 48 h at 4 °C under constant stirring. After fixation, the specimens were

dehydrated in increasing grades of isopropyl alcohol (IPA) (70%, 80%, 95%, and 100%),

each for 24 h and finally cleared with two cycles of xylene (100%) treatment, each for 24 h.

The samples were then processed for embedding.

3.7.1 MMA embedding

The embedding of bone samples was carried out in two stages: (a) infiltration and (b)

polymerisation. Each NBF-fixed bone sample was infiltrated by successive treatments with


46

35 mL solutions of three types of methyl methacrylate (MMA) (Merck, Fairlawn, NJ, USA) –

MMA I (60% MMA, 35% butyl-methyl-methacrylate (BMA) and 5% methyl benzoate mixed

with 1.2 ml of polyethylene glycol 400), MMA II (MMA I + 0.4% (w/v) benzoyl peroxide)

and MMA III (MMA I + 0.8% (w/v) benzoyl peroxide). Each infiltration step was carried out

at 4 °C for 3 days under constant stirring for efficient infiltration. The bone samples were

then placed in 5 ml glass vials (Borosil Glass Works, Ahmedabad, India) containing a 3 ml

bed that had been previously prepared by mixing MMA III + 0.6% (v/v) N, N dimethyl-p-

toluidine and polymerised at 4 °C for 24 h in the presence of nitrogen gas. The bone

specimens were oriented so that the implant surface was at right angle to the axis of the

implant in order to get sagittal sections of the implant area. The embedding polymer (MMA

III + 0.4% N, N dimethyl-p-toluidine, v/v) was poured on top of the bone specimen up to the

neck of the vial and capped immediately without any exposure to air. Complete

polymerisation of the resin was carried out in the absence of UV exposure by carefully

placing the glass vials in a −20 °C freezer (LG, South Korea) for 48–72 h without disturbing

the orientation of the specimen. The low temperature polymerisation reduces the heat

generated by the polymerisation reaction. Finally the polymerised blocks with the bone

specimens were removed from the vials by breaking the glass and stored at room temperature

till further use.

3.7.2 Bone sectioning

The embedded bone sections of 10 μm thickness were prepared using a rotatory saw

microtome (Leica SP1600, Leica Biosystems, Wetzlar, Germany). For classical stains, the

sections were mounted on glass slides (Blue Star, Polar Industrial Corporation, Mumbai,

India) freshly coated with Mayer’s egg albumin (egg white:glycerol; 1:1) freshly diluted in

order to achieve good adherence of the sections to the slides. For histochemical and


47

immunohistochemical analysis the sections were mounted on a freshly prepared 0.1% (w/v)

poly L-lysine (diluted in distilled water) (Sigma–Aldrich, St. Louis, MO, USA) coated glass

slides. Sections collected on poly-L-lysine coated slides were kept at 37 °C for 1 h to

facilitate good adherence of the sections to the slides. The mounted sections were stored at

room temperature until further use.

3.7.3 Deplasticisation

Deplasticisation of the bone sections were carried out using 2-methoxy-ethyl-actetate

(Sigma–Aldrich, St. Louis, MO, USA) three times for 20 min each and finally washed in

distilled water (DW). Deplasticised sections were treated with different staining protocols to

visualise the cells and ECM of the newly formed tissue.

3.8 Histology

3.8.1 Hematoxylin and Eosin

After deplasticizsation, sections on albumin-coated slides were washed twice in deionised

water each for 10 min, and then stained with hematoxylin for 6 min. This was followed by

washing with tap water for 10 min or until the tissue colour turned bluish followed by eosin

staining for 45 s. Stained sections were dehydrated using 95% IPA three times for 2 min each

and with 100% IPA for 2 min. The sections were mounted using Permount® (Fisher

Scientific, Waltham, MA, USA).

3.8.2 Tartrate Resistant Acid Phosphatase (TRAP) assay

Sections on poly-L-lysine coated slides were incubated in acetate buffer (pH 5.0) for 20 min

at room temperature and then in acid phosphatase–substrate solution containing naphthol AS-

MX phosphate (Sigma–Aldrich, St. Louis, MO, USA) and Fast Red (Sigma–Aldrich)


48

dissolved in acetate buffer at 37 °C for 1–4 h. Incubation time for every slide was optimised

by checking the bright red coloration of osteoclasts. The stained sections were washed in

deionised water and mounted in Kaiser’s glycerol gelatine (Merck, Darmstadt, Germany).

3.8.3 Masson Trichrome

Staining was done on sections attached to albumin-coated slides. The sections were stained

with Weigert’s hematoxylin for 6 min followed by washing in 1% acid–alcohol (7–8 dips)

followed by running tap water wash for 10 min. The sections were then stained with Biebrich

scarlet–acid fuschin solution for 2 min, rinsed in three changes of deionised water and placed

in 5% phosphotungstic acid for 15 min followed by two changes of deionised water. Finally,

the sections were stained in light green solution for 7 min, rinsed twice with deionised water

and treated with 1% acetic acid solution for 3–5 min. The stained sections were dehydrated

using 95% and 100% IPA and two changes of xylene each for 5 min and mounted with

Permount® (Fisher Scientific, Waltham, MA, USA).

3.8.4 Sirius Red Staining

After deplasticisation, the bone sections were placed in 2% KOH for a period of 30 min for

maceration. The sections were then transferred to 1.5% Alizarin Red solution (pH 7.4) until

the sections turned red under the microscope. Excess stain was removed by 1%-2% KOH.

The sections were washed in DW and mounted using Kaiser’s glycerol gelatin (Merck,

Darmstadt, Germany).


49

3.9 Immuno-Histochemistry

3.9.1 Collagen type-I and Vimentin

Col-I immunostaining was performed to study the ECM distribution laid by osteoblasts in the

peri-implant site. The sections were placed on poly-l-lysine coated slides, and deplasticised.

Sections were then rehydrated by washing twice in Tris buffered saline (TBS) for 5 min. This

was followed by antigen retrieval step, using proteinase K solution, at 37°C for 20–30 min.

Blocking of endogenous peroxidase activity was done by incubating them in 3% hydrogen

peroxide for 30 min at room temperature prior to staining procedure. After this, the sections

were washed twice in TBS and incubated in 3% bovine serum albumin (BSA) for 30 min to

prevent non-specific binding of antibodies. For antibody staining, the sections were divided

into two groups- (a) one group was used for incubation with primary anti-collagen I and anti-

vimentin antibodies; (b) the other was used as a control, where no incubation of first antibody

was done, but the sections were immersed in TBS. Sections in the first group were incubated

with primary antibody for monoclonal IgG1 anti-Col antibody (Sigma–Aldrich, St. Louis,

MO, USA) and monoclonal IgG1 anti-vimentin antibody made in mice (Sigma–Aldrich, St.

Louis, MO, USA) at 1:500 dilutions in room temperature for 45 min and 1 h, respectively,

and washed three times in TBS for 5 min each. Sections from both groups (experimental and

control) were further incubated with horseradish peroxidase (HRP) conjugated goat anti-

mouse antibody (Sigma–Aldrich, St. Louis, MO, USA) for 30 min at room temperature and

then washed in TBS three times each for 5 min. 3, 3-Diamino-benzidine (DAB) (Roche

Diagnostics, Indianapolis, IN, USA) was used as the substrate (according to the

manufacturers’ instructions) for visualisation of peroxidase enzyme activity on the sections.

Incubation reaction was done until the section turned brown in colour. The sections were

carefully observed under the microscope to prevent over-staining, washed in DW and

mounted using


50

3.10 Microscopy

Images of the stained sections were taken using Zeiss Axioplan 200M microscope (Carl

Zeiss, Oberkochen, Germany) fitted with an Axiocam Mrc digital camera and the images

were processed via the Axiovision® Version 5 software provided by the instrument supplier.

Chapter 4. Effect of Zirconium, Calcium and Strontium Contents on Implant-induced Osseointegration

51

Chapter 4

Effect of Zirconium, Calcium and Strontium Contents on Implant-Induced osseointegration

4.1 Introduction

Because of their varied attractive properties, many metals and their alloys have been

considered as biomedical implants. The idea of biodegradable alloys initiated with the

necessity for secondary surgery for removal of the implanted material and their unsuitability

in load bearing applications. A number of Mg based alloys, prepared with alloying elements

from the same group of the periodic table, have been reported for dental replacements and

orthopaedic applications [8, 13, 14]. Mg enriched materials have been studied as oral

implants [149] and as filler materials in extracted sockets [150] in animal models to evaluate

their influence in restoration or replacement of dental abutments. The application of alloys

and implant materials, placed into the extracted sockets immediately after tooth extraction

have reported beneficial effects [151, 152]. However, such techniques could not demonstrate

their capability of maintaining the bony crest in its original shape for long periods of time

because of their high corrosion and degradation in vivo. The biological and corrosion

properties of Mg alloys have been studied in detail [153-156], and their application as

orthopaedic implants has been widely accepted due to (a) their characteristic biodegradability

and biocompatibility in vivo [153-155], and (b) their established role in bone formation e.g.

ability to influence mineral metabolism in the bone matrix and promotion of osteoblast

specific cell signalling in vivo, without causing inflammatory reactions in the neighbouring

tissues [36]. In addition, mechanical properties such as elastic modulus and compressive yield

strength of many Mg based alloys closely match with that of natural bone tissue [157].


52

In spite of the above mentioned advantages, Mg-based alloys present some significant

challenges in their usage as bio-implants, for example (a) many Mg containing implants

corrode quickly at the physiological pH range of 7.4-7.6 [8, 158], and (b) they release

hydrogen gas around the implant area leading to loss of mechanical integrity even before the

tissue is healed and the new bone is mineralised [159, 160]. The performance of Mg based

alloys is also affected because of the instability of the protective hydroxide film on their

surface that dissolves in aqueous environments [161]. Using appropriate combinations of

alloying elements with the base alloy can significantly reduce these defects of Mg based

alloys. An ideal alloying element composition would stabilise the hydroxide film on the

surface, increase their corrosion resistance and mechanical properties and thus improve their

biocompatibility and bio-efficacy. In this direction, alloys of Mg with Sr, rare earth elements,

Ca, aluminium (Al), trace levels of manganese (Mn), zinc (Zn), Zr, and silicon (Si) have been

used to make new generations of orthopaedic implants [162-169]. An important factor while

designing these bioactive implant surfaces is the biological response of cells to the alloys

which ultimately correlates with the success of implants in the host tissue. The idealised

biological efficacy of an implant would be where the implant material gets totally

amalgamated with the newly formed osseous tissue and thereafter it disintegrates into the

blood stream without causing damage to the vital organs or losing its functionality [170].

The main aim of the present study was to check the stability and in vivo cellular response to

Mg alloyed with Zr, Sr and Ca and to evaluate the influence of these divalent cations on the

in vivo compatibility of these alloys. We hypothesise that inclusion of Sr in Mg alloys can

change their surface properties and can control the potential of these alloys to interact with

the osteoblast cells. This can lead to osteogenesis, osteoinduction and osseointegration

leading to ECM formation around them upon implantation. The addition of Zr, Sr and Ca to

Mg can influence their degradability and biocompatibility in different ways. Taking these


53

facts into consideration, we report the effect of Mg alloys on the in vivo degradation, bone

forming ability and mineralisation in addition to the impact of these alloys on the BMC and

BMD of the implanted animals.

4.2 Results

4.2.1 Mechanical and surface properties

4.2.1.1 Alloy composition and physicochemical characterisation

The detailed alloy compositions of the four substrates are given in Table 4.1.

As can be seen the substrates were almost free of contaminants such as Al, Fe, Mn and Si and

the relative contents of base (Mg) and alloying (Zr, Sr and Ca) metals were as per the desired

values in every substrate. The XRD data (Fig 4.1) showed that all alloys contained a

dominant and primary α-Mg phase. In addition to this, in Mg-2Zr-5Sr and Mg-Zr-2Sr alloys,

a strong and intermediate peak for the Mg17Sr2 intermetallic phase could be seen; this peak

was absent in the other two alloys. Intermetallic Mg-Zr peaks were absent in alloys – this

could be attributed to their low solubility in the presence of Mg and distribution of α-Zr

particles in Mg matrix, as has been explained by others [172]. Almost all the alloys exhibited

good mechanical strength that matched the properties of normal bone tissue, making these

Alloy Chemical composition (wt%)

Zr Sr Ca Al Si Fe Mn Mg

Mg-5Zr 4.88 - - 0.02 0.01 0.01 0.01 Balance

Mg-5Zr-Ca 4.41 - 0.91 - - 0.01 0.01 Balance

Mg-2Zr-5Sr 1.89 4.8 - 0.03 0.02 0.01 0.01 Balance

Mg-Zr-2Sr 0.92 1.82 - 0.04 0.02 0.02 0.01 Balance

Table 4.1. Nominal composition of Mg-based alloys [171].


54

alloys suitable and potential candidates for temporary implant applications. The compressive

strengths of Mg-5Zr, Mg-5Zr-Ca, Mg-2Zr-5Sr and Mg-Zr-2Sr alloys were estimated to be

237.8, 255.7, 209.7 and 242.9 MPa respectively, which was within the range of the

compressive strength of normal cortical bone (164-240 MPa) [158], except Mg-5Zr-Ca. As

can be seen, addition of Ca to the Mg-Zr base alloy in Mg-5Zr-Ca, made it stiffer and more

inflexible than the other three alloys. Based upon the compressive strengths and the

composition of the alloys we decided to compare the bone inducing activity of Mg-5Zr-Ca

vs. Mg-5Zr implants and Mg-2Zr-5Sr vs. Mg-Zr-2Sr implants. The first grouping was to

evaluate the role of Ca and the second for evaluating the combined roles of Zr and Sr in the

alloys.

Figure 4.1. X-ray diffraction spectra of four Mg alloys.


55

4.2.1.2 Surface morphology and in vitro hydrogen gas release

The SEM images of the alloys before incubation (Fig 4.2) in the culture media showed rough

surfaces, with scattered distribution of particles having a floral arrangement. After incubation

in culture media (Fig 4.3), we observed the appearance of cracks, pits and more roughened

edges on all surfaces that could be biologically advantageous for the implants because such

surface topographies can accelerate the osteoblast adhesion [172]. None of the alloys showed

any indication of peri-implant gas bubble formation or its shadows up to three months. It is

however possible that small amount of H2 gas was generated around the implant area within

the first few days after implantation. This release of H2 gas was considered within tolerable

limits and therefore H2 evolution from the implants was of little concern in our study.

Figure 4.2. Scanning electron micrographs of Mg alloys before incubation in the culture media

[171].


56

4.2.1.3 Surface energy

Table 4.2, fig 4.4 and fig 4.5 respectively shows the calculated free surface energy values and

shapes of individual drops, and the left and right contact angles of all the alloys. Mean

contact angles of the alloys were between of 23°-75° and the corresponding surface energies

were in the range of 40-70 μJ/mm2 as shown in Table 4.2. Based on this we concluded that

the surface energies of the alloys were in the following order - Mg-5Zr-Ca>Mg-5Zr>Mg-2Zr-

5Sr>Mg-Zr-2Sr. It is known that substrates with low surface energies are more hydrophobic

than surfaces with high surface energies [174]. Therefore in our study, since Mg-Zr-2Sr

showed the least surface energy and Mg-5Zr-Ca the most, Mg-Zr-2Sr was considered as most

hydrophobic and Mg-5Zr-Ca as most hydrophilic.

Figure 4.3. Scanning electron micrographs of Mg alloys after incubation in the culture media

for 7 days [171].


57

Alloy Liquid Left CA Right CA

Mg-5Zr

Water 25.25° 0.92 27.25° 0.49

Glycerol 43.9° 1.27 46.65° 0.07

Surface Energy (SE) 66.76 μJ/mm2 66.53 μJ/mm2

Average SE 66.64 0.162 μJ/mm2

Mg-5Zr-Ca

Water 23.3° 0.71 24.3° 0.99

Glycerol 45.7° 0.57 48.1° 1.98



Mg-2Zr-5Sr

Water 43.45° 0.64 41.65° 2.33

Glycerol 53.65° 6.15 54.35° 3.04



Mg-Zr-2Sr

Water 63.85° 1.76 64.15° 1.91

Glycerol 74.4° 2.68 73.55° 0.64

Surface Energy (Se) 42.42 μJ/mm2 41.85 μJ/mm2


Table 4.2. Mean left and right contact angle (CA) values and the respective surface energies

(SE), with statistical standard deviations, of all the four Mg-based alloys.


58


surface energy. The contact angles (CA) were determined using water and measured within

60 seconds of the droplet release from the syringe probe.


surface energy. The contact angles (CA) were determined using glycerol and measured within

60 seconds of the droplet release from the syringe probe.


59

4.2.2 In vivo bone formation

4.2.2.1 Histological analysis of new bone formation

The histological study of the implant induced bone properties at 12 weeks after implantation

of the alloys gave us two-fold approach: on one hand the focus was on identifying the cell

types within the newly induced bone tissue by doing HE and TRAP staining and on the other

hand, features of the ECM of new bone by Col immunostaining and MT staining was

depicted. HE staining showed no significant features of local inflammatory responses around

the implants but the fibrous encapsulation of the Mg-5Zr and Mg-5Zr-Ca implant induced

new bone (NB) indicated inflammatory response towards these implants (Fig 4.6).

HE staining of the Mg-5Zr and Mg-5Zr-Ca induced bones also showed a non-uniform bone

formation pattern that was rich in fibroblast like cells and an ECM that was filled with

remnants of degraded alloy particles. These features of Mg-5Zr and Mg-5Zr-Ca implant

induced bones were indicative of “distance osteogenesis” behaviour. In contrast Mg-2Zr-5Sr

Figure 4.6. Hematoxylin and eosin stained bone sections (100×) of Mg alloys 3 months post-

implantation. The dots represent the implant–bone interface [171].


60

& Mg-Zr-2Sr implant induced bones showed firm adhesion of the new tissue to the implant

surface and a uniform osteoblast rich bone tissue, which represented the features of “contact

osteogenesis”.

TRAP staining of the peri-implant site showed a significantly higher number of osteoclasts in

the Mg-2Zr-5Sr and Mg-Zr-2Sr induced bones as compared to Mg-5Zr and Mg-5Zr-Ca (Fig

4.7), indicating more appropriate bone remodelling behaviour in the former. We could also

notice the complete degradation of the Mg-2Zr-5Sr and Mg-Zr-2Sr implant materials, which

could be responsible for the activation of osteoclasts. Between the Mg-2Zr-5Sr and Mg-Zr-

2Sr implants the total number of TRAP-positive cells was more in the Mg-2Zr-5Sr induced

bones than in Mg-Zr-2Sr induced bones. The poor degradation of the Mg-5Zr and Mg-5Zr-Ca

implants even after 12 weeks indicated their lesser osteoclast inducing capacity.

Figure 4.7. Tartrate-resistant acid phosphatase staining of bone sections (100×) 3 months

post-implantation. The red coloured cells represent osteoclasts at the peri-implant interface.

The dots represent the implant–bone interface. Imp, implant site; OC, osteoclasts [171].


61

4.2.2.2 Immuno-histochemical analysis of newly formed bone

A positive Col immunostaining could be seen in the ECMs of all the new induced peri-

implant bones but its levels were higher in Mg-2Zr-5Sr and Mg-Zr-2Sr induced bones (Fig

4.8). Collagen type-I is one of the most abundant proteins of the mineralised bone matrix and

it is a marker of early osteoblast differentiation [138]. It is synthesised in the osteoblasts and

its secretion in the matrix indicates formation of an osteoid matrix.

MT staining (Fig 4.9) was used to differentiate the unmineralised osteoids (red staining) from

the mineralised ECM (blue staining) of mature bone tissue in the implant area. The bone

response around the implants Mg-5Zr and Mg-5Zr-Ca were separated by a gap interface rich

in fibrous tissue. The thickness of the fibrous tissue enclosing the implant Mg-5Zr is 135 μm

Figure 4.8. Immunostaining of implanted bone sections (100×) by col-I, 3 months post-

implantation. Panels show osteoblasts lining the newly formed bone. The osteoblasts

occupied significantly higher bone surface near the implant site in Mg-2Zr-5Sr and Mg-Zr-

2Sr. The dots represent the implant–bone interface [171].


62

and that of Mg-5Zr-Ca is 176 μm. The cellular details of the dense fibrous envelope formed

is well depicted in the MT stained sections where the well-structured collagen fibres are

stained blue-green in colour and the cell nuclei in red. The peri-implant site of these implants

showed superior bone contact after 3 months of healing period as depicted in Fig 4.9. These

images are depictive of mature woven bone that distinguishes between osteoblasts and

fibroblast mineralisation. Implant Mg-Zr-2Sr represented prominent osteogenic activity with

mature woven trabecular bone formation after 3 months with implant site filled with spongy

bone and no degradable particles illustrating uniform disintegration of the implant. The newly

formed bone was completely mineralised in Mg-Zr-2Sr while Mg-Zr-2Sr showed

unmineralised bone (osteoid) in red when stained with MT. The osteoblasts lining the

trabecular surface were pre-dominant around the peri-implant interface of the implants Mg-

2Zr-5Sr and Mg-Zr-2Sr as shown by MT stained images.

Figure 4.9. Masson’s Trichrome stained sections (100×) of magnesium alloys 3 months post-

implantation. The dots represent the implant bone interface [171].


63

4.2.2.3 Radiography and bone density measurements

The mean BMC and BMD values of the implant induced bones (three animals for each alloy)

and three control animals recorded at monthly intervals for three months are shown in Figs

4.10A and 4.10B, respectively. BMC values for control animals did not show any increase

with time whereas in all implant-containing bones, the values increased significantly by the

third month when for Mg-2Zr-5Sr and Mg-Zr-2Sr implants they were higher than Mg-5Zr

and Mg-5Zr-Ca. Interestingly, BMD values for all the implants remained more or less similar

to control values showing no significant change. This difference in the variation of BMC and

BMD values implies that after 3 months the total area (or volume) of healed bone in Mg-2Zr-

5Sr and Mg-Zr-2Sr implant bearing wounds was bigger than the healed areas of control or

Mg-Zr-2Sr implant bearing bones. Further, this data reflects the osteoinductive property of

the Mg alloys. Hydrogen evolution rates were calculated based on immersion tests performed

by incubating the substrates in SBF and Fig 4.10C shows the data for all the four alloys. As

expected, all the substrates exhibited release of hydrogen gas upon immersion in SBF, which

was in direct correlation with their immersion time. Mg-5Zr-Ca showed maximum hydrogen

evolution at 72 hrs indicating that it had corroded and dissolved entirely in SBF by that time.

In contrast, the Sr containing alloys showed a more optimized hydrogen evolution and

degradation behaviour in relation to those alloys that did not contain Sr. A representative X-

ray image of implanted area for the Mg-Zr-2Sr implant is shown in Fig 4.10D. We also

noticed that after three months of implantation, the implanted pellets of the alloy had

degraded and their edges had become too fuzzy to be recognised. The degradation levels of

the implants were variable from alloy to alloy which became more evident in the histological

analysis.


64

Figure 4.10. Physiological, physicochemical, and anatomical effects of the four Mg alloys

used for implantation. Panels A and B show BMC and BMD, respectively, of the implanted

animals after 1, 2, and 3 months post-implantation for control and experimental animals. Data

shown are mean values taken from three experimental animals in each time point. Panel C

shows hydrogen evolution rates of magnesium alloys upon incubation in simulated body fluid

for 7 days. Panel D shows X-ray radiograph of a representative implanted live animal (Mg-

Zr-2Sr) after 3 months implantation [171].


65

4.2.2.4 Blood cell counts and serum biochemical measurements

Total blood-cell count and serum biochemical analysis of the rabbits performed 3 months

after implantation gave the information on inflammation, liver and kidney functions. The

blood examination results are listed in Table 4.3. The rabbits implanted with Mg-5Zr and

Mg-2Zr-5Sr showed slight increase in lymphocytes (LYM), hematocrit (HCT), mean

corpuscular volume (MCV) and mean cell hemoglobin concentration (MCHC). An increase

in platelet (PLT) count was found in Mg-Zr-2Sr. However, these elevations do not indicate

any kind of disturbances in the blood count. All the animals were in good physical condition

and free from any kind of diseases within the prescribed implant period. Furthermore, it did

not indicate any kind of systemic inflammatory reactions. Serum biochemical values of urea,

uric acid, creatinine, serum glutamic–oxaloacetic transaminase (SGOT), serum glutamic

pyruvic transaminase (SGPT), alkaline phosphatase (ALP), total protein, magnesium,

calcium and chloride from rabbits implanted with Mg alloys are listed in Table 4.4. The

kidney-related parameters include urea, uric acid and creatinine levels in the serum. Liver

functions could be monitored by checking SGOT, SGPT, ALP and total protein whereas

normal body electrolytes were monitored, by checking Mg, Ca and Cl levels in the serum. All

the alloys showed elevated values for urea as compared to the control, except Mg-Zr-2Sr.

This could indicate that the degradation of Mg alloys caused a limited effect on kidney

functions. Serum ALP and Ca showed elevated values beyond the normal recommended

levels. All the other serum parameters were within the normal range. This suggests that by 3

months, the liver function was not affected by degradation of these alloys. Balance of

electrolytes in the body is necessary for normal functions of cells and organs. The unaltered

electrolyte balance signifies that the in vivo degradation of Mg alloys did not cause any

disturbance to the normal electrolyte levels of the implanted subjects.


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Sl. No. Agent Unit Control Mg-5Zr Mg-5Zr-Ca Mg-2Zr-5Sr Mg-Zr-2Sr Range

1 WBC X103/mL 9.03 ± 2.07 7.07 ± 0.38 6.16 ± 2.50 7.63 ± 2.47 10.52 ± 1.69 5.1 - 9.7 2 LYM % 77.15 ± 16.00 77.03 ± 1.39 27.25 ± 1.34 87.20 ± 5.49 57.62 ± 17.05 39 - 68 3 RBC X106/mL 5.92 ± 0.28 6.50 ± 0.16 5.52 ± 0.35 6.03 ± 0.46 4.76 ± 0.75 5.3 - 6.8 4 HGB g/dL 12.98 ± 1.15 14.27 ± 0.46 12.18 ± 0.69 10.80 ± 2.19 11.18 ± 2.56 9.8 – 14.0 5 HCT % 40.38 ± 0.54 50.37 ± 1.33 35 ± 3.25 45.53 ± 1.42 35.20 ± 2.02 34 - 43 6 MCV fL 68.43 ± 4.20 77.53 ± 0.15 63.45 ± 1.87 75.77 ± 3.63 75.21 ± 10.49 60 - 69 7 MCH pg 21.88 ± 0.87 21.90 ± 1.05 22.01 ± 0.16 18.23 ± 3.74 23.27 ± 2.17 20 - 23 8 MCHC g/dL 32.15 ± 3.28 28.30 ± 1.45 34.80 ± 1.27 24.07 ± 4.39 31.54 ± 5.97 31 - 35 9 PLT X103/mL 418.67 ± 237 472.33 ± 160 1219.50 ± 104 479.00 ± 224 1085.28 ± 316 158 - 650

Table 4.3. Blood cell count of control and experimental animals 3 months after implantation. WBC- White Blood Cells; LYM- Lymphocytes;

RBC- Red Blood Cells; HGB- Hemoglobin; HCT- Hematocrit; MCV- Mean Corpuscular Volume; MCH- Mean Corpuscular Hemoblobin;

MCHC- Mean Corpuscular Hemoblobin Concentration; PLT- Platelet.


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Parameters Control Mg-5Zr Mg-5Zr-Ca Mg-2Zr-5Sr Mg-Zr-2Sr Range Liver functions Urea (mg/dL) 24.91±2.186 35.68±8.426 49.51±9.736 22.83±4.173 25.17±1.914 5.00-25.00

Uric acid (mg/dL) 0.76±0.016 0.82±0.185 0.69±0.084 1.59±0.634 1.13±0.559 1.0-4.3 Creatinine (mg/dL) 1.03±0.046 1.23±0.195 1.39±0.126 0.76±0.062 1.27±0.369 0.5-2.6 Kidney functions

SGOT (U/L) 22.48±2.410 19.18±2.333 61.50±10.136 31.29±3.358 19.92±9.581 10.0-86.0 SGPT (U/L) 28.20±2.870 37.78±1.333 65.42±16.494 37.57±10.667 43.79±2.822 20.0-120.0 ALP (U/L) 116.02±9.770 125.80±2.921 105.13±27.555 89.62±1.324 105.81±0.584 25.0-65.0

Total protein (g/dL) 4.30±0.470 6.52±3.344 3.40±0.301 5.16±0.026 8.42±0.888 5.0-7.5 Electrolytes

Magnesium (mmol/L) 2.50±0.138 2.36±0.534 2.21±0.385 2.61±0.019 2.51±0.387 2.0-5.4 Calcium (mg/dL) 15.95±0.759 17.95±0.002 18.69±1.297 14.85±2.589 16.75±0.059 5.6-12.1

Chloride (mmol/L) 95.08±12.371 95.61±22.144 91.65±4.134 117.48±4.614 105.73±2.681 92.0-120.0

Table 4.4. Values of serum biochemical parameters of control and experimental animals after 3 months implantation. SGOT - Serum

glutamic oxaloacetic transaminase; SGPT - Serum glutamic pyruvate transaminase; ALP - Alkaline phosphatase.


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4.3 Discussion

4.3.1 Analysis of alloying and matrix properties

Implant stability plays a vital role in determining the biological efficacy of endosseous

prostheses. The material stability criteria become even more critical in the case of

biodegradable implants because it is essential for the implant to induce sufficient and quality

bone tissue around itself before its properties are diminished and the implant itself

decomposes. The surfaces of metal-based implants are highly reactive in biological systems

and numbers of reactions are observed at the implant-bone interface that can determine the

stability of the implant and quality of osseous regeneration. Some of the important factors

that facilitate optimum osseointegration include – physicochemical properties of the material

surface, bioactivity of alloying elements, rate and kinetics of implant degradation and the

type of interaction between implant material and surrounding tissues.

4.3.2 Role of individual alloying elements in implant stability

Our work has mainly focused in studying the stability of four different Mg alloys containing

different quantities of Zr, Sr and Ca, and in correlating the surface energy of each alloy to its

capacity for osteoinduction and integration at the implant site in rabbit femur bones. Based on

our results we have demonstrated that Zr-Sr alloys (Mg-2Zr-5Sr and Mg-Zr-2Sr) are superior

biomaterials for orthopaedic applications than alloys that contain only Zr or a combination of

Zr and Ca. It can be seen from the XRD data in Fig 4.1 that there is a formation of Mg17Sr2

intermediate phase in Sr containing alloys (Mg-2Zr-5Sr & Mg-Zr-2Sr). Mg17Sr2 is reported to

be the most Mg-rich stable compound in the Mg-Sr system [162], which can improve the

corrosion resistance of Mg based alloys. Inclusion of Zr in Mg alloys acts as a grain-refining

agent, which causes strengthening of the alloy by formation of α-Zr phases and grain

boundary strengthening [175]. In our study Ca or only Zr containing alloys (Mg-5Zr and Mg-


69

5Zr-Ca) were unstable and upon degradation in vivo they showed a lot of residual alloy

material at the implant site (Figs 4.6 and 4.7). This has not been reported earlier because the

stability of Ca containing Mg alloys has been studied only in the absence of Zr [167, 176-

178] and Mg alloys containing only Zr have not been reported so far.

4.3.3 Role of individual alloying elements in bone formation

The elemental constitution of the alloying metals also affects the properties of newly induced

bone as several charged groups (O, OH- etc) on the alloy surface contribute to the overall

electronegativity of the implant material [179]. In recent report using Ti based implants it has

been shown in addition to the electronegativity of the implant surface that several other

factors can also influence the osteoinduction process in vivo [180]. The advantages and

limitations of individual alloying metals have also been described in our previous report

[181] where we have shown that alloying processes improved the mechanical strength of

implant materials and made their properties similar to those of the bone tissue. Based upon

these criteria we can conclude from our results that Mg-5Zr-Ca implant, which showed high

compressive strength, was not suitable for osseointegration whereas Mg-2Zr-5Sr and Mg-Zr-

2Sr implants that contained Sr tend to resemble more to natural bone in their properties due to

decreased compressive strength.

4.3.4 Analysis of cell and matrix properties in the new bone

For assessment of bone formation in vivo the two most important properties of implant-

induced NB tissue are osteoinduction and osseointegration. We have evaluated these two

processes by histological and immuno-histochemical staining of the peri-implant bone

sections.


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4.3.3.1 Proximity of new bone and implant surfaces

In contact osteogenesis, the NB formation occurs on the surface of the implant itself. The

osteoblasts secrete a collagen-rich bone matrix directly on the implant surface as they

differentiate into osteocytes. In later stages the matrix is mineralised as calcified collagen and

a matrix free area separate the old bone from the NB [182]. In distance osteogenesis, NB is

formed over the old bone surface and the differentiating osteoblasts/osteocytes and the

secreted matrix is all seen in close association with the surface of the old bone. Due to this

pattern of cells and bone matrix, a space is generated between the NB and the implant

surface, which gets occupied with dying or undifferentiated osteoblasts that are unable to

migrate from the bone surface [182].

As mentioned earlier Mg-2Zr-5Sr and Mg-Zr-2Sr alloys induced NB exhibited contact

osteogenesis whereas Mg-5Zr and Mg-5Zr-Ca induced NB had distance osteogenesis

properties, which are highlighted in the H&E staining patterns shown in Fig 4.6. In addition,

the presence of differentiated osteoclasts, which is indicative of better bone resorption

activity, in the Mg-2Zr-5Sr and Mg-Zr-2Sr induced bones (Fig 4.7) suggested that the quality

of bone tissue around these implants was superior to the other bone types. In contrast we

observed hyperactivity of osteoclasts in Mg-5Zr-Ca induced bones, which led to deformed

and weaker bone tissue.

4.3.3.2 Extent of bone mineralisation

The facilitation of bone formation around the Mg-based implants is associated with the

release of free Mg2+ ions during the biodegradation process, which can contribute to the

enhanced activity of osteoblasts and a consequent increase in the quantity of ECM in the peri-

implant bone. In our study we analysed this activity by Col immunostaining and MT staining.

Collagen type-I is an important marker of better quality osteogenesis [183] and we could


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observe this in Mg-2Zr-5Sr and Mg-Zr-2Sr induced bones (Fig 4.8). The extent of

mineralisation is shown by deep blue staining of MT and DXA results (Figs 4.10A & 4.10B).

The blue stained sections distinguishes the mineralised and non-mineralised ECM. DXA

measurements of BMC and BMD gave us the idea of bone growth and accumulation on and

around the implanted site. From the DXA measurements, most of the component of bone

accumulation that is associated with change in bone size is detected as bone mass and bone

density. Our results gave good BMC and BMD for the implanted femur bones, which are in

line with the results of histology and immuno-histochemical studies.

Further, the impact of bone formation based on blood and serum biochemical analysis has

been reported in this research. The elevated blood platelet numbers in Mg-Zr-2Sr in the third

month (Table 4.3) indicates platelet activation, which can make cells aggregate and form a

clot. Stable clot formation is required for secretion of biochemical components for

osteoconduction [184].

4.3.5 Role of surface energy in in vivo biological studies

The surface energy and wettability of the implant material is significant in determining cell-

biomaterial interactions and it significantly influences biological events at the sub-cellular

and cellular level (eg., protein adsorption, cell attachment and spreading etc.) [185]. In our

study, Mg-Zr-2Sr alloy exhibited lowest surface energy which could be attributed to lower

weight percent of Zr and Sr (1 and 2wt% respectively) and hence more availability of free

Mg2+ ions in the vicinity of this alloy. Further, lower surface energy levels of Sr containing

Mg-2Zr-5Sr and Mg-Zr-2Sr alloys in comparison to Mg-5Zr and Mg-5Zr-Ca alloy surfaces

signifies the role of Sr in controlling the surface energy levels of these alloys. Since Mg-2Zr-

5Sr and Mg-Zr-2Sr alloy-induced NB tissues exhibited “contact osteogenesis” properties,

whereas Mg-5Zr and Mg-5Zr-Ca alloy-induced bone tissues showed “distance osteogenesis”


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properties, it can be predicted that Sr containing Mg-Zr alloys with lower surface energy

could induce better bone formation in vivo than high surface energy containing Sr deficient

Mg-Zr alloys.

4.3.6 Recommended strategies for material evaluation and significance of Sr levels in

Mg-Zr implants

The surface integrity of the alloys decides the rate and extent of bone formation,

osseointegration and osteoconduction. In vivo evaluation becomes necessary to understand

how the implant will perform in a physiological environment. Various physicochemical

factors can play equally important roles in new bone induction [180] and hence no single

property should be considered as the only criterion for selecting the implant material for

orthopaedic applications. Our study highlights the importance of including Sr and excluding

Ca in optimal combinations of Mg-Zr alloys to obtain good osteogenic outcome in rabbit

bones. The clinical significance of our study undoubtedly lies in alloying Sr with Mg-Zr

based biodegradable alloys and that 2wt% of Sr alloying in Mg-Zr system can improve the

bone remodelling rate in vivo. Although we have used a total of 30 numbers of animals for

the in vivo experiments, more detailed investigation on bone formation on the Mg-Zr-Sr/Ca

implants would need a cohort of animal study before the implants can go for the clinical trial.

Also, supplemental in vivo investigation on larger animal models is required to determine the

efficacy of the Mg alloy system under more complex biological conditions. Nonetheless, Mg-

Zr alloys containing Sr as a complementary bivalent cation in optimum quantities can be used

to promote efficient osseointegration for orthopaedic implant applications.


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4.4 Summary

We have studied a new class of biocompatible and biodegradable implants, composed of Mg

as a base alloy and containing varying quantities of Zr, Sr and Ca. Some basic

physicochemical properties of these alloys have been described earlier [181]. Here we have

estimated the surface energy of the alloys by measuring the contact angle of polar solvents on

the alloy surface, and correlated it with their bone forming capacity in the peri-implant area

by histological and radiological evaluation. It was found through our physico-chemical

investigations that inclusion of Sr into the Mg-Zr alloy increases both the corrosion resistance

and stability of these alloys. Presence of Sr leads to formation of Mg17Sr2 intermediate phase

that strengthens the alloy in Mg-rich system. The compressive strength of these alloys

resemble more to the natural bone upon addition of Sr. Our results show that optimal contents

of Zr and Sr increase the surface energy of the alloys and their capacity to simulate “contact

osteogenesis” in the peri-implant area. The inclusion of Ca without Sr reduces the surface

energy on the alloy surface and decreases its bone inducing activity. The osteogenesis,

osteoinduction and osseointegration of Sr-containing alloys could be seen by mineral and

trabecular bone formation on the implant surface. Further, the DXA, blood and serum

biochemical measurements prove that the bone growth and mineral accumulation on the

implant surface is best for the Sr-containing alloy. The pattern of contact osteogenesis

remained the same in both the Sr-containing Mg alloys. However, the higher concentration of

Sr (5wt%) gave greater level of bone formation probably due to the presence of higher

amount of secondary phase which could later grow uncontrollable as compared to the lower

concentration (2wt%) as seen by histological and immuno-histochemical studies of these

alloys. Through our physico-chemical, biochemical, histological and immuno-histochemical

findings, we report that the presence of Sr gives Mg-Zr implants the property of contact

osteogenesis. Although more detailed in vivo studies are required to confirm the superior


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bone healing capacity of Sr, the present results indicate that Sr in Mg-based alloys notably

increases early bone apposition in rabbit cancellous bones and that 2wt% Sr is the optimum

concentration for Mg-Zr-Sr alloys to get the best osseointegration by influencing the surface

energy of Mg-Zr alloys and their rate and extent of osseointegration. Further, it also helps in

complete degradation of the implant within the specified period of surgery. In future, our

results can guide the formulation of more efficient biodegradable alloys for orthopaedic

applications.

Chapter 5. Collagen Type-I Leads to In Vivo Matrix Mineralisation and Secondary Stabilisation

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Chapter 5

Collagen Type-I Leads to In Vivo Matrix Mineralisation and Secondary Stabilisation

5.1 Introduction

The success of orthopaedic implants depends entirely on the bone-implant interface. Implant

osseointegration in the injured bone site decides the healing and regeneration within the bone

tissue. The improper healing of wounded bones can result in bone fragility, calcification and

abnormal bone quality and quantity. These differences in bone healing can be resolved by

improving the implant stabilisation, which can be brought about with the help of

biodegradable alloy implants. Since the discovery of magnesium (Mg) as a biodegradable

material having combined mechanical, electrochemical, and biological features similar to

natural bone, several reports have been published regarding Mg as orthopaedic implants [12,

162, 171, 186]. It is present at high concentrations in cell and takes part in various cellular

activities [5]. Mg is also a daily dietary requirement [187] and the dissolved Mg2+ ions are

easily absorbed in our body [188]. Further, the mechanical properties of Mg are comparable

with that of cancellous bone [8, 189]. However, the rapid dissolution in chloride-containing

aqueous solutions [190] and low corrosion resistance of Mg and its alloys, leading to earlier

degradation than the actual period of bone healing process still limit the use of Mg-based

alloys as clinical implants.

We have alloyed Mg with Zirconium (Zr) and Calcium (Ca) to improve its rate of

osseointegration and implant stabilisation in a host tissue. Ca is reported to be a favourable

non-toxic Mg alloying element and its addition can significantly enhance the mechanical

properties and corrosion resistance of Mg alloys [92]. Zr has a powerful grain refining effect

on Mg alloys, which improves the corrosion and mechanical properties of these alloys [181].


76

Mg-Zr-Ca alloys are expected to be a promising candidate for use as an implant material. On

the other hand, many reports have demonstrated that Collagen type-I (Col-I) imparts

favourable effects on attachment, function and growth activities of osteoblast cell lines [138,

191]. Besides Col-I, other extra cellular matrix (ECM) proteins such as fibronectin, through

their Arginyl-Glycyl-Aspartic acid (RGD) peptide sequence, also influence osteoblast

differentiation [192]. Incidentally, MC3T3-E1 cells have been reported to contain both the

collagen binding integrin subunits α2β1, as well as fibronectin binding integrin subunit α5β1

[193]. Col-I has also been presented to display osteoconductive properties in vivo [194, 195].

Although potentially beneficial in stimulating new bone formation, there is insufficient

information about the cumulative effects of Col-I coating and Mg, Zr and Ca as alloying

elements. Their performance in osseointegration and osteoconduction as in vivo implant

would elucidate the capacity of Mg-based alloys to bridge the pre-prosthetic gaps that arise

from the poor stability of alloy implants. Successful osseointegration depends on the primary

stability of the alloy implant, which further influences the secondary stability. Primary

stability is the stability shown by alloy material upon implantation. Once the reconstruction

of bone tissue is initiated around the implanted material, the implant undergoes secondary

stabilisation, characterised by its ability to promote osseointegration. This study describes a

combinatorial approach of alloying Mg with Zr and Ca with the application of a bioactive

coating of Col-I on the alloy implant surface. We have also reported in our previous work

that the uncoated Ca-based Mg-Zr alloys corrode rapidly and hence are not bio-effective in

osseointegration [171], but their physicochemical properties are enhanced when coated with

Col-I thereby improving the bio-efficiency of these implants. Our findings demonstrate how

Col-I coating Mg-Zr-Ca alloys give secondary stability to the implant by making them more

osteoconductive over a period of 1 and 3 months implantation in rabbit femur and promote

early and enhanced osseointegration.


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5.2 Results

5.2.1 Experimental animals

The mean body weights of the implanted male New Zealand white rabbits, each for 1 and 3

months is given in Table 5.1.

5.2.2 Alloy microstructure and mechanical properties

The XRD pattern (Fig. 5.1A & 5.1B) shows the phase constituents of Mg-based alloys. The

α-Mg phases on the Mg-5Zr alloy surface could be seen with traces of unalloyed Zr. There

are reports indicating that Ca addition leads to the formation of a thermal stable second phase

Mg2Ca, which results in grain refinement of Mg casting [196, 197]. It is also noticeable that

an intermetallic Mg2Ca phase coexists with α-Mg in the Mg-5Zr-Ca alloy. This is confirmed

by microstructure images shown in Fig. 5.1C & 5.1D, showing intermetallic Mg2Ca

precipitating along the primary α grain boundaries with a small amount of Ca phase. The

grain size of Mg–5Zr-Ca is also seen to be much finer than that of Mg-5Zr. However, the

microstructure of Mg–5Zr shows some unalloyed Zr particles (arrows) due to low solubility

of Zr in Mg (only 3.8 wt%) [198]. The grain boundaries of Mg-5Zr-Ca are also rougher and

broader than that of Mg-5Zr. The compressive strengths of Mg-5Zr and Mg-5Zr-Ca alloys

were measured to be 237.8±10.1 and 255.7±7.2 MPa, respectively, making both the alloys

behave close to the natural bone. Hydrogen evolution test was carried out to study the

Alloy Before surgery (n=3) (Kg)

After 1 Month (n=3) (Kg)

After 3 Months (n=3) (Kg)

Mg-5Zr 2.540±0.285 2.807±0.463 2.912±0.243

Mg-5Zr-Ca 2.436±0.207 2.844±0.362 2.946±0.047

Table 5.1. Measured body weights of experimental animals for Col-I coated Mg-5Zr and Mg-

5Zr-Ca alloy implants after 1 month and 3 months implantation. 3 animals were used for each

alloy for each time period of implantation.


78

variation in corrosion rates with exposure time and the results are shown in Fig. 5.1E. Mg-

5Zr-Ca alloy displayed a high hydrogen production rate in SBF as compared to Mg–5Zr alloy

and it degraded almost entirely after 72 h of immersion time.

Figure 5.1. Panels A and B show the phases of Mg-5Zr and Mg-5Zr-Ca alloys, respectively,

by X-Ray Diffraction. The α-Mg dominant phases and the unalloyed Mg(103)/Zr(103) phase

could be seen on the surface of both the alloys. Secondary Mg2Ca phase could be seen on the

surface of Mg-5Zr-Ca alloy. Panels C and D show the optical microstructures of Mg-5Zr and

Mg-5Zr-Ca alloys, respectively, indicating unalloyed Zr particles along the grain boundaries.

Panel E displays the hydrogen production rate of Mg-5Zr-Ca alloy showing higher amount of

hydrogen release which increases with time.


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5.2.3 Collagen coating properties

The purity of the rat-tail tendon collagen prepared by us is shown in Fig. 5.2 (panels A and

B). The coverage and integrity of the coating, as seen by FT-IR spectra, is demonstrated in

Fig. 5.2, panel C. The SEM analysis of the exposed substrate-coating-interface indicated that

the average coating thickness was about 1-1.5 μm (Fig. 5.3).

Figure 5.2. Coomassie stained SDS-PAGE (Panel A) of Col-I extracted from rat tail on 8%

resolving polyacrylamide gel. The gel extracts were then transferred to PVDF membrane for

western blotting with mouse anti-collagen (Panel B). All three characteristic bands (α1, α2

and β) that make a functional collagen were visible. Panel C shows the representative Fourier

Transform-Infra Red (FT-IR) spectrum. The peaks for all the four amide groups

corresponding to Col-I could be seen, including amide I region corresponding to the C=O

absorption, amide II corresponding to CH2 and CH3 absorptions, amide III corresponding to

C-N and N-H absorptions and amide IV corresponding to C-O and C-O-C absorptions of

carbohydrate moieties. The amide I region (1600-1700 cm-1) of the spectrum of (Pro-Pro-

Gly)10 is where the collagen peptide is known to exist in a triple helical state.


80

5.2.4 Surface roughness and surface energy

The surface roughness measurements showed similar values for the respective groups i.e.,

uncoated and coated Mg-5Zr and Mg-5Zr-Ca alloys (Table 5.2). Increased surface roughness

was seen in the coated alloys.

Fig. 5.4 and Table 5.3 respectively shows the shapes of individual drops on the bare and

coated alloys and left and right contact angles with the calculated surface energy values. The

surface energies were calculated from the corresponding mean left and right contact angles

and were in the range of 16-67 μJ/mm2. It is known that substrates with low surface energies

are more hydrophobic than those with high surface energies [174]. The coated Mg-5Zr-Ca

alloy showed the lowest average surface energy value of 18.99±1.97 μJ/mm2 making this

alloy the most hydrophobic while the bare Mg-5Zr-Ca alloy without Col-I coating is the most

hydrophilic.

Alloy Ra Uncoated Ra Coated

Mg-5Zr 0.129±0.00488 0.683±0.0157

Mg-5Zr-Ca 0.146±0.00566 0.614±0.00672

Table 5.2. Surface roughness of the bare and Col-I coated Mg-based alloy discs. Values were

obtained from duplicate experiments using the alloy discs.

Figure 5.3. Thickness of the coated collagen on the alloy surface, visualised by SEM.


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Alloy Liquid Left CA Right CA

Bare Mg-5Zr

Water 25.25° 0.92 27.25° 0.49

Glycerol 43.9° 1.27 46.65° 0.07


Bare Mg-5Zr-Ca

Water 23.3° 0.71 24.3° 0.99

Glycerol 45.7° 0.57 48.1° 1.98


Coated Mg-5Zr

Water 81.25°±0.49 82.65°±0.21

Glycerol 73.05°±0.64 70.30°±0.14

Average SE 31.35°±0.14 μJ/mm2

Coated Mg-5Zr-Ca

Water 100.70°±4.67 100.95°±5.59

Glycerol 96.65°±0.07 97.80°±1.98

Average SE 18.99°±1.97 μJ/mm2

Table 5.3. Individual surface energy of Col-I coated Mg-based alloys calculated from the

respective left and right contact angles measured using water and glycerol. Contact angle

measurement was done within 60 seconds after the liquid drop touched the alloy surface.

Each alloy disc was used for measuring the contact angles for 3 times with 2 μL of liquid

droplet on 3 different positions on the alloy surface, both with water and glycerol.

Figure 5.4. Shapes of water and glycerol droplets used to estimate the contact angles of

uncoated and Col-I coated Mg-based alloys. The contact angles were recorded within 60

seconds of the droplet touch on the alloy surface.


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5.2.5 Surface morphology by SEM

Fig. 5.5 (A-D) shows the surface morphology of uncoated and coated Mg-based alloy discs

observed by SEM. Apart from irregular cracks and floral arrangements on the uncoated alloy

surface there was no obvious difference in their morphology. The coated alloy surface

showed presence of thin folds on their surface because of the dried layer of coated collagen.

5.2.5 Cell viability

The cell viability ratio (Fig. 5.5E) of MC3T3-E1 osteoblast cells after incubation for 12 h in

the coated alloy extracts gave better results than those in the uncoated alloy extracts. It is

evident from this point that Col-I coating made the alloy surfaces more compatible for

osteoblast adhesion and viability.

5.2.7 Histology and immunohistochemistry

The ossification and secondary stabilisation around the Col-I coated Mg-based implants was

checked by HE and TRAP staining. Fig. 5.6 (A, B, E and F) elucidates the implant-bone

Figure 5.5. Scanning electron micrographs of the uncoated (Panels A and B) and Col-I coated

(Panels C and D) Mg-based alloys. Panel E shows the cell viability of MC3T3-E1 osteoblasts

on the uncoated and coated Mg-based alloys. ***p<0.001.


83

interface by HE staining, 1 and 3 months post-implantation. The coated Mg-5Zr alloy after 1

month implantation showed a continuous fibrous layer around the implant area and a small

layer of newly formed trabecular bone around it. After 12 weeks post-implantation, this

fibrous layer becomes thinner and more amount of new bone can be seen in place of it. In

coated Mg-5Zr-Ca implant, the formation of a rigid layer of new bone can be seen only after

1 month implantation which indicates obvious ossification further leading to secondary

stabilisation upon 3 months implantation. In Fig. 5.6 (C, D, G and H), the TRAP positive

osteoclasts start becoming visible in the coated Mg-5Zr-Ca implant after 1 month

implantation itself which becomes denser after 3 months. Higher number of osteoclasts in the

vicinity of Mg-5Zr-Ca implant indicates better bone resorption and hence more appropriate

bone remodeling around this implant. The degradation products are also visible as black

residues which can be the reason for heightened osteoclast activation in the peri-implant area.

The poor degradation of coated Mg-5Zr implant even after 12 weeks indicates its lesser

osteoclast-inducing capacity. Fig. 5.7 shows mineralisation of the newly formed bones by

MT staining and Col-I immunostaining, respectively. The unmineralised osteoid and the

mineralised ECM were stained in red and blue, respectively, in MT staining of the bone

tissue implants. The coated Mg-5Zr implant after 1 month showed some fibrous tissue and

partially mineralised new bone trabeculae, distant from the implant area while a broader,

more interconnected and fully mineralised bone close to the implant surface is visible for

coated Mg-5Zr-Ca implant 1 month post-implantation. Since the mineral content of the

woven bone is high, the 3 months implanted images indicate facilitation of rapid bone

formation and mineralisation.


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Figure 5.6. Images of bone sections stained with HE (Panels A, B, E and F) and TRAP

(Panels C, D, G and H) after 1 month (Panels A-D) and 3 months (Panels E-H) implantation

with coated Mg-5Zr and coated Mg-5Zr-Ca alloys. Note the rigidity of new bone in the

coated Mg-5Zr-Ca implanted 1 month bone section indicating stronger and more mature bone

formation within a short period of time. The coated Mg-5Zr-Ca alloy shows good expression

of osteoclasts only after 1 month implantation (Panel D) which becomes denser after 3

months implantation (Panel H) compared to the coated Mg-5Zr alloy.


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Figure 5.7. Mineralisation of the newly formed bone shown by Masson’s Trichrome staining

of bone sections after 1 month (Panels A and B) and 3 months (Panels E and F) implantation,

and Collagen-I immunostained images of bone sections after 1 month (Panel C and D) and 3

months (Panels G and H) implantation. Matured mineralised new bone could be seen for the

coated Mg-5Zr-Ca alloy as compared to the coated Mg-5Zr alloy after 3 months

implantation. In the Collagen stained images, the 3 months sections (Panels G and H) showed

uniform arrangement of collagen stained osteoblasts which was seen to be haphazardly

positioned in 1 month sections (Panels C and D).


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The new bone mineralisation is supported by Col-I immunostaining, which shows dense Col-

I stained osteoblast lining around the implanted alloys. Even though Col-I positive osteoblast

cells are uniformly distributed around both the coated alloys for 1 month implantation, its

levels were higher in 3 months implanted coated Mg-5Zr-Ca implant. It can also be seen from

Fig. 5.8 that the 3 months implanted coated alloys show a very dense and well-formed

network of vimentin indicating the integrity of cytoplasm and cytoskeletal interactions in the

ECM of the implanted alloys.

The 3 months coated implant-B induced new bone shows the most uniformly arranged

vimentin network clearly suggesting its role in better and stronger bone formation. The new

bone formation by uncoated Mg-based implants has been shown in Fig. 5.9 and demonstrates

very less new-bone formation. The surrounding fibrous layer of the implant area in HE

staining indicated inflammatory response towards uncoated implants. The new bone matrix is

seen in a distance away from the implant surface, giving rise to the phenomenon of distance

Figure 5.8. Vimentin immunostained images of coated Mg-5Zr and coated Mg-5Zr-Ca alloys

implanted bone sections after 1 month (Panels a&b) and 3 months (Panels c&d) implantation.

Noticeably uniform and dense vimentin network is seen for 3 months implanted sections.


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osteogenesis. TRAP staining showed less number of osteoclast activation around the

uncoated implants as compared to the Col-I coated implants indicating less amount of bone

remodelling around the bare implants. MT staining showed very less new bone mineralisation

of the uncoated Mg-5Zr-Ca implant, which becomes dramatically enhanced after Col-I

coating as seen in Fig. 5.7.

Figure 5.9. Histology images of uncoated Mg-5Zr and Mg-5Zr-Ca implanted bone sections

after 3 months implantation. Panels A and B- HE staining, Panels C and D- TRAP staining

and Panels E and F- Masson’s Trichrome staining.


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5.2.8 BMC/BMD by DXA and blood cell count

The quantitative in vivo new bone mineralisation was deduced from the mean BMC/BMD

values of the implanted femur obtained by DXA radiography after 1 and 3 months

implantation (Fig. 5.10). The BMC values showed higher mineral content for the coated

alloys for both the implant periods. These values were highest for coated Mg-5Zr-Ca implant

as compared to the control and coated Mg-5Zr implant. The BMC values were lowest for the

uncoated Mg-5Zr-Ca implant upon 1 month implantation. However, there was no significant

change for the BMD values of 3 months implanted alloys as compared to the control. This is

possibly because the healed areas in these femur bones were much bigger than the healed

areas of the control.

Figure 5.10. DXA radiography detailing the BMC and BMD of the uncoated and Col-I

coated Mg-based alloys after 1 month and 3 months implantation. A significant raise in the

BMC/BMD of 3 months implanted bone for coated Mg-5Zr-Ca implant directly suggests its

enhanced osseointegration capacity giving rise to better bone content and density. 1M - 1

Month, 3M - 3 Months. *p<0.05, **p<0.01.


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Table 5.4 lists the total blood-cell count of the rabbits performed before surgery, and after 1

and 3 months of implantation giving information on inflammation and immune mechanism of

the body. A significant increase in platelet (PLT) count was found for coated Mg-5Zr-Ca

implant after 1 month. This could be probably due to higher platelet activation because of

better bone induction and bone formation. The number comes down to normal level after 3

months implantation period. Hence, this elevation does not indicate any kind of disturbances

in the blood count.

5.3 Discussion

The surface characteristics and its influence on bone formation and mineral apposition is

undeniably an important feature of the cell response. However, such data provide inadequate

information regarding implant-host tissue interactions and their biological effects. To date,

various strategies have been employed to improve this interaction by modifying the surface

characteristics such as surface energy, roughness, topography and chemistry of the target

alloys [179]. In the present study we have surface coated Ca-comprising Mg-5Zr alloys with

Col-I and hypothesised that an enhancement in osteoinduction would lead to in vivo

secondary stabilisation. Our results prove that coated Mg-5Zr-Ca implant demonstrate

enhanced trabecular bone volume and mineralised bone matrix leading to implant secondary

stabilisation. The elevations in surface roughness and hydrophobicity of the coated Mg-5Zr-

Ca implant may have played a role in this development. A major concern about using these

implants resides in its tendency to react very fast in a physiological pH. The Mg-based

implant surface undergoes a shielding outer layer of corrosion referred to as “passivation”

[199].


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Blood cells Before surgery 1 Month implantation 3 Months implantation Range

Control Mg-5Zr Mg-5Zr-Ca Control Mg-5Zr Mg-5Zr-Ca RBC (X106/uL) 6.44±0.43 5.85±0.72 6.37±1.21 6.41±0.49 6.87±0.37 6.82±0.64 6.75±0.22 5.3 - 6.8 WBC (X103/uL) 6.87±1.32 13.26±2.35 9.33±2.03 7.91±3.37 8.10±0.57 9.22±1.01 11.82±4.08 5.1 - 9.7

LYM% 85.12±8.13 63.05±25.38 83.08±2.29 69.21±6.58 58.75±2.05 87.13±4.76 82.97±1.65 39 - 68 HGB (g/dL) 14.22±0.68 12.85±1.06 13.95±2.89 13.71±1.67 13.35±0.92 13.92±0.78 13.48±0.49 9.8 – 14.0

HCT (%) 53.03±1.13 39.66±7.58 50.85±7.33 46.53±3.16 43.30±2.26 54.20±3.91 53.22±3.28 34 - 43 MCV (fL) 82.65±7.28 67.53±4.66 80.15±3.70 72.60±0.71 63.15±0.07 80.00±2.26 78.85±2.24 60 - 69 MCH (pg) 22.03±0.47 21.95±0.92 21.76±0.42 21.28±0.97 19.40±0.28 20.40±0.80 19.93±0.09 20 - 23

PLT (X103/uL) 598.00±176 645.20±124 474.50±136 1271±135.76 204±22.63 345.83±85.56 391.17±244.89 158 - 650

Table 5.4. Blood cell count of the control and experimental animals implanted with Col-I coated Mg-5Zr and Mg-5Zr-Ca alloy implants. Cell

count was performed before surgery and after 1 month and 3 months implantation period for the individual implanted alloys from the control and

experimental animals.


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When exposed to air, this phenomenon produces a magnesium oxide (MgO) layer that

prevents the alloy surface from further reactions. However, this layer dissolves very fast in a

physiological pH and hence the high chloride content of the host system causes corrosion on

the implant surface. In our study, although Col-I coating imparted osteoinducing property to

the alloy implants, the corrosion and degradation factors after coating remained

underestimated. One aspect of Col-I coated alloy corrosion can be the presence of defects

such as cracks, pits and vacancies inside the coating film, which enhances the corrosion rate

[200], and might weaken the effect of coating. In our study, the degradation of a material

happened because of its own solubility in the aqueous media and because of the reaction

happening between the ions present in the media and the corrosion products of the material.

Due to these activities, the in vitro degradation of the alloy is faster because of the

availability of more aqueous media. However, the in vivo degradation is expected to be

slower as there is less exposure of the implanted material to the aqueous media. During this

period, the induction of new bone happens rapidly, which provides strength to the peri-

implant area, and the original strength of the degrading material becomes insignificant.

5.3.1 Col-I coating and secondary stabilisation

Our previous studies have demonstrated that uncoated Mg-5Zr-Ca alloys suppress bone

turnover [171]. In the current study, the bone turnover, osseointegration and secondary

stabilisation was achieved by twice its rate compared with the normal and uncoated Mg-

based implants. The histology results indicate direct new bone apposition with the implant

surface and its mineralisation 3 months post-implantation after Col-I coating. Although there

remains an intervening fibrous layer between the new bone matrix and the implant surface in

1 month implanted alloy samples, it can be speculated that with time Col-I coating could help


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in the surface fixation and secondary stabilisation of the Mg-5Zr-Ca alloy implant, which is

clearly visible in 3 months implanted sections.

The success of orthopaedic implants depends on implant stability, which can be divided into

primary and secondary stability. The nature of the alloy decides the implant primary stability,

which depends on the mechanical characteristics of the bone and the type of implant used

[201]. They play a major role in the early events of endosseous healing and contribute to

improved secondary stabilisation, which is achieved by contact osteogenesis [182]. It induces

the formation of new bone through osteoconduction and gradual bone remodelling resulting

in a stable new bone matrix at the bone-implant interface. This mechanism can be seen in our

histology outcomes of 3 months implanted coated alloys, wherein the new bone formation

can be seen in close proximity to the implant surface (Fig. 4E and 4F), which gradually

progresses to mineral apposition (Fig. 5E and 5F). In contrast, the uncoated alloys

demonstrate the mechanism of distance osteogenesis [171, 182], which causes stress

shielding resulting in a number of critical clinical issues such as early implant loosening,

damage of adjacent anatomical structures etc. [71].

The role of Col-I in bringing this stabilisation to the coated alloys can be illustrated by its

property of RGD peptide-cell interactions via integrin α2β1 [193]. Under appropriate

circumstances, Col-I act as effector substance similar to a peptide hormone or cytokine that

activates some of the differentiation-promoting activities associated with the ECM of

osteoblast cells and stimulate their differentiation. Our data suggests fabrication of a well-

formed and rigid trabecular bone and its mineralisation in coated Mg-5Zr-Ca implant without

any inflammation after 4-weeks of implantation, similar to the uncoated group after 12 weeks

of implantation [171]. These results imply that it is the availability of a connecting factor for

cells to dissolving Ca2+ ions, and not the Ca concentration, that contributes to osteoinduction.

In our work, this role has been played by Col-I. It should, however, be emphasised that this


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enhancement in new bone induction does not reflect any improvement in the mechanical

characteristics of the implant material. Nonetheless, a significant difference in the physico-

chemical properties of coated alloys could be seen in our data. The elevated surface

roughness and surface hydrophobicity of the alloy implants indicate induced adsorption of

serum proteins upon contact with blood plasma. This mechanism can accelerate the formation

of a stable bone-implant interface formed by reaction of the biologic cement-line matrix with

the sensitive protein layer on the surface of the implanted area. This fact is supplemented by

DXA of the experimental animals that gives us the BMC/BMD profiles of the implanted area

for the quantitative bone growth and mineral accumulation on the implant surface. In our

study, it is found to be highest in the coated Mg-5Zr-Ca implant in both 1 month and 3

months implanted animals. This suggests better deposition of bone matrix on a pre-existing

material surface and hence improved bone mineral density.

5.3.2 Col-I coating and bone remodelling

Osseointegration and implant stabilisation take place simultaneously with bone remodelling

that occurs via a defined sequence of events [202] within the host bone and subsequently

within the woven bone formed in the peri-implant interface. With time, this remodelling

creates increased contact between the implant surface and the newly formed bone. This leads

to formation of a mature cortical bone, which is the desired end result. Woven bone consists

of loosely packed collagen fibers of varying size in a random spatial arrangement. Our data

demonstrate woven bone formation in the form of trabecular bone (Fig. 4A and 4B) in 1

month implanted animals which later progresses to cortical bone formation (Fig. 5E and 5F)

in 3 months implanted animals. This study, along with the reports published earlier [203],

suggest that over time, Col-I coated Mg-based alloys show remodelling and improved

biological interlocking between implant and the host bone. Since more osteoclasts are


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required for bone remodelling, it is not surprising from our data that the trabecular bone of 3

months coated Mg-5Zr-Ca implant can remodel faster than the coated Mg-5Zr implant. This

process of bone formation and resorption is coupled spatiotemporally and the required

microscopic architecture for maintaining the mechanical properties of bone is preserved.

Remodelling also depends on the preliminary biological processes of the implant area eg,

blood coagulation, inflammation, and tissue formation. Platelets are the first cells to come in

contact with the implant surface immediately after contact with blood [204]. Activation of

platelets results into clot formation, which is necessary for the mechanical and biochemical

components required for osteoconduction. The protein conformation of the host blood also

determines the availability of the bioactive peptide sequences located within the protein for

incoming cells [205].

5.4 Summary

The mechanical properties of Mg-5Zr and Mg-5Zr-Ca alloy implants are comparable to that

of the natural bone; although Mg-5Zr-Ca alloy has a high corrosion rate in vitro, its

performance as an implant material could be stabilised by surface coating with Col-I. Col-I

coating enhances the surface energy and hydrophobicity of these alloys, and strongly

influences the protein binding capacity onto the alloy surface leading to better osteoblast

activity. Histology, immuno-histochemistry and radiology results suggest that Col-I coating

enhances the rate of osseointegration with good level of new bone formation after only a

month of implantation. It can be comprehended from our histology results that with time,

Col-I coating helps in surface fixation and secondary stabilisation of the Mg-5Zr-Ca alloy

implant. From our studies, we cannot absolutely determine whether the improved

osseointegration, mineralisation and implant secondary stabilisation is solely due to Ca


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content of the Mg alloys or due to the altered surface topography, but can however be

comprehended due to the combined effects of Ca ions and Col-I coating.

Chapter 6. Strontium Content and Collagen-I Coating Influence Osteogenesis and Bone Resorption

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Chapter 6

Strontium Content and Collagen-I Coating Influence Osteogenesis and Bone Resorption

6.1 Introduction

Magnesium (Mg) alloys are attractive for several biomedical applications, ranging from

dental prostheses [206] to cardiovascular implants [207]. They exhibit mechanical properties

that are similar to human bone tissue such as similar material density and compressive

strength [8]. In addition biocompatible Mg alloys are non-toxic and biodegradable – their in

vivo biodegradation progressively stimulates their own activity and minimises their re-

accumulation and toxicity [8].

The role of other metals in Mg alloys has been shown to influence their bone forming

properties [6, 162, 166, 208]. Among other metals, the use of Sr in implants has become

popular because strontium ranelate (SrR) has been approved for the treatment for

osteoporosis, under the trade name of Protelos [209, 210]. Our earlier studies had

demonstrated improved in vivo biocompatibility, osseointegration and bone remodelling by

modulating the contents of Sr and Zr in Mg-Zr-Sr implants [171, 211]. Other studies have

also indicated the role of Sr content in implants for better bone formation [212, 213],

improved bone resorption [214, 215] higher bone collagen synthesis [212], more precipitation

of hydroxyapatite [216, 217] and increased hardness of trabecular and cortical bones [218].

Sr has been shown to have some negative effects on bone formation also such as defective or

abnormal mineralisation of the bone matrix [219-221], hence the use of Sr in bone implants

needs to be re-evaluated. The use of implants with bioactive coatings, such as RGD peptides

or Collagen Type-I (Col-I), also stimulates osteoinduction [222]. These coatings enhance the

interaction between the implant surface and extracellular matrix (ECM) at the implant site


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thereby increasing mineralisation of the pre-existing bone and making it more rigid and load-

bearing.

In this study we have analysed the in vitro and in vivo properties of Col-I coated Mg-Zr

alloys containing high (5wt%) or low (2wt%) quantities of Sr. We have evaluated the dose

and time-dependent influence of Sr quantity and Col-I coating of the implants. Our

investigation indicated that Sr content of implants influenced bone specific gene expression

in vitro and implant-bone apposition during in vivo bone formation; Col-I coating of implants

enhanced the in vivo mineral deposition in newly formed and pre-existing bones.

6.2 Results

6.2.1 Cell viability, cell adhesion and mineralisation

The MC3T3-E1 osteoblasts showed good compatibility and viability with the Col-I coated

Mg-Zr-Sr alloys. Fig. 6.1 shows the cell viability ratio with the uncoated and coated alloy

extracts.

Figure 6.1. Cell viability ratio of MC3T3-E1 osteoblasts cultured with alloy extracts of bare

and Col-I coated Mg-Zr-Sr alloys and incubated for 24 hours. *p<0.05, **<p<0.01,

***p<0.001.


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Fig. 6.2 depicts the SEM images of morphology and mineralisation of osteoblasts on the

uncoated and coated alloy surfaces. The osteoblasts exhibited cell death after incubation for

24 h on the uncoated surface of no-Sr alloy. However, after coating the alloys with Col-I and

when cultured under similar conditions, better cell adhesion demonstrating extensive network

and mineral deposition was observed.

6.2.2 In vitro Expression of bone markers

The comparative gene expression profiles of bone related markers in osteoblasts, after their

growth for 1 or 7 days in different uncoated alloy extracts were measured by RT-PCR (Fig.

6.3). Expression of all genes was clearly detectable in cells after 7 days of growth in all alloy

extracts, however, it was higher in cells grown in extracts of Mg-2Zr-5Sr as compared to

those grown in extracts of Mg-Zr-2Sr. This indicated that higher Sr containing alloys had

better gene inducing ability directed towards bone differentiation. Interestingly BMP-2 gene

expression was detectable at 24 h and 7 days but its levels were not significantly different

among the alloys at both time points.

Figure 6.2. Morphology of MC3T3-E1 osteoblasts on bare and coated Mg-Zr- based alloy

surfaces after 24 h of incubation.


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Figure 6.3. The expression of genes related to bone specific markers, proteins related to bone

ECM and proteins of non-collagenous bone matrix after 24 h (1d) and 7 days (7d) of

incubation in the respective coated-alloy extracts. *p<0.05, **p<0.01.


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6.2.3 In vivo bone inducing activity

Fig. 6.4 shows HE stained images of the newly formed bone around the Mg implants, after 1

and 3 months implantation. In the coated implants, initiation of bone formation could be seen

early after 1 month implantation (Panels A-C), which gives rise to mature and fully

developed trabecular bone after 3 months (Panels D-F). Very little or no discontinuity

between the cement line matrix and the newly formed bone could be seen for coated Sr-

containing implants, whereas a large gap could be seen for coated no-Sr implant. The

Figure 6.4. Hematoxylin and Eosin staining of bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys after 1 month and 3 months post-implantation. Panels A-C:

Col-I coated 1 month study, Panels D-F: Col-I coated 3 months study, Panels G-I: Uncoated 3

months study.


101

uncoated alloys also showed good amount of bone formation by Sr-containing implants, after

3 months implantation (Panels G-H).

The TRAP-positive cells were minimal in the sections of coated implants (Fig. 6.5, panels A-

F). However, the number of TRAP stained cells was seen to be higher in the bone sections of

uncoated implants, especially for high-Sr alloy, suggesting excessive bone turnover (Fig. 6.5,

panel I). This difference in staining, between the alloys containing low and high-Sr and

between the uncoated and coated implants, indicates the role of Sr concentration and Col-I

coating in bone induction and turnover by Mg-Zr-Sr alloys.

Figure 6.5. TRAP staining of osteoclasts for bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys after 1 month and 3 months implantation. Panels A-C: Col-I

coated 1 month study, Panels D-F: Col-I coated 3 months study, Panels G-I: Uncoated 3

months study.


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6.2.4 Implant induced bone mineralisation

Fig. 6.6 shows Masson’s Trichrome staining, demonstrating coated (panels A-F) and

uncoated (panels G-I) peri-implants. Homogeneous mineral deposition on the new bone could

be seen in all the implant types, with superior bone mineral in coated high-Sr implant, after

both 1 month and 3 months (panels C & F).

Interrupted and heterogeneous bone mineral deposition was seen by the uncoated implants 3

months post-surgery (Panels G-I). Irrespective of the Sr concentration in uncoated alloys,

bone mineralisation could be seen in all the implant types. However, the time- and coating-

dependent bone mineral deposition by the coated implants could have been influenced due to

Figure 6.6. Masson’s Trichrome staining of bone sections implanted with Col-I coated and

uncoated Mg-Zr-based alloys 1 month and 3 months post-implantation. Panels A-C: Col-I

coated 1 month study, Panels D-F: Col-I coated 3 months study, Panels G-I: Uncoated 3

months study. The blue stain indicates collagen matrix of mineralised bone.


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controlled release of Sr ions by Col-I coating in the vicinity of the implant area. This result

matches the Alizarin Red staining of coated implants (Fig. 6.7) in which the coated low-Sr

implant showed partial Alizarin Red staining for the same 3 months period of implantation,

whereas coated no-Sr implant displayed very little staining. The Ca content of newly formed

bone indicates its extent of maturity, and the bone response observed in the sections of coated

implants suggests the influence of Sr content in bone formation and its degree of

calcification.

6.2.5 Bone density and mineral concentration

The BMC and BMD of the implant sites give us an idea about bone healing from the mineral

content and density of the newly formed bone. As seen in Fig. 6.8, there is a significant

increase of BMC and BMD in the coated high-Sr implant site, compared with the uncoated

and coated alloys having no Sr or lower concentration of Sr, in both the implant periods (1

Figure 6.7. Alizarin Red staining of bone sections implanted with Col-I coated Mg-Zr-based

alloys after 1 month and 3 months implantation. Panels A-C: Col-I coated 1 month study,

Panels D-F: Col-I coated 3 months study. The red stain indicates accumulation of calcium in

mineralised bone.


104

month and 3 months). These results agree with our histology data, which confirms that Col-I

coated high-Sr implant is superior to other uncoated and coated alloys in enhancing bone

mineral formation in rabbit femur.

6.3 Discussion

There has been significant progress in the development of biodegradable alloys that exhibit

good biocompatibility and osteoinducing properties [223, 224]. Mg based alloys have proven

to be useful in this direction because of their properties mentioned above (see references in

Introduction). However, further optimisation of their constitution and modifications of their

surface characteristics are necessary to make them biologically and physico-chemically more

suitable for medical applications. Extending upon some of our earlier results on Mg based

alloys [171, 181, 211], we have enumerated the role of Sr concentration and Col-I coating in

Mg-Zr-Sr alloys with reference to their in vitro and in vivo bone inducing potential.

Figure 6.8. Quantitative in vivo mineralisation of the peri-implant area by evaluation of femur

BMC and BMD of experimental animals implanted with Col-I coated- Mg-Zr-based alloys

after 1 month and 3 months implantation. *p<0.05.


105

6.3.1 Sr concentration and Col-I coating enhance osteogenesis

Our results in Fig. 6.1 and Fig. 6.2 showed that high Sr concentration in Mg-based alloys is

associated with superior viability and proliferation of osteoblasts on their surfaces. These

alloys also induced higher expression of genes related to bone markers (Fig. 6.3). The level of

bone formation and its maturation was found to be affected by time up to three months, as

seen from the histology and DXA results (Figs. 6.4-6.8). Since the collagenous matrix of

newly formed bone is depicted by Masson’s Trichrome staining, our results indicate that

increased Sr concentration in Mg alloys influences the collagenous matrix formation by

osteoblasts, without inducing any negative effects in long-term matrix mineralisation, as also

stated by Barbara et al. (2004) [225]. These results are also consistent with previous studies

showing that Sr increases trabecular bone formation in vivo in rats, mice and humans [226].

From the histology results, no defects in bone apposition and calcification were observed.

This homogeneity could be because all the existing bone around the peri-implant area

consisted of new bones, formed during the period of implantation. Also, there is a decreased

expression of osteoclast markers in the bone sections of high-Sr-containing implants,

indicating the inhibitory effect brought upon by Sr concentration on osteoclast resorption,

without any cytotoxic effects. Although the precise mechanism of the role of Sr concentration

on osteogenesis and bone turnover is not fully understood, the increased bone formation and

its suppressed turnover, within the specified implant period, suggest that Sr is dose

dependently taken up and distributed by the osteoblasts in newly formed bone tissue.

6.3.2 Role of Sr concentration in bone formation and resorption

The Sr ions released from the degrading implant possibly affects the osteocytes, which in turn

controls the activity of osteoblasts and osteoclasts [227]. The mechanical pressure applied by

the newly forming bone matrix is sensed by osteocytes. Less activity in the surrounding peri-


106

implant area activates the osteoclasts for bone resorption or inhibits the osteoblasts for bone

formation. These phenomena are the response to mechanical stimuli, and any factor that

affects these stimuli can affect bone formation and/or resorption. Sr can replace these

mechanical stimuli, as osteocytes devoid of any external stimuli enhance the osteoclast bone

resorption in vivo, as seen in our uncoated Sr-containing implant sections and reported by

others [228, 229]. Considering the site of bone resorption and the position of osteoclasts, the

applied pressure signals the osteoclasts to reduce their activity in the presence of certain

concentration of Sr, which in our study has been shown to be 5 wt%. The activated osteocytes

also produce osteoblast-stimulating factors, including BMP-2, for bone formation. Also, the

action of Sr on bone resorption can be associated with a Ca-like effect on bone, wherein high

concentrations of Ca have been shown to inhibit osteoclast activity in vitro [230] and in vivo

[231].

6.3.3 Role of Col-I in enhanced bone formation

The enhanced osteogenesis can also be associated to Col-I coating. Availability of coating on

the Mg alloy surface leads to gradual release of degrading Sr ions upon implantation. The

slow release of ions not only promotes osteoblast proliferation, but also facilitates the

precipitation of apatite, thus increasing the mechanical strength of the bone-implant interface

[216, 217]. Consistent with this interpretation, Barbara et al. (2004) [225] also found that Sr

uptake is directly proportional to Mg uptake and free Mg ions act as stimulators of osteoblast

activity [171]. In our study, the gradual release of Sr ions due to Col-I coating creates

availability of free Mg ions for longer periods throughout the implantation. This could have

contributed to superior osteoblast activity and a subsequent increase in ECM quantity in the

Sr-containing peri-implant bone area. Hence, Col-I acts as a connecting factor between the


107

alloy surface and the extent of availability of degrading ions to the peri-implant area required

for bone ECM quality and quantity.

6.3.4 BMC/BMD and bone mineralisation

Our findings, along with the previous report [171], raise the possibility that high Sr content (5

wt%) in Mg alloys may have beneficial effects on osteogenesis, with decreased bone

resorption, while maintaining enhanced new bone mineralisation, when coated with Col-I.

However, it affects not only the process of new bone formation but also the bone mineral

deposition, as seen from the histology and BMC/BMD results of DXA radiography in our

study. However, further studies need to be carried out to understand the precise impact of Sr

concentration on bone mineralisation. It would be valuable to conduct future work to estimate

the precise amount of new bone that is formed around the implant. Measuring the deposition

of Sr ions in the newly formed bone will provide us with a more accurate knowledge of Sr

incorporation in the implant site from biodegrading Mg implants. Gene expression for

osteoblasts grown in coated-alloy extracts was not quantified due to the potential effect of

coated Col-I in the growth of cells. Nonetheless, our study helps in understanding the basic

optimum elemental constituents required in Mg alloys for their application as orthopaedic

implants.

6.4 Summary

Our work provides experimental evidence that Mg-based alloys containing high-Sr (5 wt%)

and coated with Col-I, enhances the rate of osteogenesis and reduces bone resorption over

short term implantation, resulting in increased bone mass. This gives rise to increased

trabecular bone volume without altering mineralisation process of the newly formed bone.

Such alloys can help in the prevention of bone loss and increase bone mass and maturity.


108

This work confirms that prolonged exposure of Sr to the implant site results in better bone

formation and lower bone turnover, without any deleterious effects of long-term Sr exposure

on bone mineralisation. We also confirm the effect of Col-I coated high-Sr containing Mg-

based alloys in bone resorption, by altering the in vivo osteoclast properties. Our data

describe a positive effect of Col-I coated high-Sr containing Mg alloy (Mg-2Zr-5Sr) on

osteoblasts and mineral deposition and negative effect on osteoclasts and bone resorption.

Chapter 7. General Discussion

109

Chapter 7

General Discussion

7.1 Highlights of Research

The potential advantages of Mg-based materials over other non-resorbable biomaterials,

especially for orthopaedic applications, are obvious. However, in spite of significant research,

there remain challenges with the successful implementation of Mg-based materials in a

variety of dental and orthopaedics applications. Many of these challenges are related to

corrosion and early degradation. This chapter discusses the highlights of this work in favour

of Mg biomaterials, in the context of alloying, surface modification and osseointegration, as

implant material in orthopaedics.

7.1.1 Impact of Mg alloying elements in bone formation and osseointegration: the

requirement to alloy Mg- strength versus toxicity

Pure Mg does not provide the necessary mechanical and corrosion properties for implant

applications. Therefore, potential alloying elements need to be carefully considered. Of all

the available alloying elements to date including Al, Zn, RE, Li, and Mn, we chose to alloy

Mg with Zr, Ca and Sr because of the advantages described in chapters 4, 5 and 6. Since our

studies were typically short-term, the results may have been influenced directly or indirectly

by corrosion of the alloy itself, especially in in vitro tests. In such cases, it is practically

impossible to isolate the effect an increased corrosion rate might have had on the perceived

toxicity of the investigated alloy. While still providing the requisite mechanical properties,

these alloying elements can provide mechanical or corrosion benefits when alloyed with Mg.

Sr proved to be the most biocompatible Mg-alloying element, although studies using other

concentrations of Sr should continue (especially over long-term usage) before they can be


110

developed fully. This has been suggested by the formation of the Mg17Sr2 intermediate phase

in Sr containing Mg based alloys, which improves their corrosion resistance. In our study, Ca

or only Zr containing Mg alloys were unstable and upon degradation in vivo, they showed a

lot of residual alloy material at the implant site. This could also lead to a preliminary

conclusion that formation of the Mg17Sr2 intermediate phase gives in vivo stability to the

alloys and leads to osseointegration of Mg-Zr-Sr implants.

7.1.2 Impact of Col-I coating on osseointegration and implant stability

An area that has recently received wide attention in the field of Mg biomaterials is the study

of surface modification methods to minimise their degradation rates. For coating of

biodegradable Mg alloys, the appropriate criterion is that they are biocompatible and fully

degradable. This becomes important for understanding what occurs to the implant material

itself over the implant life cycle. In the case of biodegradable materials like Mg-based alloys,

coating cannot be a solution to corrosion as the alloy itself has to be accessible to the

surrounding bone tissue for new bone formation and its regeneration. Ideally, the coating

should itself degrade gradually, helping to control the overall corrosion process while leaving

no harmful traces. The Mg alloys used in this study has been functionalised with Col-I

protein for assisting in bio-acceptance of these alloys, if not to improve the corrosion

properties. Being a biomolecule, Col-I does not play an important role in preventing

corrosion of Mg-based alloys; however it could enhance the osseointegration and stability of

the implanted alloys for a longer period, until complete healing of the bone is observed.

Interestingly, it could increase the proximity of new bone formation to the implant surface for

Ca-containing Mg alloys (Chapter 5), affecting the secondary stabilisation of implanted

alloys and stimulating mineral deposition in Sr-containing Mg-based alloys (Chapter 6).


111

7.2 Work Required to Connect the Gaps

While effective designing of materials is the standard criteria in engineered bone substitutes,

it is also of utmost importance for a clinically relevant device to integrate biologically during

its existence with the host tissue. This section summarises the points that connect the gaps

between defining the precise role of Mg-based alloys in bone tissue-engineered alternatives

and other available materials used in bone grafting techniques and recent advances in

biomaterials research.

7.2.1 Live cell imaging for studying the cellular dynamics to degrading alloys

To understand the influence of degrading alloys on cellular behaviour, the cellular dynamics

(size, polarity, movement) and nuclear dynamics (size, circularity, movement) and the

temporal changes in these two aspects of the cells need to be studied. This can be done by

tagging the focal adhesion complexes of the cells with fluorescent proteins and use as

markers for quantitative analysis of the adhesion sites and rates following time-lapse

experiments.

7.2.2 Measurement of in vivo corrosion or biodegradation rate

There are many in vitro corrosion measurement methods described in the literature, including

quantification of pH values, Mg ion release, hydrogen evolution, mass or volume changes,

and potentiodynamic polarisation [232]. Several reports have included measurement of Mg

degradation in commonly used inorganic media with their inorganic ingredients similar to

those of plasma [233], such as phosphate buffer solution (PBS), hanks balanced salt solution

(HBSS), and simulated body fluid (SBF) and cell culture media (e.g., DMEM) [234].

However, the difference between the environments under in vitro and in vivo conditions

suggest probable difference in the degradation behaviour of the alloys, owing to higher


112

deposition of ions in the vicinity of the degrading alloy under in vitro conditions. Due to this,

measurement of the degradation rate under in vivo conditions, after the alloy has been

implanted into the host tissue, becomes necessary. Micro-computed tomography (μCT)

analysis is a necessary criterion in this aspect as an established monitoring protocol to

compare the volume changes of the implanted alloys to assess their corrosion rates in vivo. It

is appropriate to compare the densities of implants while they are still inside the host tissue,

as the corroded portion has lower substrate density than its residual part. pH measurement is

also a user-friendly and a quick qualitative assessment that can be applied for monitoring the

corrosion rates through the determination of OH– concentration in the extracts, urine being

the target extract under in vivo conditions, as the dissolution of Mg is accompanied by the

release and precipitation of OH–.

7.2.3 Measurement of bone bonding strength

Bone bonding is an important feature that describes the strength of the newly formed bone on

the cement line matrix. The understanding of bone bonding can be elucidated from the

resorption surface created by osteoclasts at the BRC (Fig 2.8) and the initial matrix synthesis,

which occurs during the de novo bone formation. This can be achieved by focusing on the

structure of the resorbed bone surface and the cement line-ECM that occupies the interface

between new bone and old bone during the process of bone remodelling.

Chapter 8. Conclusions and Recommendations for Future Work

113

Chapter 8

Conclusions and Recommendations for Future Work

8.1 Conclusions

Mg alloys as biodegradable implants are promising replacements for currently available

implant materials and have become popular due to advancements in technology, allowing

improved control of corrosion. Although this has led to success in vascular applications,

some of the most important considerations regarding biological responses are required to

ensure the continual and effective development of Mg based biomaterials for clinical

purposes. In many cases, it is realistically impossible to determine the effect an increased

corrosion rate might have on the perceived toxicity of the investigated alloy. This thesis

addresses the biological and cellular properties imparted by Mg-based alloys upon

implantation into a rabbit model, and investigates the role of Mg and its alloying elements in

cellular behaviour and their response. In summary, the following conclusions can be drawn

from this thesis:

1. Effect of zirconium, calcium and strontium contents on implant-induced

osseointegration

All four alloys, namely Mg-5Zr, Mg-5Zr-Ca, Mg-2Zr-5Sr and Mg-Zr-2Sr, demonstrated

mechanical properties close to natural bone. Notable differences in contact angles and

calculated surface free energies were seen for all alloy surfaces, Mg-Zr-2Sr being the most

hydrophobic and having the lowest surface energy. The inclusion of Sr gives Mg-Zr implants

the property of contact osteogenesis. The addition of Sr also notably increases early bone

apposition in rabbit cancellous bones. The optimum concentration for the best

osseointegration of Sr for Mg-Zr-Sr alloys has been found to be 2wt%. The influence in the


114

rate and extent of osseointegration is probably due to the physico-chemical changes brought

about by the change in surface energy of Sr-containing Mg-Zr alloys. Hence, Zr and Sr have

a positive impact on the mechanical and biological properties of Mg-Zr-Sr alloys. Ca has a

negative impact on osseointegration because of its rapid corrosion property. The surface

topography of Mg-Zr-Sr alloys is compatible as in vivo implants and these materials can

generate new bone when used as bio-implants. It also helps in complete degradation of the

implant within the specified period of surgery

2. Collagen type-I leads to in vivo matrix mineralisation and secondary stabilisation

Col-I coating of Ca-containing Mg-Zr based alloys enhances their in vivo biocompatibility

and osteogenesis capacity. The mechanical properties of Mg-5Zr and Mg-5Zr-Ca alloy

implants are comparable to that of natural bone, as demonstrated from their phase

constituents, microstructure and compressive strengths. The negative impact of Mg-5Zr-Ca

on osseointegration and implant stability could be improved by surface coating with Col-I,

which helps in surface fixation and secondary stabilisation. Col-I coating enhances the

surface energy and surface hydrophobicity of these alloys and strongly influences the protein

binding capacity onto the alloy surface, leading to better osteoblast activity. Col-I coating

enhances the rate of osseointegration with good levels of new bone formation after only a

month of implantation, as demonstrated by histology, immuno-histochemistry and radiology.

Our results provide new information on the quality of bone mineralisation and suggest an

earlier onset of bone remodelling process compared with the uncoated counterpart alloys.

3. Strontium content and collagen-I coating influence osteogenesis and bone resorption

For developing new Sr-containing Mg-Zr based biodegradable implants with high

osteogenesis and mineralisation properties, the Sr levels should be ≥5 %. Our results confirm


115

that Mg-based alloys containing high Sr content, which in our study is 5wt%, and coated with

Col-I reduces bone resorption and increases bone formation and osteogenesis during shorter

periods of time. Lower yet prolonged exposure of Sr to the implant site, by having high

concentration of Sr and surface coating with Col-I, has helped in better bone formation and

lower bone turnover. This enhancement in bone formation is followed by the absence of any

negative effect of long-term exposure of the degrading particles on bone mineralisation. Such

alloys can help in prevention of bone loss and increase in bone mass and bone strength. We

also confirm the effect of Col-I coated Mg-2Zr-5Sr alloy in bone remodelling by altering the

in vivo osteoclast properties. Further, the enhanced osteogenesis due to Col-I coating affects

primarily the initial stages of bone healing. This can indicate decreased susceptibility for

infections and earlier initiation of wound recovery.

8.2 Recommendations for Future Work

One of the most demanding challenges towards systemic development of future Mg

biomaterial is related to the fact that a strong knowledge of material science, corrosion

engineering, and biological interactions is required to fully understand the underlying

biology. The following directions can be looked into for future focus in Mg biomaterials.

1. The in vitro tests are done in closed conditions and are devoid of dynamic environment

like the human body. As a result, the conditions alter rapidly due to which many processes

are affected, including corrosion, pH changes, the degrading metal ion concentration,

solubility of the corroding products and hydrogen gas evolution. Consequently, there is a vast

difference between the in vitro and in vivo results. To make a more realistic conclusion from

our studies, prolonged in vitro tests should be performed under conditions that can mimic

physiological conditions and which can be compared in parallel with the in vivo studies.


116

2. The reliability of long-term in vitro and in vivo studies in terms of basic aspects, such as

toxicity imparted by the degrading alloying elements, has to be addressed to date. Without

this information it may be premature and unethical to conclude any kind of short term as well

as long term in vivo studies related to biological and cellular response.

3. The long-term in vivo study of the degradation and wear resistance of metals can be

associated with pseudo-tumours. The formation of pseudo-tumours can be seen as soft-tissue

masses relating to post-surgery effect due to the elevated levels of metal ions in serum,

especially when the implanted alloys have minimal wear-resistance, as in the case of Mg

alloys. Although excessive accumulation of Mg ions in the body is almost impossible as they

are excreted out from the body, the degradation of the alloys may lead to accumulation of

other elements that have been used as alloying agents to strengthen the corrosion resistance of

Mg alloys. Pseudo-tumours may involve a cytotoxic response and a delayed hypersensitivity

response to the degrading alloy particles. A long term study of the implants in higher animals

that are more phylogenetically related to humans would give us a better idea of the tumour

response of host tissue to the degrading alloy particles.

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