Osseointegration of novel silver-doped hydroxyapatite

coated and acid-etched titanium implants in an ovine

model

Dasun Abeygoonawardana

A thesis in fulfilment of the requirements for the degree of

Master of Science

Prince of Wales Clinical School

Faculty of Medicine

March 2017

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Date Abstract

Despite advances in technology and surgical procedures, periprosthetic infection of orthopaedic implants remains a common and costly cause of implant failure and revision. Novel materials and coatings have been developed in an effort to improve implant resistance to the onset of infection, including those utilising the antimicrobial properties of silver ions. The aims of this study were to evaluate the osseointegration of a novel silver-doped hydroxyapatite coating and a second acid-etching process, applied to a titanium implant substrate implanted using a well-established ovine model. The subsequent results from the silver-doped coating could be compared to conventional hydroxyapatite coatings, to evaluate the effect of silver ions upon potential bone ongrowth. In this study, the silver HA coated implants demonstrated significantly higher shear stresses of mechanical pushout when compared to the acid etched implants after 4 weeks and 12 weeks in situ. Similarly, the bone-implant contact percentage was significantly higher for the silver HA coated implants indicating greater osseointegration at both timepoints. When compared to conventional HA coated implants in an identical ovine model, there was no significant difference observed in mechanical pushout or bone- implant contact of the silver HA coated implants. These results indicate that the novel silver-doped HA coating does not adversely affect the osseointegration of titanium implants. Therefore, this represents a viable solution for an antimicrobial implant coating which continues to facilitate effective osseointegration following implantation.

ii Table of Contents

List of Figures ...... v

List of Tables ...... vii

Acknowledgements ...... viii

Chapter 1. Introduction ...... 9

1.1 Infection and revision of orthopaedic implants ...... 9 1.2 Economic impact of prosthetic infection ...... 10 1.3 Mechanisms of prosthetic infection ...... 11 1.4 Antimicrobial properties of silver ions ...... 14 1.5 Use of silver in current implant technology ...... 15 Chapter 2. Method ...... 18

2.1 Study Design Summary ...... 18 2.2 Implant Manufacture ...... 19 2.3 Animal Preparation ...... 21 2.4 Cortical Site Implantation ...... 21 2.5 Cancellous Site Implantation ...... 23 2.6 Post-operative Recovery ...... 24 2.7 Harvest and Sample Preparation ...... 25 2.8 Mechanical Testing ...... 28 2.9 SEM Sample Preparation ...... 30 2.10 SEM Imaging...... 32 2.11 Histomorphometry ...... 33 2.12 Histology ...... 34 Chapter 3. Results ...... 35

3.1 SEM Coating Characterisation ...... 35 3.2 Surgery and Harvest ...... 37 3.3 Radiography ...... 38 3.4 Mechanical Testing ...... 43 3.5 SEM Imaging...... 48 3.6 Histomorphometry ...... 48 3.7 Histology Findings ...... 51

iii Chapter 4. Discussion ...... 55

4.2 Comparison of results with other ovine model studies ...... 57 4.3 Effect of surface topology upon bone ongrowth ...... 61 4.4 Current developments of silver in implants...... 68 4.4.1 Functionalised HA Coatings ...... 68 4.4.2 Silver-doped Strontium coatings...... 72 4.4.3. Agluna® material ...... 76 4.4.4 PorAg® coating ...... 80 4.5 Silver concentrations of current implants and coatings ...... 81 4.6 Ag HA compared to conventional HA ...... 83 4.7 Study Limitations and Future Studies ...... 87 Chapter 5. Conclusions ...... 90

References ...... 92

Appendix A Histomorphometric Results ...... 97

Results from W2602...... 98 Results from W2603...... 99 Results from W2604...... 100 Results from W2605...... 101

iv List of Figures

Figure 1 Three cortical implants in tibial diaphysis of animal W2603 showing implant flush with bone surface ...... 22 Figure 2 Implantation overview of ovine model showing four implantation scenarios in cancellous sites and line-to-line implantation in bicortical sites [54] ...... 23 Figure 3 Step drill used in surgery to provide a range of implantation scenarios within a single site ...... 24 Figure 4 Isolated femur and tibia (both left and right) from animal W2603 ...... 26 Figure 5 Isolated axial segment from tibial diaphysis containing cortical implant ...... 27 Figure 6 Sample W2602L4 following sectioning, showing medial half (W2604L4-M) ...... 28 Figure 7 Raw SEM image of Specimen W2604R5L at 80x magnification. Implant (blue arrow) is displayed as white, while bone (red arrow) is displayed as grey. Voids (circled) are represented as darker regions...... 32 Figure 8 SEM images of Specimen W2604R5L showing raw (top) and processed (below) versions. Green represents regions of bone contact, red represents bone voids. Total bone contact is 85.2%, as indicated in top left corner of processed image...... 33 Figure 9 SEM imaging of acid etched surface (Group 1) (x50 magnification). Texturing and evenly distributed ridges achieved through acid-etching process visible ...... 35 Figure 10 SEM imaging of Ag HA (Group 2) coating (x50 magnification). Unique surface of plasma-spray coating visible in the form of compact undulations and microstructures...... 35 Figure 11 SEM imaging of acid etched surface (Group 1) (x1000 magnification). Network of ridges, crevices and nanopits visible at this level of magnification...... 36 Figure 12 SEM imaging of Ag HA coating (x250 magnification). Small voids and porous regions visible between the plasma-sprayed structures of the Ag HA coating...... 36 Figure 13 Harvested tibia from W2605 showing visible cancellous (blue arrow) and cortical (red arrow) implantation sites. Not evidence of foreign body reaction or implant rejection surrounding implants in situ ...... 38 Figure 14 Left limb radiographs of animal W2602, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 39 Figure 15 Right limb radiographs of animal W2602, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 39 Figure 16 Left limb radiographs of animal W2603, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 40 Figure 17 Right limb radiographs of animal W2603, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 40 Figure 18 Left limb radiographs of animal W2604, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 41 Figure 19 Right limb radiographs of animal W2604, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 41 Figure 20 Left limb radiographs of animal W2605, AP(left) and lateral (right) views, showing good implant alignment at all sites ...... 42

v Figure 21 Right limb radiographs of animal W2605, AP(left) and lateral (right) views. Implant alignment was sufficient at all sites...... 42 Figure 22 Force-displacement curve output from custom Matlab script. Green line indicates stiffness (N/mm) taken from initial linear portion. Shaded blue region indicates integral area calculated for energy to failure (J). Peak load is the force value corresponding to junction between blue and white shaded regions...... 44 Figure 23 Mean shear stress of pushout at 4-week and 12-week timepoints for acid-etched and Ag HA implant groups. Error bars indicate standard deviation. Ag HA implants had significantly higher shear stresses at both timepoints...... 47 Figure 24 10x magnification of W2605L5L, showing Haversian systems (arrowed) and dense woven bone (circled) in proximity of Ag HA coating ...... 51 Figure 25 10x magnification of W2605R5M, showing rounded and enlarge osteoblasts layered upon newly formed bone (arrowed) at various sites ...... 52 Figure 26 Newly formed bone on lateral implant face and internal thread (arrowed) in Ag AH group ...... 53 Figure 27 Increased frequency of voids (arrowed) and reduced bone-implant contact (boxed) in acid-etched implant group ...... 54 Figure 28 Mean Shear Stress of cortical implants in ovine model at 4 week and 12 week timepoints. Results from current study highlighted in purple (4 weeks) and green (12 weeks). Acid-etch titanium had lower shear stresses than all coated titanium substrates...... 59 Figure 29 Mean Bone Implant Contact of cortical implants in ovine model at 4 week and 12 week timepoints Results from current study highlighted in purple (4 weeks) and green (12 weeks). BIC for PS Ti** was markedly lower than other plasma-sprayed variants...... 60 Figure 32 Mean shear stress of pushout at 12 weeks in ovine model of Ag HA and Conventional HA coatings ...... 86 Figure 33 Mean BIC in cortical bone at 12 weeks for Ag HA and Conventional HA coatings ..... 86

vi List of Tables

Table 1 Study design indicating group (Group 1: Acid-etched, Group 2: Ag HA) and timepoint per implantation site. Femur and Tibia indicate cancellous sites. Cort 1, 2 and 3 indicate cortical sites in descending order from the top of tibia...... 19 Table 2: 4-week Mechanical Testing Results showing calculated p-value for each parameter . 45 Table 3 SPSS Output of Independent t-test for Mechanical Testing at 4 weeks. All parameters were found to be significantly different (p<0.05) between the groups (red boxes), excluding the average cortical thickness of the tested specimens (blue box). This indicates variations can be attributed to the quality bone-implant interface...... 45 Table 4: 12-week Mechanical Testing Results showing calculated p-value for each parameter 46 Table 5 SPSS Output of Independent t-test for Mechanical Testing at 12 weeks. All parameters were found to be significantly different (p<0.05) between the groups (red boxes), excluding the average cortical thickness of the tested specimens (blue box). This indicates variations can be attributed to the quality bone-implant interface...... 46 Table 6 Bone ongrowth evaluation of cortical specimens. Mean percentage ongrowth of the Ag HA implants was consistently higher than that of acid-etched implants at both timepoints. ... 49 Table 7 ANOVA Output for cortical samples, testing bone ongrowth at all timepoints. BIC is dependent upon implant type and timepoint, but variation between groups is independent of time in situ ...... 49 Table 8 Bone ongrowth evaluation of cancellous specimens. Mean percentage ongrowth of the Ag HA implants was consistently higher than that of acid-etched implants at both timepoints...... 50 Table 9 ANOVA Output for cancellous samples, testing bone ongrowth. BIC is dependent upon implant type, but independent of timepoint, and variation between groups is independent of time in situ ...... 50 Table 10 Comparison between current study and previous HA study in same ovine model ..... 84

vii Acknowledgements

I would like to thank Dr. Matthew Pelletier for his constant guidance and expertise throughout the entire duration of this ordeal research project. I am also indebted to Dr. Bill Walsh and Dr. Nicky Bertollo who lent their time, advice and priceless experience to many aspects of this work. I am extremely grateful to Dr. Declan Brazil for going out of his way to make this whole thing possible.

I can’t even begin to express my appreciation for the unwavering support of

Santhara and my family, so I’ll just leave it at that.

Most of all, I want to acknowledge my Seeya, who wanted this for me more than anybody. I know he would have been equal parts thrilled and relieved to see it finally happen.

viii Chapter 1. Introduction

1.1 Infection and revision of orthopaedic implants The popularity of orthopedic implants in surgical interventions is expected to continue its steady increase seen during recent decades, with the demand for primary total hip and knee and knee projected to reach 572,000 and 3.48 million procedures in the USA alone, by 2030 [1]. Similarly, revision surgeries associated with these primary procedures are also set to grow by up to 601%, based on 2005 rates [1]. The economic drains and impact upon patient health associated with a revision procedure has been well documented [2], and as such the developments aimed at improving primary implant success represent a unique and significant challenge in the field.

In Australia, 17.1% of revisions of primary conventional total hip replacements in 2013 were due to infection (n = 1534) [3]. These rates were even higher in primary total knee replacements, with 22.2% of revisions in 2013 due to infection (n = 3038) [3]. Similar trends were observed in the United Kingdom, with 12% of all hip revisions (total 10,040) and 22% of all knee revisions (total

6,009) occurring as a result of infection in 2012 [4].

Although infection represents a major cause of Total Joint Replacement (TJR) revision, the actual incidence of infection as a proportion of all joint replacement procedures remains low. An analysis of NIS data in the USA from

1990 – 2004 conducted by Kurtz et al [5] recorded infection rates of 1.04% for both hip arthroplasties and knee arthroplasties, despite the number of knee arthroplasty revisions being higher (5,838 vs. 3,352 revisions). This calculation did not include arthroplasty devices explanted as the first stage in 2-stage revision procedures to treat infected joints. In 2001, the infection burden 9 (defined as the ratio of revisions for infection to the total number of arthroplasties) was calculated at 1.99% for total hip and 2.05% for total knee arthroplasties. The study data was updated following the release of more recent

NIS data in 2010. By this time the infection burden rates had increased to

2.21% for total hip arthroplasty and 2.32% for total knee arthroplasty, with an average infection burden rate from 2001-2010 of 2.20% for total hip and 2.25% for total knee arthroplasties [6]. The increase may partly be explained by a significant decline in length of hospital stays, which reduces the likelihood of early identification of infection during the initial post-operative stay.

The trends noted in the NIS data reflect similar incidence rates in other global centers. Data from the Australian Joint Registry indicates a cumulative incidence of revision due to infection of approximately 0.6% following primary total hip replacement and 1.0% following primary total knee replacement, up to

13 years after the procedure [3] with further studies recording infection rates within this 0.6% - 2.25% range [7-10]. Infection may occur at any period through the life of the implant, however it is generally observed that the vast majority of infections develop within the one to two years postoperatively [8-

12].

1.2 Economic impact of prosthetic infection Despite the low incidence of infection when compared to the total number of joint replacement procedures performed, treatment and revisions for infected prostheses are a costly, resource intensive exercise, particularly when compared to revisions for other causes such as loosening. The financial burden associated with infection of total joint replacements has been well documented.

10 Klouche et al [13] performed a retrospective analysis of 424 primary THA, 57 non-infected THA revisions and 40 THA revisions due to infection. The study examined the medicosurgical costs of each treatment, considering human resources, prescriptions, medicotechincal and service costs. Revision of the infected THA cost 3.6 times more than the primary procedure, whilst an aseptic revision was 1.4 times the primary procedure. Revisions for septic devices also required longer hospital stays (mean 30.6 days) when compared to an aseptic revision (8.9 days). Similar results were obtained in a study by Peel et al [14], examining both hip and knee procedures, which demonstrated the cost of managing a prosthetic joint infection case was 3.1 times the cost of the primary arthroplasty. Treatment and management of infection entailed greater number of re-admissions, more additional surgery, longer hospital stays (31.6 days vs.

7.9 days), with a mean cost of $69,414 compared with $22,085. Similar increases in financial burdens are documented in economic studies by Kapadia et al in the US and Garrido-Gomez et al in Spain [15, 16]. It is estimated that in the UK alone, the cost of infected revisions to the NHS is in the order of £200 million per annum [17].

1.3 Mechanisms of prosthetic infection Given the devastating impact of prosthetic joint infection on patient outcomes, and the difficulty and cost of treatments or revision, significant developments have been made to create strategies aiming to prevent infections in the surgical setting. This includes advances in operating standards, patient isolation to confine pathogenic strains, control of potential personnel and environmental sources of contamination during surgery, as well as peri-operative prophylaxis to combat infection establishment [18, 19]. Whilst effective, these strategies do 11 not directly address the ongoing issue of bacterial adhesion to implant materials, which serves as the genesis of periprosthetic joint infections.

The mechanisms of periprosthetic joint infection differ from other forms of systemic infection, due to the presence of an inanimate biomaterial substrate.

Whilst biomaterials are not biologically inactive at an atomic level, they are inanimate and therefore present a surface susceptible to colonisation by opportunistic bacteria [20]. The surrounding regions of an implant are known to have local immune depression, which further increases the vulnerability to microbial colonization [21, 22]. In orthopaedic applications, damaged or traumatised tissues caused by wear particles or friction at the implant surface are also, by extension, inanimate substrates susceptible to microbial colonisation, and subsequent development of infection [20, 23]. The presence of a foreign material has also been shown to dramatically reduce the dose of contaminating microorganisms required to cause infection at the site, up to a factor of 100,000 times in one animal study [24-27].

Generally, organisms responsible for infection originate from the skin flora of the patient or physician during the procedure, which enter the open wound and migrate through incision channels to the implant surface [22, 28]. The microorganisms most commonly responsible for infection have been identified as coagulase-negative staphylococci (CoNS) and S. aureaus, which together reportedly account for approximately 50-65% of infection cases [26]. An evaluation by Campoccia et al [23] of nearly 800 clinical isolates from prosthetic infections since 2000 found that roughly four out of five infections were caused by staphylococci, whilst S. aureas and S. epidermidis specifically accounted for

12 34% and 32% of infection isolates respectively.

Implant infection is instigated by initial bacterial adhesion to the biomaterial surface [22]. Surfaces exposed to the biological environment immediately acquire a film of proteins, which attract bacteria as a source of nutrition.[20,

22]. Subsequent prosthesis-specific infection requires active interaction between the biomaterial surface and the bacteria. This can include the bacterial production of surface adhesins which attach to the protein films on the implant, or production of a polysaccharide “slime” to create an adhesive biofilm on the implant surface [22, 23, 28-30]. Biofilm contains a multi-level community of bacteria more protected from the host immune system, and which are less susceptible to conventional antimicrobial agents [22, 23, 28]. Isolation of biofilms on infected hip prostheses has identified CoNS, S. aureaus as well as other microorganisms present [31].

The unique pathogenesis of prosthetic joint infection has driven the development of preventative strategies targeting the tissues surrounding the implant/biomaterial interface. This site represents the area of initial bacterial adhesion to the implant, which is a crucial step in the development of infection

[23]. Current developments focus upon alteration of the outer-layer chemistry or surface topology of the biomaterial, to disrupt initial microbial adhesion whilst maintaining the required mechanical and biocompatible properties of the material. Potential surface modifications include coatings with surfactants or proteins to create adhesion resistant surfaces, as well as bioactive coatings such as chitosan which have large-spectrum antimicrobial properties [32, 33]. The use of a coating doped with silver ions is another example of a bioactive coating,

13 which facilitates controlled elution of metal ions into the interstitial space to help prevent microbial adhesion to the implant surface.

1.4 Antimicrobial properties of silver ions The broad uses of silver ions against microbial action has been utilised for centuries, with current uses in a range of applications including wound dressings, urinary catheters and water systems[34, 35]. Silver is often described as oligodynamic, meaning it is able to generate a bactericidal effect at minute concentrations., with reports of bactericidal activity at concentrations as low as 35 parts per billion [36, 37]. The Ag+ ion is an effective antimicrobial agent as it remains non-toxic to human cells [38] whilst displaying antiseptic properties [39]. When used as a bioactive coating, silver ions have been shown to reduce initial bacterial adhesion and colonisation upon the implant surface, thereby disrupting the development of biofilm [40, 41]. Chaw et al [42] demonstrated that even low concentrations of silver, whilst unable to exhibit antimicrobial effects within biofilms, were capable of destabilizing the biofilm matrix of S. epidermidis bacteria. Similar results have been recorded by Stobie et al [41].

Additionally, controlled elution of ions from the coating can provide an antimicrobial effect in the interstitial joint space surrounding the implant, with the ions performing a bactericidal function. The mode of action of the silver ion is multifaceted, but generally believed to rely upon binding of the Ag+ cation to the negatively charged bacterial cell wall, denaturing the membrane and ultimately leading to loss of structural integrity and function, cell lysis and death [37, 43, 44]. Silver ions are able to bind to molecular groups of enzymes in

14 the cell wall, thereby disrupting key cell metabolic processes such as respiration or ion transport [45]. Once within the bacterial cell, silver ions can further bind to proteins and inhibit cell DNA synthesis and function [46]. The affinity of silver ions to proteins has lead to suggestions that impregnation of silver ions into a coating may be more effective than a direct surface coating alone, as surface silver ions can be readily deactivated by the surface proteins of a device following implantation [41, 47].

Excessive amounts of silver compounds or long term treatment with silver can lead to argyria, a discolouration of skin or tissues, which is otherwise not harmful [48]. Further, silver toxicity or argyrosis can be resolved with the cessation of therapy, which will occur naturally in the case of elution’s from silver coatings on implants [49]. Levels of silver concentration associated with argyrosis are reported to range from 4-6g silver, while in comparison Gosheger et al [50] estimates the silver concentration required for coating a total femoral replacement to be approximately 1.15g, indicating the risk of developing systemic argyrosis is unlikely.

1.5 Use of silver in current implant technology In vivo studies of silver-coated implants have supported the use of silver ions to reduce incidence of periprosthetic infections. An animal study by Gosheger et al

[50] examined the in vivo antimicrobial efficacy of silver coated megaprostheses, and the toxicological side effects of the coating. In this study

30 rabbits were implanted with either a titanium or silver-coated titanium endoprostheses, and subsequently infected with S. aureus. Silver coatings had

99.7% purity, and were applied by galvanic deposition to a layer thickness of

15 10- 15 µm, with a 0.2 µm thick layer of gold to act as a cathode driving the release of Ag+ ions. At 90 days postoperative follow-up, the silver group showed significantly lower infection rates (7% vs. 47% p<0.05) when compared to the titanium group. A second group of 10 rabbits were implanted with silver-coated endoprostheses to evaluate the toxicological side effects of the coating. The results from this group showed elevated silver concentrations in blood and the organs, with no pathological changes in laboratory parameters, or histological change in organs. This indicates that the silver-coating is able to successfully reduce infection rates without adverse toxicological side effects. A similar study by Collinge et al [51] utilised stainless steel pins coated with silver, and implanted in sheep infected with S. aureaus. At 19 days follow-up, 84% of the uncoated pins were infected, compared with 62% of the coated pins.

Additionally, motion of the implants was found to correlate highly with infection at the site, with silver-coated implants displaying much lower quantities of biofilm when examined under electron microscopy.

Early outcomes of the use of silver ions in clinical settings have shown promising results with respect to infection control. A retrospective evaluation conducted at the Royal Orthopaedic Hospital in Birmingham, UK examined the incidence of early periprosthetic infection in high-risk patients treated with silver-treated custom endoprostheses [52]. 170 patients were evaluated in two equal groups, one with silver treated implants and the other a control group, with data collected at 3, 6, 9 and 12-month post-operative follow-up visits. Post- operative infection rates were halved in the silver treated group (11.8% vs.

22.4%). Additionally, treatments for infection were more effective in the silver

16 treated group than controls (success rates with DAIR were 70% vs, 31.6%).

Therefore, mid-term results of the study associated silver-treated endoprostheses with lower rates of early periprosthetic infection, and superior outcomes for subsequent treatments and revisions.

17 Chapter 2. Method

2.1 Study Design Summary Implants were evaluated for osseointegration using a standard bilateral ovine model developed by Walsh and Bruce [53]. This model facilitates both bicortical implantation sites at the tibial midshaft, and cancellous implantation sites at the distal femoral condyle and proximal tibia, within a single animal. A total of

40 implants divided into 2 implant groups were evaluated in this study. Group

1 comprised of the acid etched titanium implants, whilst Group 2 consisted of the silver-doped hydroxyapatite (Ag HA) coated implants. An overview of the study is designed, including implant group used at each site, in presented in

Table 1.

Osseointegration was evaluated at 4-week and 12-week time points. Each time point consisted of 2 animals, for a total of 4 animals in this study. The bilateral ovine model utilises the rear legs of the animal for implantation. Each leg provides 3 bicortical implantation sites and 2 cancellous implantation sites.

Therefore, utilising 2 animals per time point allows 12 bicortical sites (n=12 specimens per group) and 8 cancellous implantation sites (n=8 specimens per group) to be evaluated at each time point.

Implant evaluation at each timepoint comprised of mechanical testing and histomorphometric analysis. Mechanical testing was performed on cortical samples to evaluate the shear stress of pushout from cortical bone. This provides an indication of the strength of the new bone/implant interface and therefore the success of osseointegration of the implant. Histomorphometric analysis allowed the quantification of bone ongrowth along the implant

18 perimeter to compare the extent of osseointegration and bone adhesion in each group per timepoint.

Table 1 Study design indicating group (Group 1: Acid-etched, Group 2: Ag HA) and timepoint per implantation site. Femur and Tibia indicate cancellous sites. Cort 1, 2 and 3 indicate cortical sites in descending order from the top of tibia.

Left Leg Right Leg

Animal Femur Tibia Cort1 Cort2 Cort3 Femur Tibia Cort1 Cort2 Cort3

W2602 1 2 2 1 2 1 1 1 2 1

W2603 1 2 2 1 2 1 1 1 2 1

W2604 1 2 2 1 2 1 1 1 2 1

W2605 1 2 2 1 2 1 1 1 2 1

Shaded cells indicate 4-week timepoint. Remaining cells are 12-week timepoint.

2.2 Implant Manufacture Two titanium implant groups were evaluated for osseointegration in this study.

All implants consisted of a Ti6Al4V substrate, in the form of a cylindrical dowel

20mm in length. Dowels intended for Group 1 (acid etching) had a dimeter of

6mm, whilst dowels intended for Group 2 (Ag HA coating) had a diameter of

5.6mm. The 0.4mm reduction in diameter of the Group 2 implants was intended to allow the application of the Ag HA coating to a target thickness of

0.2mm, such that final implant diameter remained constant across both groups.

A standard M4 metric thread was tapped longitudinally into all dowels to a depth of 5mm, which facilitated the attachment of instruments required for the surgical implantation of the devices. A total of forty (40) Ti6Al4V dowels were manufactured for this study by Signature Orthopaedics Ltd (Sydney, Australia) before undergoing further processing, such as surface treatments, coating applications and sterilisation. 19 Group 1 (acid etched) implants were not coated, but subject to a proprietary nanotexturing process of the Ti6Al4V substrate through exposure to an acid by

KKS Ultraschall AG (Steinen, Switzerland). All external surfaces of the dowel were exposed to the acid etching process. Group 1 contained 20 implants in total.

Group 2 (Ag HA coating) implants first underwent additional processing in the form of grit-blasting of the external cylindrical surface. Subsequently, they were plasma-sprayed with a silver-doped hydroxyapatite (Ag HA) coating by

Accentus Medical (Didcot, UK) which covered all external surfaces of the dowel, excluding the cylindrical end face surrounding the threaded hole. The coating was applied to a target uniform thickness of 0.1mm through proprietary processes developed by Accentus Medical, to ensure the diameter of the final coated implant was 6mm. Group 2 also contained 20 implants in total.

Following final processing all implants were provided non-sterile by the respective suppliers. Initial SEM imaging was performed to evaluate and compare the implant surfaces prior to implantation. Imaging was completed using a Hitachi S-3400 SEM (Hitachi High-Technologies Corporation, Tokyo,

Japan), with images taken up to 1000x magnification at various points along the implant.

Implants underwent a final cleaning process at Signature Orthopaedics Ltd before sterilisation through gamma irradiation from 25-40 kGy (Steritech Pty

Ltd, Narangba, Australia). Implants were therefore provided at surgery clean, sterile and individually packed.

20 2.3 Animal Preparation Four skeletally mature sheep (cross-bred Merino Wethers, 18 months) were used in this study with ethical consent from the institutional Animal Care and

Ethics Committee. Each animal was fasted from both food and water for 24 hours prior to undergoing the surgical procedure. Animals were rendered unconscious via an intramuscular injection of Zoletil (Virbac, Carros, France), whilst they remained in the holding pen, on the day of the surgery. When the animal had lost consciousness, it was position in a left lateral position with neck extended upon on a table for subsequent intubation. An 8.5Fr endotracheal tube was introduced and the cuff was inflated with 10mLs of air, whilst the tube was fastened securely to the lower jaw of the animal. Animals were given 02 at a rate of 4L/min and 1.5-2.5% Isoflurane using an anaesthetic machine. Throughout the procedure, animals continued to receive the above, as well as 1g of Cephalothin (intravenous) and 5mLs of Bencillin as antibiotic prophylaxis (intramuscular, Troy Laboratories, Glendenning, Australia).

Animals further received 4mLs of Carprofen (Rimadyl, Zoetis, NJ, USA) and

1mL of buprenorphine (Temgesic, 0.324mg) to provide pain relief prior to the surgical procedure.

2.4 Cortical Site Implantation Bicortical implantation sites were selected along the tibial mid-shaft of the rear legs of each animal. A 3cm surgical incision was made approximately 50mm from the articular surface along the anteromedial aspect of the tibia. Following exposure of the periosteum, further dissection was performed to visualise the underlying cortical bone of the tibial diaphysis. Bicortical implants were implanted in a line-to-line scenario. Therefore, a bicortical hole was created

21 using increasing drill diameters up to a maximum of 6mm. A threaded insertion instrument and surgical hammer were then used for impaction of the implant into the implantation site. The implant was impacted until it was flush with the surface of the cortical bone, as shown in Figure 1. This procedure was repeated for the remaining two bicortical implants, with a spacing of 20mm between implantation sites. The surgical site was finally closed using resorbable sutures.

Figure 1 Three cortical implants in tibial diaphysis of animal W2603 showing implant flush with bone surface

Due to slight variations in coating thickness of the Group 2 Ag HA implants, it was observed that some Group 2 implants were more difficult to impact into the line-to-line implantation scenario than others. All implants were subject to final inspection by the suppliers prior to receipt. However, the dimensional tolerance of ±0.1mm on the thickness of the Ag HA coating may have contributed to the tactile variation in ease of insertion.

22 2.5 Cancellous Site Implantation Cancellous sites were chosen at the medial aspect of the distal femoral condyles and proximal tibia. A 1cm incision was made and muscle dissected to expose the bone beneath. Cancellous implants were evaluated in several different implantation scenarios encompassed within a single implantation site. These included press-fit (interference), line-to-line, a 1mm gap and 2mm gap either side of the implant. Multiple implantation scenarios were achieved by the use of a step drill, as shown in Figure 2 and Figure 3. This drill incorporated stepped diameters of 4mm, 6mm, 8mm and 10mm along its length, providing the required hole diameters for the implantation scenarios described. Note that the press-fit (interference) scenario was achieved by over drilling the step- drilled hole using a 5.5mm drill bit resulting in steps at 5.5mm, 6mm, 8mm and

10mm.

Figure 2 Implantation overview of ovine model showing four implantation scenarios in cancellous sites and line-to-line implantation in bicortical sites [54]

23 Implants were inserted using a threaded insertion instrument and surgical hammer for impaction into the implantation site, until flush with the surface of the cancellous bone. Following implantation, tissues were reflected and the site closed with resorbable sutures. There were no complications associated with the implantation of the cancellous implants.

Figure 3 Step drill used in surgery to provide a range of implantation scenarios within a single site

2.6 Post-operative Recovery Following completion of the surgery, animals were allowed to rest until they were deemed to be breathing independently of the anaesthetic machines. The endotracheal tube was removed when the animal was consciously chewing on the tube. Animals were initially transferred to a cage until they sufficiently recovered to attain a standing position. At this point, they were deemed fit to be transferred to pen, shared with another sheep, and with feed and water supply.

Animals were observed in their pen environment to ensure they were recovering appropriately from the surgical procedure. Any signs of infection,

24 haematoma or swelling, bleeding or ill health could indicate issues associated with the implantation. Within the pen, animals were allowed to move unrestricted and with full weight bearing, without the use of splints. Post- operative analgesic relief was provided for the first two days following the surgery, in the form of 4mls of Carprofen, SC. One animal, W2604, showed signs of a haematoma approximately 5cm in diameter on the right tibial midshaft. However, this did not appear to impede the animal’s mobility or present distress.

2.7 Harvest and Sample Preparation At the required time point, animals were euthanised through an overdose of sodium pentobarbitone (Lethabarb, Virbac, Australia) delivered by lethal injection to the jugular vein. The cull was confirmed by an absence of muscular twitches of the eyelids in response to finger clicking and other auditory stimuli.

Animals were transferred to the University wet labs where the hindlimbs of the animals were harvested through direct separation at the hip joints. The lungs, liver, kidneys and spleen of each animal were also harvested, with no evidence of adverse reactions noted. All external tissues including skin, muscles and ligaments were systematically stripped from the harvested limbs using a scalpel, in order to isolate the femur and tibia containing the surgical implantation sites, as shown in Figure 4.

25 Figure 4 Isolated femur and tibia (both left and right) from animal W2603

Following the removal of soft tissues to isolate the femoral and tibial bones, radiographs of the harvested bones were taken using a Faxitron (Faxitron,

Wheeling, IL) with digital plates (Agfa Healthcare, Mortsel, Belgium). Two X- rays were taken, in the anteroposterior and lateral views, encompassing the 5 implantation sites on each hindlimb.

Each individual implant was then isolated from the main bone in order to undergo the required mechanical testing and subsequent SEM imaging. The harvested bone was sectioned into several segments, each containing one implant. Sectioning was performed by hand using a band saw to cut axially through the bone at the required points. Cancellous implants from the femoral and tibial condyles were isolated to include the entire implant and surrounding bone. These cancellous samples were placed immediately into a glass jar and fixed in cold phosphate 10% buffered formalin, to facilitate further testing at a later date. 26 Cortical implants were initially isolated in three axial segments taken from the tibial midshaft, as shown in Figure 5. All axial segments included the bicortical ring of tibial bone, and generally elements of internal bone tissues and marrows.

Figure 5 Isolated axial segment from tibial diaphysis containing cortical implant

The axial segments of the cortical implants then underwent further processing in preparation for sectioning into medial and lateral specimens. The superior and inferior surfaces of cortical bone surrounding the implant were subject to grinding and polishing, to ensure they were parallel to implant in the sagittal plane. These parallel surfaces allowed the cortical bone to be mounted in the

Bueler low speed saw, such that the implant was fixed perpendicular to the diamond-coated wafering blade. The Bueler low speed saw was then used to section each cortical implant into medial and lateral halves, as shown in Figure

6.

27 Figure 6 Sample W2602L4 following sectioning, showing medial half (W2604L4-M)

2.8 Mechanical Testing Cortical implants were mechanically tested to evaluate the shear strength of the bone-implant interface at each time point. At the 4-week time point, mechanical testing was performed on both medial and lateral cortical implant specimens i.e. 12 samples per group. At the 12-week time point, mechanical testing was performed only on the medial cortical implant specimens, i.e. 6 samples per group. The lateral cortical implant specimens of the 12-week time point were not used in pushout testing in order to preserve the bone-implant interface for SEM analysis of the interaction between the implant surfaces and surrounding bone.

Prior to pushout testing, samples were further processed to remove periosteal bone surrounding the implant, and create a flat face of cortical bone perpendicular to the axis of the implant. This was achieved through polishing

28 of the specimen, until the outermost circular face of the implant was flush with the surface of the cortical bone. Removal of the periosteal bone was crucial in ensuring there was no residual bone impeding the progress of the implant, such that the measured force during the pushout test relates entirely to the shear interaction between implant and bone, rather than axial obstruction with periosteal bone.

Following sample preparation, the cortical bone thickness of each specimen was measured using digital callipers at two locations either side of the implant, and recorded for pending shear stress calculations.

Specimens were loaded onto the test bed of a calibrated uniaxial servohydraulic test machine (858 Mini Bionix ®, MTS Systems Inc, Minneapolis, Minnesota,

USA). Specimens were placed on a 5mm thick stainless steel support jig containing an 8mm hole, which would capture the implant upon pushout from the cortical bone. Specimens were placed face down upon the flat surface ground perpendicular to the implant, such that the implant was now vertical along its longitudinal axis. Above the specimen, a stainless steel push-out pin was attached to the actuating arm of the MTS unit. The push-out pin included a 6mm spigot designed to contact the implant’s axial circular face, and generate the pushout force. Specimen were aligned such that the implant was directly below the push-out pin and situated over the hole in the support jig, with a minimum distance of 1mm between the edge of the hole and the implant.

The MTS unit was programmed to apply a compressive axial force upon the implant via the push-out pin, at a rate of 0.5mm/min. Axial load and displacement data was captured at a sampling rate of 50Hz. The axial force was 29 applied until a peak pushout load was achieved, indicated by increasing applied displacement with decreasing load. Recorded data was evaluated by plotting the overall force-displacement curve of each specimen, to present the peak shear force during pushout. Further calculations using this data were performed to evaluate stiffness, energy to failure and proof resilience using a custom script developed for Matlab R2009a (Mathworks Inc, MA, USA).

2.9 SEM Sample Preparation All samples, both cancellous and cortical, were embedded in polymethylmethacrylate (PMMA) in preparation for SEM imaging and histology analysis. Prior to fixation in PMMA, samples were initially placed in jars of 10% cold phosphate buffered formalin. For cancellous specimens, this was done immediately after isolation of each implant from the harvested bone.

For cortical specimens, this was done following pushout testing. The 12-week lateral cortical implants which did not undergo mechanical testing were placed in the formalin immediately after sagittal sectioning into the medial and lateral halves.

After the initial 10% buffered solution, samples were transferred through a series of containers containing increasingly concentrated ethanol solutions, in order to sequentially dehydrate the specimens. Specimens were held in each concentration of ethanol for 7 days, before transfer to incrementally higher concentration from 70% to 100%. Following dehydration, specimens were polymerised in PMMA and allowed to cure for several days before further processing.

30 Embedded samples underwent further preparation to facilitate imaging via

SEM. From their embedded state, the implant surfaces were isolated to exposed along the longitudinal axis such that the bone-implant interface could be assessed. The embedded samples were initially processed using the band saw to rapidly remove material not required for evaluation including PMMA and excess bone and tissue. Care was taken not to damage the implant with the band saw blade, and to maintain sufficient PMMA for clamping in fixtures and jigs during later processing stages.

Specimens were subsequently loaded onto the Bueler low speed saw, and aligned with the wafering blade such that the implant would be sectioned along its longitudinal axis. The initial sectioning was then performed to divide the implant into two halves. Following the initial sectioning, the sample was translated 450µm to allow a second, parallel cut to be made. The end result was a section of embedded implant, with a thickness of 450µm, exposing the bone- implant interface along the longitudinal midline of the implant.

A thickness of 450 µm was chosen as it has previously been demonstrated to be an effective thickness for imaging embedded samples through SEM, without excessive distortion caused by the PMMA medium. Following sectioning, the implant face of all SEM specimens was polished using a fine grade of polishing paper to enhance the clarity of the SEM image and reduce the likelihood of artefacts or “cloudiness” during subsequent SEM imaging.

31 2.10 SEM Imaging Specimens were assessed with back scattered electron microscopy (BSEM) imaging on a Hitachi S-3400 SEM (Hitachi High-Technologies Corporation,

Tokyo, Japan). Each implant required a series of images to cover both the upper and lower perimeter of bone-implant contact regions. The number of images required varied from four to six, depending on the magnification selected to provide the clearest image for histomorphometric analysis.

Magnification was varied between 40x and 80x across all samples in both timepoints. Examples of the image series used are presented in Figure 7.

Images were saved for subsequent histomorphometric analysis.

Figure 7 Raw SEM image of Specimen W2604R5L at 80x magnification. Implant (blue arrow) is displayed as white, while bone (red arrow) is displayed as grey. Voids (circled) are represented as darker regions.

32 2.11 Histomorphometry SEM images were evaluated to quantitatively define the extent of bone

ongrowth for each specimen, as a percentage of the total bone-implant

interfacing perimeter. The SEM images highlighted metal as white, voids as

black, and bone/coating materials as shades of grey. An existing in-house script

for MATLAB R2009a (Mathworks Inc, MA, USA) was used to identify the bone-

implant interface of each image and calculate the total length of the perimeter,

using the variation in pixel colour corresponding to various structures within

the image. Regions of bone ongrowth or bone voids were manually defined by

the operator along the length of predetermined perimeter. The script then

calculated the percentage of perimeter designated to have bone ongrowth

present, as a function of the total length. Figure 8 presents an example of an

SEM image following histomorphometric analysis, with the final ratio of bone

ongrowth to total bone-implant interface presented in the top-left corner (0.825

in the example pictured).

Figure 8 SEM images of Specimen W2604R5L showing raw (top) and processed (below) versions. Green represents regions of bone contact, red represents bone voids. Total bone contact is 85.2%, as indicated in top left corner of processed image. 33 2.12 Histology Embedded samples selected for histology analysis were sectioned using a SP

1600 Microtome (Leica, Germany). After mounting the specimen, an initial planed section was performed to ensure the superior surface of the sample was perfectly flat. A glass slide was secured to the sample using a UV adhesive

(Loctite, Dusseldorf, Germany) and a subsequent sectioning was performed to create a 30µm thick section. The section was then cleaned and stained using methylene blue and basic fuschin to highlight bone and fibrous tissues. Two histology sections were taken from each selected sample.

Histology slides were viewed under a light microscope (Olympus, Tokyo, Japan) to observe new bone formation and cell characteristics at both the immediate implant-bone interface, as well as the surrounding extremities of the specimens.

34 Chapter 3. Results

3.1 SEM Coating Characterisation Scanning electron microscopy of the two implant surface topologies are presented in Figure 9 to Figure 12, for two magnification levels, prior to implantation.

Figure 9 SEM imaging of acid etched surface (Group 1) (x50 magnification). Texturing and evenly distributed ridges achieved through acid-etching process visible

Figure 10 SEM imaging of Ag HA (Group 2) coating (x50 magnification). Unique surface of plasma-spray coating visible in the form of compact undulations and microstructures. 35 Figure 11 SEM imaging of acid etched surface (Group 1) (x1000 magnification). Network of ridges, crevices and nanopits visible at this level of magnification.

Figure 12 SEM imaging of Ag HA coating (x250 magnification). Small voids and porous regions visible between the plasma-sprayed structures of the Ag HA coating. 36 At both magnification levels, there is a clear distinction in surface topologies of each group, as it to be expected when comparing a coated and uncoated surface.

Under 1000x magnification the nanotextured features of the acid-etched implant, including large ridges and a smaller network of nanopits and crevices can be clearly observed. Contrastingly, the plasma-sprayed surface of the Ag

HA coating upon the titanium substrate has resulted in a thicker surface coating with an undulating structure creating larger voids and prominences in a random distribution. The addition of elemental silver to the hydroxyapatite coating was not observed to result in an observable variation in surface structure or topology when compared to a conventional plasma-sprayed HA coating.

3.2 Surgery and Harvest There were no complications or issues arising from the surgical procedures or subsequent harvest at the designated time points, as described in the Method.

As previously noted, a haematoma was observed on the right hind limb of animal W2604 prior to harvest at the 12-week end point. However, this did not appear to impede mobility or cause any distress to the animal, therefore no further action was deemed necessary.

Following euthanisation of the animals, the right and left hind limbs were harvested and photographed for later reference. Upon removal of the surrounding skin and tissues from the bones, the implantation sites were clearly visible beneath the surface of tissue at both cortical and cancellous sites, as shown in Figure 13. There were no signs of infection or adverse reactions e.g.

37 discolouration or coating debris associated with the introduction of the implants into the animal’s limbs.

Figure 13 Harvested tibia from W2605 showing visible cancellous (blue arrow) and cortical (red arrow) implantation sites. Not evidence of foreign body reaction or implant rejection surrounding implants in situ

3.3 Radiography

Radiographs of the isolated femoral and tibial bones were taken in the anteroposterior and lateral views, encompassing all 5 implant sites on each limb. The X-Ray images were useful in confirming implant positions and alignment, as well as verifying the absence of post-operative fracture. There were no significant differences noted in the radiographs alone between implant groups at either time point.

Figure 14 to Figure 21 presents the radiographs of the harvested limbs in this study. All radiographs present the isolated femoral and tibia components from each animal. No irregularities or complications were noted upon review of the radiographs.

38 Figure 14 Left limb radiographs of animal W2602, AP(left) and lateral (right) views, showing good implant alignment at all sites

Figure 15 Right limb radiographs of animal W2602, AP(left) and lateral (right) views, showing good implant alignment at all sites

39 Figure 16 Left limb radiographs of animal W2603, AP(left) and lateral (right) views, showing good implant alignment at all sites

Figure 17 Right limb radiographs of animal W2603, AP(left) and lateral (right) views, showing good implant alignment at all sites

40 Figure 18 Left limb radiographs of animal W2604, AP(left) and lateral (right) views, showing good implant alignment at all sites

Figure 19 Right limb radiographs of animal W2604, AP(left) and lateral (right) views, showing good implant alignment at all sites

41 Figure 20 Left limb radiographs of animal W2605, AP(left) and lateral (right) views, showing good implant alignment at all sites

Figure 21 Right limb radiographs of animal W2605, AP(left) and lateral (right) views. Implant alignment was sufficient at all sites.

42 3.4 Mechanical Testing Sectioned cortical implant specimens were subject to pushout testing to evaluate the shear strength of the bone-implant interface. At the 4-week time point, both medial and lateral cortical specimens underwent mechanical testing. At the 12-week time point only medial cortical specimens were mechanically tested, in order to preserve the bone-implant interface of the lateral specimens for further evaluation through SEM and histomorphometry.

Shear stress (N/mm2, i.e. Pa) was defined as the peak axial pushout force acting over the total bone-implant contact surface area. Contact surface area was calculated as the cylindrical surface area of bone surrounding each implant, given by:

= x d x l

Where:

= ()

= . 6

= ℎ ()

Interface stiffness (N/mm) was defined as the gradient of the initial linear portion of the force-displacement curve. Energy to failure (J) was defined as the integral of the force-displacement curve from zero to the point of peak load.

These values were calculated from the raw force-displacement data using a custom script developed for Matlab R2009a (Mathworks Inc, MA, USA), as illustrated in Figure 22.

43 Figure 22 Force-displacement curve output from custom Matlab script. Green line indicates stiffness (N/mm) taken from initial linear portion. Shaded blue region indicates integral area calculated for energy to failure (J). Peak load is the force value corresponding to junction between blue and white shaded regions.

Results of mechanical testing are presented in Table 2 to Error! Reference source not found. and Figure 23. Statistical analysis was performed using

SPSS Statistics 23 (IBM, New York, USA). Each mechanical parameter was compared between groups at identical timepoints using an independent samples t-test. Differences were considered to be significant for a P value <

0.05. The Levenes Test for Equality of Variances was deemed to be violated when it’s significance was < 0.05, and thus homogeneity of variance was not assumed between groups. Full outputs from SPSS are presented in

Table 3 and Error! Reference source not found..

At both time-points, the mechanical performance of the Ag HA implants was significantly greater than those of the acid etched samples. The mean shear 44 stress of pushout at 4-weeks for the Ag HA implants was 8.35 MPa (SD = 3.89

MPa), compared with 3.11 MPa (SD = 3.24) for the acid etched implants. A similar variation was noted at 12 weeks, with the Ag HA implants recording mean shear stress of 17.84 MPa (SD = 5.11 MPa) compared with 5.22 MPa (SD

= 2.13 MPa) of the acid etched implants.

Table 2: 4-week Mechanical Testing Results showing calculated p-value for each parameter

Parameter Group 1 (Acid Etched) Group 2 (HA Ag Coating) P

Samples (n) 12 12

Peak Force (N) 248.61 (261.06) 675.34 (269.44) 0.001

Displacement to max (mm) 0.32 (0.19) 0.50 (0.20) 0.040

Stiffness (N/mm) 1551.34 (1134.49) 2795.04 (828.78) 0.006

Energy to Failure (J) 42.11 (44.17) 168.06 (158.27) 0.020

Average Cortical Thickness (mm) 4.18 (0.65) 4.65 (1.76) 0.393

Shear Strength (MPa) 3.11 (3.24) 8.35 (3.89) 0.002

Table 3 SPSS Output of Independent t-test for Mechanical Testing at 4 weeks. All parameters were found to be significantly different (p<0.05) between the groups (red boxes), excluding the average cortical thickness of the tested specimens (blue box). This indicates variations can be attributed to the quality bone-implant interface.

45 Table 4: 12-week Mechanical Testing Results showing calculated p-value for each parameter

Group 1 Group 2 P Parameter (Acid Etched) (HA Ag Coating)

Samples (n) 6 6

0.000 Peak Force (N) 642.39 (386.05) 2133.10 (571.11)

0.035 Displacement to max (mm) 0.33 (0.12) 0.65 (0.28)

0.029 Stiffness (N/mm) 3157.73 (1188.90) 5876.60 (2323.67)

0.01 Energy to Failure (J) 103.92 (86.07) 620.56 (271.96)

Average Cortical Thickness 0.653 6.20 (1.97) 7.14 (4.11) (mm) 0.002 Shear Strength (MPa) 5.22 (2.18) 17.84 (5.11)

Table 5 SPSS Output of Independent t-test for Mechanical Testing at 12 weeks. All parameters were found to be significantly different (p<0.05) between the groups (red boxes), excluding the average cortical thickness of the tested specimens (blue box). This indicates variations can be attributed to the quality bone-implant interface.

46 Figure 23 Mean shear stress of pushout at 4-week and 12-week timepoints for acid-etched and Ag HA implant groups. Error bars indicate standard deviation. Ag HA implants had significantly higher shear stresses at both timepoints.

All mechanical parameters measured were observed to be significantly greater for the Ag HA implants, when compared to acid etched, across both timepoints.

The exception to this was the average cortical thickness measurements, which were not found to be significantly different between the two groups for either timepoint (P=0.393 and P=0.653). This provides the implication that the intrinsic parameters calculated using the average cortical thickness measurement, specifically shear stress of pushout, represents the true differences between implant performance in each group. It further highlights that the method used to obtain cortical thicknesses was a robust a repeatable exercise.

47 3.5 SEM Imaging SEM images were processed using a custom MATLAB script, which was able to calculate the percentage of bone-implant contact based upon a predefined trace of calculated perimeter.

It was observed that the MATLAB script was occasionally unable to successfully identify the entire bone-implant contact perimeter within a single image. This generally occurred when the level of contrast between bone and implant was similar, thereby increasing the difficulty in differentiating between the two within practical program thresholds. In these cases, a single

SEM image was analysed in multiple sections, each containing one part of the total bone-implant perimeter. Different regions of the image were isolated by varying the program thresholds to alter the sensitivity to variation in pixel colour.

3.6 Histomorphometry A summary of ongrowth results for cortical and cancellous samples are presented in Table 6 and Table 8 respectively. An analysis of variance

(ANOVA) was conducted using SPSS Statistics 23 (IBM, New York, USA) to test the bone ongrowth performance of each test material at both timepoints.

There was significant variation in bone ongrowth between the groups, and between samples at each time point (P = 0.000), noted for both cortical and cancellous samples. The variation between the bone ongrowth ratios for each group appear to be independent of the time in situ, as indicated by a P-value of

0.362 for cancellous and 0.372 for cortical samples (P > 0.05). This is demonstrated in Table 7 and Table 9. Full histomorphetric results are presented in Appendix A. 48 Table 6 Bone ongrowth evaluation of cortical specimens. Mean percentage ongrowth of the Ag HA implants was consistently higher than that of acid-etched implants at both timepoints.

Animal ID Parameter Group 1 (Acid Etched) Group 2 (Ag HA) Samples (n) 6 6 W2602 Mean On Growth (%) 51.69 68.69 SD 8.51 8.17 Samples (n) 6 6 W2603 Mean On Growth (%) 51.85 63.18 SD 11.81 5.41 Samples (n) 6 6 W2604 Mean On Growth (%) 67.12 81.52 SD 9.36 10.36 Samples (n) 6 6 W2605 Mean On Growth (%) 76.42 81.47 SD 6.43 4.34 Samples (n) 12 12 4 Week Mean On Growth (%) 51.77 66.03 TOTAL SD 9.82 7.25 Samples (n) 12 12 12 Week Mean On Growth (%) 71.77 81.50 TOTAL SD 9.06 7.57

Table 7 ANOVA Output for cortical samples, testing bone ongrowth at all timepoints. BIC is dependent upon implant type and timepoint, but variation between groups is independent of time in situ

49 Table 8 Bone ongrowth evaluation of cancellous specimens. Mean percentage ongrowth of the Ag HA implants was consistently higher than that of acid-etched implants at both timepoints.

Animal ID Parameter Group 1 (No Coating) Group 2 (Ag HA) Samples (n) 2 2 W2602 Mean On Growth (%) 33.88 53.83 SD 8.00 8.53 Samples (n) 2 2 W2603 Mean On Growth (%) 27.78 55.42 SD 3.18 4.84 Samples (n) 2 1 W2604 Mean On Growth (%) 37.48 54.42 SD 21.13 0.00 Samples (n) 2 2 W2605 Mean On Growth (%) 43.26 58.15 SD 15.37 11.90 Samples (n) 4 4 4 Week TOTAL Mean On Growth (%) 30.83 54.62 SD 6.10 5.74 Samples (n) 4 4 12 Week TOTAL Mean On Growth (%) 40.37 56.90 SD 15.45 8.69 Table 9 ANOVA Output for cancellous samples, testing bone ongrowth. BIC is dependent upon implant type, but independent of timepoint, and variation between groups is independent of time in situ

50 3.7 Histology Findings Histologically, both implant groups demonstrated signs of effective osseointegration into the ovine bone, particularly at the 12-week timepoint. A qualitative evaluation of the histological slides supported the quantitative findings the bone-implant contact assessments, in that the Ag HA implants generally demonstrated a greater degree of bone proliferation and contact when compared to the acid-etched implants.

The woven bone surrounding the Ag HA implants was characterised by increased density and higher continuity when compared to that surrounding the acid etched implants. Figure 24 illustrates the presence of Haversian systems surrounded by a large number of osteocytes observed in the proximity of the Ag HA coating, indicating remodelling and the rapid establishment of mature bone.

Figure 24 10x magnification of W2605L5L, showing Haversian systems (arrowed) and dense woven bone (circled) in proximity of Ag HA coating 51 The acid etched implants also showed signs of osteoblastic activity at numerous sites both adjacent to the implant surface and at a greater distance from the immediate implant site. Figure 25 presents an example of rounded osteoblasts in an enlarged state layered upon newly formed bone at a number of sites surrounding the acid-etched implant surface.

Figure 25 10x magnification of W2605R5M, showing rounded and enlarge osteoblasts layered upon newly formed bone (arrowed) at various sites

Despite both implant groups demonstrating effective osseointegration of implants into the ovine bone at 12-weeks, there was a clear distinction in the extent of bone-implant contact between the two groups. The Ag HA implants consistently displayed a more uniform degree of contact along the implant length, with new bone formations adhering closely to the outer surface of the coating and on occasion penetration through the coating itself toward the 52 titanium substrate below. Figure 26 and Figure 27 highlight the variation in bone ongrowth between the implant groups. The propensity of new bone formation surrounding the Ag HA implant shown in Figure 26 extends to the lateral face of the implant as well as the internal threaded surfaces of the dowel. The newly formed bone directly adjacent to the coating is observed to be tightly packed and closely adhered to the coating, with little to no voids along the implant length.

Figure 26 Newly formed bone on lateral implant face and internal thread (arrowed) in Ag AH group

Contrastingly, new bone formation in the acid-etched implant group was observed to have reduced bone-implant contact, and a more frequent appearance of voids in the newly formed bone, within the proximity of the implant surface as illustrated in Figure 27 53 Figure 27 Increased frequency of voids (arrowed) and reduced bone-implant contact (boxed) in acid-etched implant group

There was no histological evidence of foreign body reaction or other immunological responses at either time-point across each of the implant groups. The absence of multinucleated or histiocytic cells further indicates that the implant materials of both groups were effectively incorporated into the animals with no visible adverse reactions. Whilst the extent of bone implant contact regions and the quality of said contact varied between the substrate treatments, neither group displayed signs of an intervening fibrous tissue layer which often indicates inadequate osseoconductivity or potentially excessive micromotion of the implant.

54 Chapter 4. Discussion

This study aimed to evaluate the osseointegration of two distinct surface modifications of a titanium implant substrate, using an established ovine model

[53-56]. The surface treatments included acid etching to produce a nanotextured titanium surface, as well as a novel hydroxyapatite coating doped with silver with to promote release of silver ions around the implant site in situ. The implants were placed within the cancellous bone (n=4 per animal) and cortical bone (n = 6 per animal) of the tibial midshaft. A total of 4 animals were used in the study, with osseointegration assessed through mechanical testing and histomorphometric analysis at 4 weeks (n=2) and 12 weeks (n=2) after implantation. All animals showed full recovery following the surgical procedure.

Cortical implants were placed in a line-to-line fashion across the bicortical implantation site. Cancellous implantation sites were prepared using a step- drill, to create a range of implantation scenarios for each sample. The step-drill allowed implants to be assessed in press-fit, line-to-line and gap scenarios within the same site.

Mechanical testing of both groups demonstrated improvement in shear strength from 4 weeks to 12 weeks following implantation. Mean shear stress of the acid-etched implants increased significantly from 3.11MPa (SD: 3.24Mpa) at 4 weeks to 5.22MPa (SD: 2.18MPa) at 12 weeks. Similarly, mean shear stress of the Ag HA implants also increased significantly from 8.35MPa (SD:

3.89MPa) at 4 weeks to 17.84MPa SD: 5.11MPa) after 12 weeks.

55 Improvements in bone-implant contact ratios at the 12-week timepoint when compared to the 4-week timepoint were also noted for both groups, in cortical and cancellous sites. At the 12 week timepoint, the mean BIC of the cortical acid-etched implants had increased significantly to 71.8% (SD: 9.1%), when compared to the 4 week levels of 51.8% (SD: 9.8%). Similarly, there was a significant increase recorded in the mean BIC of the cortical Ag HA samples, increasing from 66.0% (SD:7.3%) to 81.5% (SD:7.6%) at 12 weeks.

The increases in shear strength over time, accompanied by significant improvements in the bone-implant contact ratios in each implant group are all indicative of an increase in the strength of the bone-implant interface occurring whilst the implant remains in situ. Improvements in shear stress over time are representative of the propensity of both surface treatments to effectively promote osseointegration, leading to subsequent bone ongrowth onto the titanium substrate.

Whilst both surfaces successfully demonstrated osseointegration capacities, with the notable absence of foreign body reactions or biological rejection within the ovine model, the Ag HA coating was observed to consistently outperform the acid etched surface when assessed at each timepoint.

The mechanical strength of the newly formed bone was significantly higher

(P<0.05) in the Ag HA implants than the acid-etched counterparts. There was no significant difference in the mean measured cortical thicknesses between the two groups at either 4-weeks (P=0.393 or 12 -weeks (P=0.653), indicating the measured shear stress values are an accurate comparison of the bone-implant interface strength. Other measured parameters including energy to failure and 56 stiffness were also higher in the Ag HA group, when compared to the acid etched group at each time point.

Similarly, the BIC ratios of the Ag HA implants were significantly higher(P<0.05) than the acid etched implants at each timepoint across both cortical and cancellous implantation sites. The findings from the histological analysis confirmed the quantitative results and demonstrated a denser, more continuous formation of newly woven bone surrounding the Ag HA coating.

The variation in implant performance and the differing abilities of each surface treatment to promote osseointegration and bone formation can likely be attributed to both the physical and chemical compositions of each surface treatment analysed in this study. As highlighted previously, there was a distinct variation in the surface topologies of each implant, produced by the inherent processing methods used to create each of surface treatments.

Further, the chemical compositions of each of implant surfaces were vastly different, due to the addition of the Ag HA coating onto the common Ti6Al4V substrate. Implant topology and roughness, at both the micro- and nano-levels, as well as the presence of bioactive coatings, have been demonstrated in previous ovine studies to influence the mechanical strength of the bone-implant interface when assessed under shear pushout conditions, as discussed below.

4.2 Comparison of results with other ovine model studies The bilateral bicortical ovine model used in this study has been well established with a strong history of use in evaluating shear stress and osseointegration of a range of implant materials, surface treatments and coatings [53-56]. Therefore, this model provides a useful method to quantifiably compare implant and

57 coating performance across a range of studies and subsequent samples properties.

Previous iterations of this ovine model have evaluated a range of implant materials and surface treatments, including but not limited to grit-blasted HA

[57] electron beam melting (EBM) manufactured titanium implants [58] plasma sprayed titanium with and without dicalcium phosphate dihydrate (DCPD) coatings [55]. The results obtained from the current study have been compared to those recorded for other implant types in the same ovine model, specifically for shear stress of cortical pushout and bone implant contact as determined by

SEM and subsequent quantification.

Figure 28 and

Figure 29 present comparisons of results from the current study to a range of other implants utilised in the ovine model. The figures compare results of the 58 acid-etched titanium (A Ti) and silver HA coating (Ag HA) from this study with three variants of plasma-sprayed titanium coatings (PS Ti* [58], PS Ti** [53],

PS Ti*** [55]) with a 150µm thick HA coating (PS Ti 150HA [53]), electron beam melted titanium (EBM Ti [54]), grit-blasted titanium (GB Ti [53]) with a

100µm thick HA coating (Gb Ti 100HA [53]) and a 200µm thick HA coating (GB

Ti 200HA [57]), plasma-sprayed titanium with a dicalcium phosphate dihydrate

(DCPD) coating (Ps Ti DPCD [55]) and finally a porous titanium coating on titanium substrate (Po Ti [56] ).

Data was evaluated from both mechanical testing and shear stress of pushout as well as the histomorphometric analysis of bone implant contact for cortical implants at both 4 weeks and 12 weeks post-operatively.

Figure 28 Mean Shear Stress of cortical implants in ovine model at 4 week and 12 week timepoints. Results from current study highlighted in purple (4 weeks) and green (12 weeks). Acid-etch titanium had lower shear stresses than all coated titanium substrates.

59 Figure 29 Mean Bone Implant Contact of cortical implants in ovine model at 4 week and 12 week timepoints Results from current study highlighted in purple (4 weeks) and green (12 weeks). BIC for PS Ti** was markedly lower than other plasma-sprayed variants.

When considering the mean shear stress of pushout, there is a distinct trend of greater shear stresses associated with rougher and/or more porous surface treatments. Across the various studies, the highest shear stresses were generally recorded for plasma-sprayed titanium surfaces, both with or without additional surface treatments, as well as the EBM titanium and porous coated titanium processes. The exception to this was the plasma-sprayed titanium surface utilised in the Svehla study [53] PS Ti**, which will be discussed in greater detail. Surface treatments which resulted in less pronounced variations in the implant surface were seen to correspond with lower shear stress of pushout, most notably the acid etched titanium from the current study as well as a grit-blasted titanium surface. The addition of a HA coating was observed to improve the shear stress when compared to equivalent substrates without the

HA coating. While addition of HA to a plasma-sprayed titanium surface did improve shear strength when compared to samples without the HA, it did not

60 increase shear stress above that achieved by EMB titanium or plasma-sprayed titanium from other iterations of the ovine study.

When evaluating bone-implant contact between the various group, a second trend was identified. Implants with the greatest bone implant contact percentages all utilised a HA coating on the outer surface of the specimen. The grit-blasted HA samples, including the current Ag HA, as well as the plasma- sprayed titanium with HA coating represented the implant groups with the greatest bone implant contact ratios. The HA coating was also seen to outperform the DPCD coating used in conjunction with a plasma-sprayed titanium surface, when considering bone implant contact alone. Despite demonstrating greater mean shear stresses during mechanical pushout, the plasma-sprayed titanium, porous titanium and EBM titanium surfaces had lower bone implant contact ratios across most of the studies when compared to the bioactive coated groups. The differences in shear stress and bone implant contact ratios could potentially be attributed to the underlying influences of surface roughness and topology for each of the various substrate treatments, as discussed below.

4.3 Effect of surface topology upon bone ongrowth The variations in bone ongrowth between implants with contrasting surface topologies and coatings highlights an important distinction between the properties of the various surface treatments. The bioactivity of HA coatings has been widely demonstrated in the literature to effectively promote bone ongrowth [59]. Therefore, the results of this comparison of bone implant contact across ovine models are in line with the expected outcomes based upon the

61 properties of HA when compared to a titanium substrate alone. However, the improved shear stresses resulting from various forms of surface treatments illustrates that the mechanical fixation and stability of the implant within the bone is not solely reliant upon bioactivity of the surface, but also the microstructures and surface interactions that occur between the implant and surrounding cells.

The influence of implant surface roughness and topologies upon cell behaviour and subsequent implant osseointegration have been the subject of several in vitro studies. Further research by Svehla et al [53] examined the influence of substrate roughness and the thickness of HA coatings upon subsequent implant shear stress up to 26 weeks following implantation. The same ovine model as described in this study was used, with different timepoints employed (4, 8, 12 and 26 weeks in situ). This study utilised titanium dowels which were divided into the following groups; grit blasted (GB) and plasma-sprayed with a HA coating of either 50µm (GB50) 100µm (GB100) or 150µm (GB150), plasma- sprayed titanium with a plasma-sprayed HA coating of 50µm (PL50), 100µm

(PL100) or 150µm (PL150), or polished uncoated titanium alloy (S).

Crystallinity of the HA coating increased with coating thickness, with crystallinity values of 30% for 50µm, 50% for 100µm and 60% for 150µm thick coatings. Surface analysis revealed that the plasma-sprayed substrate resulting in greater average implant roughness than the grit-blasted substrates, whilst the polished titanium surfaces presented the smoothest surfaces, as expected.

At all measured timepoints, the plasma-sprayed substrates were found to display superior performance to the grit-blasted substrates, which in turn were

62 superior to the polished titanium implants. Shear stress values and bone implant contact percentages were greater in the plasma-sprayed substrate groups, when compared to the grit-blasted substrates. The thickness of the HA coating, and hence the quantity of hydroxyapatitie applied to the implant substrate, was not found to have a significant influence on the bone implant contact of the specimens, although an upward trend was noted in bone implant contact between 4 and 26 weeks when comparing 100µm and 150µm thick coatings on both substrate types. Failure during mechanical pushout at 4 weeks occurred between the newly formed peri-implant bone at the drill hole interface, whilst at all later time points failure occurred at the HA coating-bone interface. The findings from this study regarding HA coating thickness were noted to correlate with other published results [60] in which underlying substrate metal and topographies were seen to be generally be the driving factor determining shear strength and osseointegration. However, in the absence of a HA coating, as shown by the polished titanium implants, bone ongrowth and shear stress was significantly lower than when a HA coating was applied. In the smooth implants, BSEM analysis revealed minimal bone implant contact, and the presence of a soft tissue layer at the bone-implant interface which inhibits adequate bone contact with the implant.

The results of the current ovine study are found to be congruent with these previous findings. The grit-blasted titanium substrate with Ag HA coating had superior shear stress and bone implant contact when compared to the uncoated acid-etched titanium samples at both 4 week and 12 week timepoints. Whilst surface roughness measurements were not taken in this study, the presence of

63 HA was shown to enhance implant fixation, as has been widely reported in other animal and clinical studies [61-64].

The texturing of the titanium surface achieved by the acid-etching process for implants in this study creates local topography allowing for a mechanical interlock with newly formed bone, as observed in SEM imaging of the samples prior to implantation. The surface roughness of titanium has further been previously demonstrated to influence the process of osteoblast differentiation, including events such as spreading and proliferation. An in vitro study by

Sammons et al [65] evaluated the interaction between rat calvarial osteoblasts and a range of titanium dental implants with varying microstructured surfaces.

The titanium implants included grit-blasted and/or acid-etched, plasma- sprayed, anodised and smooth (as machined) surfaces. Surface roughness (Ra) ranged from 0.81µm (as machined) to 3.5µm (plasma-sprayed). Implants were immersed in a culture medium with bone fragments from two animals placed covering the exposed implant surfaces. Samples were then incubated for either

2 weeks or 4 weeks to provide two timepoints for final analysis.

Results from this study indicated that osteoblasts attached and spread more readily on the grit-blasted/acid-etched surfaces more rapidly than others examined, with the possible exception of plasma-sprayed surfaces, which promotes these processes due to the porous microstructure created. Notably in this study, two implant variants undergoing the same surface treatment (grit- blasted/acid-etched) and with comparable surface roughness (2.41µm vs

2.75µm) had significant variation in the percentage of osteoblast cells spreading across the implant surface (p < 0.01). This variation was attributed to the

64 differences in microtopography of each surface, which could not be discerned based upon surface roughness alone.

A similar study conducted by Salou et al [66] also demonstrated the variations in osteoblast cell behaviour when encountering microstructured titanium surfaces prepared using identical processes, including the commonly used grit- blasting and acid etching sequences. The significant differences arising from seemingly equivalent surface preparations is a result of the random and uncontrolled microstructures formed when these processes are applied to the titanium substrate. The unpredictable microstructures created can have an indeterminate impact upon impact osteoblast call and protein behaviour which interact with implant surfaces at a nanometer level. The study by Salou highlights the variations in microstructured verses nanostructured surfaces by comparing the osseointegration of grit-blasted/acid-etched titanium implants

(MICRO) and nanostructured (NANO) implants in a rabbit femur model. A standard as-machined implant (MA) was also included as a control group. The

MICRO group was found to exhibit the typical random cavities and structures expected from the grit-blasting and acid-etched surface preparations. This included large cavities produced by the grit-blasting, which also displayed internal microstructures produced by acid-etching when viewed under higher magnifications. The NANO surface was achieved through a regular array of titanium oxide nanotubes approximately 37nm in diameter with a 160nm thickness. Interestingly the NANO and MA groups had similar average surface roughness of 0.6µm and 0.5µm respectively, which was relatively smooth compared to average 1.5µm surface roughness of the MICRO group.

65 Following 4 weeks after implantation, the three groups were assessed for pull-

out strength, bone-to-implant contact and bone growth. The NANO implants

were found to have better integration into the bone when compared to the

standard MICRO surface, with regard to both bone-implant-contact and bone

growth, although the improvement was not statistically significant. Both the

MICRO and NANO groups outperformed the control MA group. The

equivalence in bone anchorage performance between the MICRO and NANO

groups, despite the NANO surface being almost 3 times smoother,

demonstrates that surface roughness alone is not a determinant for potential

osseointegration of a titanium implant, a conclusion supported by other studies

comparing micro- and nano-structured titanium surfaces [67, 68]. The

nanoscale characteristics and environment of implant surfaces are a critical

factor in determining the cellular behaviour immediately following

implantation into the body [69, 70]. Whilst many studies focus upon the attachment and proliferation of osteoblast-like cells, the first cells to colonise an

implant surface are mesenchymal stem cells (MSC’s) which have the potential

to differentiate into osteoblasts, chondroblasts, mysoblasts and adipocytes [69].

Differentiation is determined by a number of specific cues including the

presence of cytokines, growth factor or the micro-environment. These cells

interact with implant surfaces and structures at a nanolevel, and as such,

nanostructures and characteristics including pore sizes and structure arrays

have the capacity to determine whether MSC’s differentiate into mature

osteoblasts leading to direct bone apposition, rather than into fibroblastic cells

which can lead to fibrous tissue layers at the bone-implant surface. Therefore,

whilst surface roughness characterises surfaces at the microlevel, the complex 66 interactions occurring at the nanolevel contributing to eventual

osseointegration of implants should also be considered when evaluating the

performance of a textured titanium implant substrate surface.

As well as effectively promoting cell behaviour at the implant interface, the inherent ability of implant surface topology to facilitate strong mechanical fixation and bone attachment can be critical in minimising micromotion of implants, to allow osseointegration to occur. Studies have demonstrated that a reduction in micromotion below 40-70µm results in favourable conditions conducive to the proliferation of bony-ingrowth, whilst motion exceeding 150µm leads to fibrous tissue on-growth [71-73]. Practical applications of this principle were examined by Harrison et al [74], who develoed a novel anchor-like surface topology created through direct metal laser sintering, which was able to provide up to 76% greater resistance to transverse motion than a porous tantalum surface structure. The anchor-like properties lead to a significant reduction in micromotion amplitudes when implanted in a bovine model, and subsequently demonstrated significantly greater bony ingrowth and primary fixation when compared to plasma sprayed implants, in an in vivo bovine study [75].

When considering only the results of the acid etched titanium implants from the current ovine study, the quantitively outcomes are in line with the results expected given the outcomes from previous studies conducted using the same ovine model or in vitro models. When considering the Svehla et al [53] study as described above, the mean cortical bone implant contact for smooth titanium implants was reported as approximately 6% at 4 weeks, increasing to approximately 18% at the 12 week timepoint. The acid-etched titanium

67 implants in this study had mean cortical bone implant contact of 52% at 4 weeks, increasing to 72% at 12 weeks. Further, there was no indication of a fibrous tissue layer at the bone-implant interface, as recorded for the smooth titanium implants, which would have reduced the potential for cellular attachment to the implant surface. However as discussed, acid-etching processes result in a random and unpredictable formation of microstructures which may not provide optimal conditions for cell adhesion or differentiation at the nanolevel.

4.4 Current developments of silver in implants

4.4.1 Functionalised HA Coatings With the antimicrobial properties of silver well established since ancient times, research into the incorporation of silver ions into a range of medical applications has been constantly ongoing in a range of fields. More recently, research has been conducted in a laboratory setting to examine the viability of incorporating silver into bioactive coatings such as HA with attempts to examine the limitations and effective compositions to achieve the balance between osteointegration and effective antimicrobial activity.

A study by Roy et al [76] examined the functionalisation of HA coatings on a titanium substrate through the addition of silver particles. This study describes the methods to introduce silver to the hydroxyapatite and assesses the biocompatibility and antimicrobial effectiveness the resultant coating. Elution profiles for ions leaching from the implant were also evaluated across each group for varying concentrations of silver ions.

68 A key feature of the study was the way in which silver particles were introduced to the coating. Silver particles were incorporated within the HA coating itself rather than having them present at the HA surface. This was achieved doping the HA powder itself, by mixing with silver oxide (Ag2O) at the desired concentrations (2%, 45 and 6% wt). Therefore, when the coating was applied via plasma spraying to the titanium substrate, it contained a uniform distribution of silver particles throughout the entire composition instead of only at the surface.

This process is equivalent to that utilised in the current ovine study in which the silver particles incorporated in the implants had been added to the HA powder prior to application on the substrate, rather than being cultivated onto a previously applied HA coating.

This differs from other variants of silver-doped HA coatings, such as those described in research by Lee and Murphy [77]. This study propagated silver particles onto the external surface of the HA by transferring citric-acid treated specimens into a silver nitrate solution. The resultant effect was a uniform covering of silver particles on the external surface of the HA. The size of the silver particles, ranging from tens of nanometers to several micrometers was dependent upon the concentration of silver nitrate and subsequent incubation time.

The experimental procedure of the Roy et al study relied upon an evaluation of the interactions between the coating and bone cells (to assess biocompatibility) and a challenge of P. aeruginosa (to assess antimicrobial activity). 4 concentrations of silver ions at 2%, 4% and 6% wt were examined, alongside HA with no silver introduced. Initially the tensile strength of the plasma-sprayed

69 coatings was evaluated, as well as the silver release profile for each level of silver concentration. Subsequently, MTT assays, ALP activity and cellular morphology was used to examine osteoblast proliferation and cellular attachment to the coating surface. The efficacy of the silver doped coatings against bacterial colonisation and adhesion was characterised by Live/Dead fluorescent staining of cultures following 24 hours exposure to the various coatings.

Results from this study indicated that the introduction of silver ions into plasma- sprayed HA resulted in a bioactive coating which was further able to provide antimicrobial activity. The tensile strength and adhesive properties of the coating were not adversely affected by the addition of silver, with comparable bond strengths across all groups, and no samples showing pure adhesive failure

(failure at coating-substrate interface) after testing. This is likely due to the Ag particles undergoing the same coating formation process as the similar sized HA particles during plasma spraying, such that the final adhesive strengths are similar for both components after mechanical bonding to the substrate.

There was no significant variation in cell proliferation between the conventional

HA and silver doped HA up to a concentration of 4% wt. after 11 days, when examined through MTT assay. The 6% wt. group however did not show any signs of cell proliferation at any stage throughout the trial, with cell morphology assessments indicating cell death. Reducing levels of cell density were also noted for increasing silver concentrations, however the variation was not statistically significant. Conversely, increased concentrations of silver in the coating improved the antimicrobial efficacy, as may be expected. Even at the 2% wt. concentration of silver ions, the majority of single or colonized bacteria which had

70 adhered to the surface of the coating were found dead. At the 4% wt. concentration, there a visible reduction in bacteria overall, with many also found to be dead.

The results from this study clearly indicate that the osseointegration and antimicrobial properties of a potential HA silver coating are largely dictated by the concentration levels of the silver ions within the applied coating. The inherent antimicrobial properties of silver can become harmful to body cells at excessive concentrations, which in turn inhibits osteoblast proliferation and cell attachment vital for osseointegration. In this research, silver concentrations of

2-4% wt. were found to display antimicrobial effects without adversely impacting osteoblast proliferation. However, at a concentration of 6% wt. silver, cell proliferation and attachment was inhibited despite the improved bactericidal effects.

Interestingly, the elution profiles for the release of silver ions from the coating were initially found to be independent of the silver concentration. Within the first

6 hours, the 2%, 4% and 6% wt. silver coating all released approximately 200ppb

Ag/Ag2+. However, beyond this initial period, an increase in silver concentration was found to correlate with an increase in the release of silver ions, with the 6% wt. samples displaying the highest concentrations at all time points. The amount of silver released was generally observed to be cumulatively less than 1% of the total silver content within the coating. However even at the lowest concentration of 2% wt., this corresponded to approximately 760ppb after 7 days, which still exceeded the minimum inhibitory concentration of 6.25ppm for antimicrobial effect.

71 The high levels of silver retention within the coating itself is significant in allowing the surface of the implant to remain resistant to bacterial adhesion and hence prevent the formation of biofilms over a longer period of time. This unique property is a result of incorporating the silver particles into the HA coating itself prior to application by plasma spray. Contrastingly, the addition of silver to the outer surface of the HA coating will provide the initial antimicrobial activity as the silver ions are released into interstitial space, but will not provide ongoing resistance against bacterial adhesion following the initial elution period. This is a key consideration when evaluating the suitability of silver doped coatings for use with orthopaedic implants, when the initial 6-hour window post-operatively is viewed as a critical period in preventing the onset of infection by harmful organisms, such as S. aureus, which generally present in early stages of implant life. early infection (occurring within 3 months) [72, 74]. Delayed infections in an open wound scenario require continued resistance and antimicrobial activity, which silver ions within a coating could provide.

4.4.2 Silver-doped Strontium coatings Contrasting results were recorded in recent work by Geng et al [78] which further evaluated the viability of incorporating the antimicrobial properties of silver into a bioactive coating for titanium substrates. This research assessed both the biological and antibacterial properties of a HA incorporating silver ions, strontium ions and a third group with both strontium and silver ions included. Strontium ions were included in this study on the basis that the silver ions alone may impair the biocompatibility of the HA coating and reduce the bone cell attachment and proliferation. As strontium is a bone-seeking element, it was hypothesized that the Sr2+ ions could be used as a secondary doping 72 chemical to offset the harmful effects of Ag2+ to ultimately create an optimally balanced anti-microbial and bioactive coating. Four coating groups were examined in this study, classified as HA, Ag 0.1 (HA with Ag2+), 10Sr (HA with

Sr2+) and Sr/Ag (HA with both Ag2+ and Sr2+).

Biocompatibility of the various coatings were assessed by estimating the cell proliferation of osteoblast-like MG63 cell seeded onto the coatings. Samples were incubated for periods of 1, 3 and 7 days, and up to 14 days for alkaline phosphatase (ALP) activity tests. Final quantification of biocompatibility was based on colorimetric MTT assays of the coating solutions, as well as ALP activity of MG63 cells on the coatings after up to 14 days of incubation.

Antibacterial activity was evaluated using agar disk diffusion methods against

E coli and S. aureaus. Prepared Ti-based coatings were placed onto agar plates seeded with the respective bacterium and allowed to incubate for 24 hours.

Following this time, the diameter of the incubation zones surrounding each coating sample was measured to define antimicrobial effectiveness. SEM and

TEM methods were also used to more deeply examine bacterial morphology following the interaction with the various coating groups.

Results from this study found that the HA coatings incorporating Ag2+ alone provided the greatest level of antibacterial activity when compared to the other groups. However, it also demonstrated the lowest biological compatibility with the lowest levels of cell proliferation and ALP activity of MG63 cells.

Conversely, the Sr10 group showed excellent cell proliferation with no antimicrobial activity. As hypothesised, the Sr/Ag group with both ion types incorporated into the HA provided a balance of effective antimicrobial activity 73 against both E. Coli and S. aureus, without compromising MG63 cell proliferation on the coating surface. The well-known benefits of each ion were able to be utilised independently within the same bioactive coating.

While the results of this research further support the use of bioactive coatings doped with Ag2+ as an antimicrobial solution, the reported adverse effect on bone cell attachment was a significant concern to the researchers. The addition of strontium to encourage osseointegration, despite the cytotoxicity of the silver ions, provided a viable alternative. The concentration of silver used in the Ag

0.1 coating developed by Geng et al [78] can be calculate based upon the specified elemental ratios of the coating itself. The measured proportions of

Ca:Sr:Ag:P in the Ag0.1 coating were 9.73:0:0.1:9.6. This corresponds to a

0.57% silver in an elemental ratio, or 1.66% wt. concentration within the coating.

As noted in the Roy et al [76] study, the concentration of silver ions alone can be adjusted to provide an appropriate balance between antimicrobial activity and biocompatibility. However, a silver concentration up to 4% wt. was observed not to adversely impact osteoblast proliferation or attachment to the implant surface. Therefore, the findings of Geng et al using a concentration below 2% appear to be contradictory. The discrepancy may be due to difference methods by which coatings were synthesized and applied to the titanium substrates. All coatings evaluated by Roy et al were produced using a plasma spray technique with silver particles incorporated into the HA powder at varying concentrations. Those tested by Geng et al study were formed using hydrothermal deposition using a range of reagents containing the desired

74 elements. While the influence of coating applications and processes upon potential osseointegration and ion release lies outside the scope of the current ovine study, it has been established that factors such as availability of surface metals, area of exposed coating surfaces and hydrophobicity of coating interfaces, all of which are affected by coating manufacturing processes, can impact final rate of ion release [76]. Therefore, the coating preparation and deposition technique used to create the silver doped HA coating in the context of the current ovine study, similar to traditional plasma spraying, should be kept in consideration when drawing comparison and evaluation against other research in the field.

While the work by both Roy et al and Geng et al provide insights into the effect of silver ions upon both beneficial and harmful cells, they are limited in their experimental design and scope to provide data regarding the mid-term performance of the coatings in an implant scenario. A strength of the current ovine study is that it allows a more comprehensive evaluation of osseointegration in a surgical implantation. Primarily, the experimental duration and study timepoints of all the methods has an overarching effect on the practical implications of the outcomes measured.

The cell based research described above defined timepoints of 1, 3, 7 and up to

14 days at a maximum. For studies based upon cell cultures and assays this is a suitable and effective time-period over which to evaluate results. It is particularly useful when considering outcomes directly influenced by elution of ions from the bioactive coating, in this instance antimicrobial effects provided by Ag2+ ions or increased bioactivity provided by Sr2+ ions. The reason for this is

75 that elution rates for ions incorporated into coatings have been shown to be greatest immediately following implantation, and gradually decline as the implant remains in situ over time, as discussed previously [77].

However, when considering the suitability of a coating for use in an implant setting, the current ovine study provides a more practical representation of an implantation scenario. The incorporation of an implant through osteointegration and bone adhesion takes several weeks, therefore the 4 week and 12 week timepoints of the current study allowed adequate time for these processes to take place. It further allowed a more comprehensive evaluation of biocompatibility and suitability of the coating which extended beyond cell proliferation. Potential adverse effects including foreign body reactions were assessed though the SEM analysis and histological imaging of each specimen in both cortical and cancellous bone. Similarly, the propensity for coating degradation because of silver doping could also be identified through these methods. Therefore, implantation of a representative implant containing the coating in this ovine model allows for a more comprehensive review of coating biocompatibility when compared to cell culture studies alone.

4.4.3. Agluna® material Ovine studies have also been performed to assess the biocompatibility and osseointegration of a silver-doped titanium surface with the tradename

Agluna® (Accentus Medical, Didcot, UK). The Agluna® material consists of a titanium substrate which has undergone an electrochemical process similar to high voltage anodization, within an aqueous solution containing silver. The resultant titanium surface contains circular pockets or reservoirs in which the

76 bulk of the ionic silver is stored. Concentrations of the silver ions can be controlled through the voltages used and duration of exposure of the material.

An ovine model has been used in a study by Coathup et al [79] to assess the osseointegration of Agluna® implants through pull-out strength and bone implant contact at 6 week and 12 weeks’ follow-up. This study further assessed antimicrobial activity through evaluating zones of inhibition and live/dead staining of bacterium grown in agar following exposure to the samples. Various implant groups were assessed in the study for both the osseointegration evaluation, as well as the assessment of antibacterial effects. For the in vitro studies examining antimicrobial activity, titanium alloy discs were divided into four groups: Polished titanium alloy (Ti), anodised titanium alloy (Ano),

Agluna® treated (Ag) and Agluna® treated with an additional conditioning step

(incubation in a culture fluid for 48 hours, Ag C). For the ovine study in which samples were implanted into the animals, grit blasted titanium alloy (Gb) and

Agluna® treated titanium at a silver concentration of 4 – 6 µg/cm2 (Ag) were evaluated at the 6 week timepoint. At the 12 week timepoint, an additional two groups were added; high dose Agluna® implants at 15 – 20 µg/cm2 silver concentration (Hd Ag) and anodised titanium alloy (Ano). 5 implants were used in each group, with 8 implants per animal (60 in total).

It was found that the non-conditioned Agluna® surface was in fact cytotoxic, however this was negated by the condition process undertaken by the Ag C samples. The number of live fibroblast cells on the surface of the Ti, Ano and Ag

C samples were significantly higher than that of the Ag samples. Further, the

Ag C samples also remained bactericidal despite allowing for fibroblast

77 attachment. When assessing the ovine model, mean shear strength during pushout at 6 weeks was significantly lower in the Ag group when compared to the Gb group (310.4N vs 561.2N, p=0.01). However at the 12 week timepoint there were no significant differences across all 4 groups for pushout shear strength. Further, bone implant contact and histological analyses did not identify any significant differences across the 4 groups at either the 6 week or

12 week timepoints.

Based upon the results of this study it was found that the conditioned Agluna® surface provided an appropriate balance between antimicrobial activity without impeding the osseointegration properties of the titanium substrate. The high dosed Agluna® samples also demonstrated equivalent bone implant contact and shear strength to the lower dosed Agluna® and untreated titanium indicating that silver concentrations up to 20 µg/cm2 did not have adverse toxic effects upon osseointegration or interfacial shear strength. This is a key consideration as the importance of identifying acceptable levels of silver within the implant material has previously been discussed, with respect to providing a suitable balance of desired qualities.

In addition to laboratory studies, Agluna® has demonstrated strong clinical effectiveness in providing an implant material which can promote osseointegration as well as assist in the prevention of post-surgical infection. A

2015 case control study by Wafa et al [80] evaluated 85 patients implanted with

Agluna® treated tumour implants between 2006 and 2011 matched against 85 patients implanted with identical untreated tumour prostheses in the same

78 period. The procedures included 50 primary reconstructions (29.4%), 7- one stage revisions (46.5%) and 41 two stage revisions (24.1%).

It was noted that overall post-operative infection rate when using Agluna® treated prosthesis was 11.8%, compared to 22.4% for the untreated implants

(p=0.033). Importantly, the Agluna® implants were most effective for patients undergoing a two-stage revision procedure in which periprosthetic infection was already present at the time of the surgery. In this cohort, the overall success rates for controlling infection using Agluna® implants was 85%, compared to 57.1% in the untreated control group (p=0.05). Cumulatively, the rate of recurrence of infection in all patients across both the one-stage and two- stage revisions was 8.47% for the Agluna® treated cohort, compared with 22.9% when using the untreated prostheses.

Another interesting observation was the improvement observed in the Agluna® group in response to conventional treatments for infection including debridement, antibiotics and implant retention (DAIR). 70% of infected

Agluna® implants (7/10) were successfully treated with DAIR, compared to

31.6% of the untreated control implants (6/19). This result indicates that the presence of silver ions surrounding the implant or within the joint space can enhance the effectiveness of traditional treatments against infection such as

DAIR.

Similarly, successful clinical evidence has been presented demonstrating the antimicrobial effectiveness of silver coated megaprostheses which are often prone to peri-prosthetic infection.

79 4.4.4 PorAg® coating An evaluation by Scoccianti et al [81] present the clinical outcomes of 48 patients implanted with a custom-made megaprosthesis (Waldemark Link) with a unique PorAg® coating. The coating is comprised of two layers: a lower basic silver layer with thickness 1µm, and a hard top layer of TiAg20N with thickness 0.1µm. The combination of the two layers creates a controlled electrochemical reaction which promotes silver ions and electrons to the surface of the coating to form an oligodynamic protective cover. The pure silver coating below leaches metal ions and non-silver particles providing bactericidal effects into the interstitial joint space surrounding the implant.

Patients were implanted between June 2010 and August 2015 with a mean follow-up of 25.9 months (range 12 to 56 months). All patients were deemed suitable candidates for the PorAg® implants as they represented cases in which infection had caused previous failure of implant, or they were currently at high risk of potential septic complications. This study aimed to evaluate the ability of the implants to prevent onset or recurrence of infection, as well as quantifying silver ion levels in the patients’ blood and urine following implantation of the device.

After two years post-operatively, no infections were recorded for any of the 12 patients who had previously not suffered septic complications. After 25 months, one patient developed an infection which was however preceded by a further two surgical procedures. In the 21 patients who had previously suffered septic complications, 2 patients (9.5%) had recurrent infection after 7 and 24 months following implantation. These results were in line with those previously

80 recorded by Wafa et al as discussed above, with a recurrent infection rate of approximately 8.5% when using silver-doped titanium implants.

4.5 Silver concentrations of current implants and coatings The mean silver concentrations in the blood when using the PorAg® implant were found to be lower than those recorded when using the similar styled

Mutars silver-treated megaprostheses. which was the subject of clinical studies by both Hardes et al [82] and Glehr et al [83]. The Mutars implant incorporates a silver coating achieved by galvanic deposition of elementary silver onto the titanium implant substrate. While both studies further supported the antimicrobial effects of a silver-coated prosthesis, they also present higher levels of silver concentration in both the blood and urine following implantation. The mean silver concentration in the blood of 20 patients evaluated by Glehr et al was 15.9µg/L, almost three times the maximum mean recorded by Scoccianti et al. Mean blood silver concertation were lower in the

Hardes et al study (1.93-12.98µg/L). However, the highest single blood silver concentration was recorded to be 56.4µg/L in one patient after 15 months, which was more than double the single peak concentration recorded by

Scoccianti et al. The studies using the Mutars implant were seen to generally have a greater variability in silver concentrations, with higher peak values, when compared to the PorAg® coating. This may be attributed to the relatively limited and stable leaching of silver ions created by the electrochemical reaction necessary to facilitate release into the joint space. By comparison, the galvanic deposition of elemental silver employed by the Mutars implant results in a more uncontrolled and random elution of ions from the coating.

81 Most notably, Glehr et al records a 23% rate of local argyria developing when using the Mutars prosthesis, although no instances are noted by Hardes et al when using the same prosthesis. Further, no instances of argyria have been reported for the PorAg® or Agluna® treated implants in a clinical setting.

While a direct correlation between blood silver concentration and the onset of argyria was not established in these studies, it is well accepted that excessive leaching of silver ions into the joint space is often responsible for the local discolouration of the skin observed in cases of argyria. The coating of the

Mutars implant is produced by the application of 0.33g to 2.89g of elemental silver applied to the titanium implant. This can be compared to the Agluna® treatment which uses a maximum of 6mgs of silver across the entire megaprostheis. Therefore, while the level of silver utilised in the Agluna® treatment has been shown to have effective antimicrobial properties, it is unlikely to result in local or systemic toxicity due to excessive silver concentration [84].

The properties of the Agluna® material are of particular significance to the current ovine study as it was used as the basis for the development of the tested silver HA coating by manufacturer Accentus Medical. Specifically, the target silver concentration of the Ag HA was designed to correlate with the previously established concentrations of the Agluna® treatment. The Agluna® treatment results in a target silver concentration of 4 – 6 µg/cm2 across the implant surface [79] with concentration up to 15 – 20 µg/cm2 also shown to have no adverse effects on osseointegration of implants in an ovine model. While the underlying production principles and manufacturing processes for the Agluna®

82 treatment are vastly different to those used to create the Ag HA coating in this study, the basis for silver dosage remains the same. The Ag HA coating applied to the titanium implants had a target concentration ranging from 2 – 10 µg/cm2 achieved by incorporating silver particles with the HA powder prior to coating application. The results of this study in which osseointegration and bone implant contact was not adversely affected by the addition of silver ions further supports the use of silver ions in implant coating at this concentration level.

4.6 Ag HA compared to conventional HA The recent developments and research into the functionalisation of osseoconductive coatings through doping with antimicrobial elements aligns closely with the aims of the current research study. In addition to facilitating a direct comparison of osseointegration properties between the two implant groups, a further point of interest from the current ovine study was the evaluation of the impact of elemental silver, when incorporated into a hydroxyapatite coating, upon osseointegration and cellular response to the implant.

The aforementioned in vitro studies examining cellular response and antimicrobial activity of silver doped coatings and materials provide a well- supported basis for the inclusion of silver into potential coatings for orthopaedic implants. However, the use of an established ovine model in the current study allows for a direct comparison of outcomes between the novel Ag HA coating and other traditionally effective osseoconductive coatings, in a controlled and repeatable environment.

83 Specifically, the outcomes of a grit-blasted titanium substrate with a 200µm thick plasma-sprayed HA coating utilised in a previously conducted study [57] are of particular relevance to the current study. The GB 200HA implant substrates were manufactured using identical processes and to equivalent dimensional specifications as the titanium substrates used in the Ag HA group.

Both samples were manufactured into cylindrical dowels, from Ti6Al4V alloy, which underwent subsequent a grit-blasting process. The application of the HA coatings was both performed through a plasma-spraying process to a target thickness of 200µm. The only variation between the groups were those necessary to incorporate silver particles into the Ag HA cohorts coating.

Subsequently, the GB 200HA provides an ideal control cohort with which to compare the influence of the addition of silver particles into a HA coating, and the ensuing effects upon osseointegration and bone cell behaviour. A direct comparison of the study parameters and design is presented in Table 10.

Table 10 Comparison between current study and previous HA study in same ovine model

Description Current Study (Ag HA) Previous Study (Conventional HA)

Implant coating Ag HA coating HA coating

Implant Substrate Ti6Al4V Ti6Al4V

Implant size 6mm x 20mm 6mm x 20mm (Diameter x length)

Implant Manufacturer Signature Orthopaedics (Aus) Signature Orthopaedics (Aus)

Target coating thickness 0.2mm 0.2mm

Coating Manufacturer Accentus Medical (UK) Accentus Medical (UK)

Study time points 4 weeks, 12 weeks 12 weeks

84 Cortical sample size per 12 12 time point

Mechanical pushout Mechanical pushout

Study endpoints SEM and histomorphometric SEM and histomorphometric

properties properties

The outcomes of shear stress of pushout and bone implant contact for cortical samples of both groups are presented in Figure 30 and Figure 31. The ovine model used to assess the GB 200HA implants entailed evaluation at 12 weeks follow-up only, therefore comparisons cannot be made between the groups for the 4-week timepoint. When considering the shear strength of pushout, the Ag

HA silver implants demonstrated a mean shear stress of 17.84MPa (SD: 5.11) after 12 weeks. The conventional HA implants had a mean shear stress of 17.89

MPa (SD: 3.47) after 12 weeks. Therefore, there was no significant difference (P

= 0.981) in the shear stresses between each group after 12 weeks post- implantation. Similarly, cortical bone implant contact for the Ag HA implants was found to be 81.5% (SD: 7.6%), whilst for the conventional HA implants it was 78.6% (SD: 4.3%) at 12 weeks. Again, there was no significant difference (P

= 0.262) observed between the two groups for cortical bone implant contact percentages at identical timepoints.

85 Figure 30 Mean shear stress of pushout at 12 weeks in ovine model of Ag HA and Conventional HA coatings

Figure 31 Mean BIC in cortical bone at 12 weeks for Ag HA and Conventional HA coatings

The equivalence of study outcomes between the Ag HA and conventional HA groups strongly supports the notion that the presence of silver particles in the

86 HA coating, at the tested concentration, does not adversely affect the inherent osseointegration properties or bioactivity of the HA. Neither shear stress of pushout nor the extent of bone ongrowth at the coated surface of the implants were significantly altered due to the presence of silver particles in the HA coating. When coupled with the evaluation of cell behaviour and formation through the histological assessment of Ag HA samples, as previously discussed, the silver HA coating process and doping concentration appear to present a viable method of incorporating silver particles into a HA coating.

The equivalent experimental outcomes of the Ag HA group when compared to conventional HA are significant in supporting the viability of Ag HA as a potential coating for orthopaedic implants. Given the known antimicrobial properties of silver ions, the evidence demonstrating the absence of impact upon bone ongrowth and formation indicates that the coating can provide an added form of defence against the risk of infection without compromising the established properties of a HA coating.

4.7 Study Limitations and Future Studies The results of the current ovine study provide a positive preliminary examination of the potential of the Ag HA coating process to incorporate silver particles into a hydroxyapatite coating without adversely affecting the osseointegration properties of the implant. An examination of the studies and literature surrounding the use of silver particles and coating further illustrates the importance of the concentration of silver in achieving a balance between effective antimicrobial activity and maintenance of osseointegration and biocompatibility. It has been shown that excessive silver concentrations may

87 lead to a harmful effect upon the body’s own cells, thereby inhibiting bone ongrowth onto the implant surface despite exceptional bactericidal properties.

This ovine study was limited in that it focussed primarily upon the bone and cell response to the silver HA coating, when compared to other potential implant surface treatments and conventional coatings. Whilst the silver concentrations used in the processes to develop the Ag HA coating are based upon those used in the clinically successful Agluna® material, there was no confirmation or quantification of the actual silver content in each of the implants used in the trial. Further, no analyses regarding the rates and quantities of silver elution from the coating surface were considered within the scope of this ovine study. Combined with the quantification of silver concentration, elution profiles of the silver ions provide a comprehensive illustration of the varying silver concentrations in both the implant coating and periprosthetic joint space.

Future studies using the Ag HA coating should aim to address both of the above study limitations. Silver concentrations on implant surfaces and coatings can be quantified through the use of inductively coupled plasma emission spectroscopy (ICP) [75] to determine the atomic weights of each element present at the implant coating surface. Similarly, the elution profiles of silver ions can be determined through the immersion of implant samples a phosphate- buffered saline (PBS) solution at 37°C [73, 75]. Over time, the silver ions will elute into the PBS, simulating release of ions from the implant into the periprosthetic joint space in vivo. Extracts of the release media can be collected at designated time points (e.g. 1, 2, 3, 5, 9, 15, 30 days etc.) and subject to the

88 same ICP analysis as above to identify the concentration of Ag2+ ions in the solution.

In addition to further quantification of the silver properties within the Ag HA coating, the focus of future studies using this coating should aim to examine the antimicrobial properties of the coating and the anticipated effectiveness in preventing infection. After the findings of the current study illustrating that the Ag HA coating does not hinder osseointegration, the demonstration of antimicrobial effectiveness will confirm the suitability of the coating as a viable solution for a bioactive coating which additionally helps to fight the onset of infection. Alternatively, if the Ag HA coating does not demonstrate adequate anti-microbial effectives, it would serve to further clarify the strong equivalence observed in the shear stress of pull-out and bone implant contact between conventional HA and the Ag HA coating.

Study methods used to assess antimicrobial effectiveness can include testing the effects of released silver ions on bacterial cultures of S. aureus and/or E.

Coli through an agar disk diffusion method [73, 75]. Following inoculation of the plates and exposure to the coated implants, the seeded plates are incubated after which diameters of the inhibition zones can be measured. Further studies to assess the impact upon bacterial morphology can be examined through SEM of bacterial cultures incubated on the surface of implant itself. This method can highlight the precise influence of the silver ions upon cell membranes and structures, and provides a useful comparison of healthy or unaffected bacterial cells in control groups.

89 Chapter 5. Conclusions

Periprosthetic infection remains a prominent cause of failure and revision of orthopaedic implants. The burden to both the patient’s wellbeing and wider socio-economic networks is acutely increased for a revision procedure, when compared to the primary intervention. Consequently, the ongoing development of implant materials and coatings to improve antimicrobial resistance remains a significant pursuit for the orthopaedic field.

This study aimed to evaluate the osseointegration of a novel silver-doped hydroxyapatite coating, and a second acid-etching process, applied to a titanium implant substrate implanted using a well-established ovine model.

Osseointegration was quantified through assessment of the shear strength of mechanical pushout, as well evaluating bone-implant contact percentages through SEM imaging of the implant perimeter. Implantation sites included both cortical and cancellous bone, with samples harvested after both 4-week and 12-week in situ.

At both the 4-week and 12-week timepoints, the silver HA coating displayed significantly greater osseointegration properties, when compared to the acid- etched implants in both cortical and cancellous bone. The mean shear stress of pushout for the silver HA implants was 8.35 MPa and 17.84 MPa at 4-weeks and 12-week respectively. In comparison, the acid-etched implants had mean shear stress of 3.11MPa and 5.22MPa, at 4-weeks and 12-weeks respectively.

Mean bone-implant contact percentages in cortical bone for the silver HA coated implants were 66.03% and 81.50%, compared with 51.77% and 71.77% for the acid-etched implants, at 4-weeks and 12-weeks respectively. Similarly,

90 in cancellous specimens, the mean bone-implant contact percentages for the silver HA coated implants were 54.62% and 56.90%, compared with 30.83% and

40.37% for the acid-etched implants, at 4-weeks and 12-weeks respectively.

While histological analysis of all specimens demonstrated osseointegration and the absence of foreign body reactions in both groups, it further highlighted a greater density and higher continuity of woven bone surrounding the silver HA implants when compared with the acid-etch implants.

The results of this study indicate that the addition of elemental silver does not adversely affect the osseointegration properties of the HA coating. This was reinforced by comparison of the outcomes of this study to that of previous studies examining osseointegration of a conventional HA coating using the same ovine model. Following 12-weeks implantation the shear stress of pushout and bone-implant contact percentage of conventional HA was 17.89MPa and

78.6% respectively. Consequently, there was no significant difference between the osseointegration properties of conventional HA and the silver HA coating.

While the results of this study indicate that the novel silver HA coating has strong potential as an antimicrobial coating, future studies will be required to effectively quantify the antimicrobial effectiveness of the coating.

Characteristics such as elution profiles, silver concentrations and verification that target microorganisms can be destroyed, in addition to the now verified ability to promote osseointegration, are all significant factors in fully assessing the suitability of this novel coating for use on orthopaedic implants.

91 References

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96 Appendix A Histomorphometric Results

Full results of the histomorphometric analysis conducted upon the series SEM images of taken for each sample are presented in the following charts. Each chart presents the results for implants taken from one animal.

The length of bone-implant contact regions at each implant section was calculated in pixels. Given the known pixel length of the entire implant perimeter, the bone-implant contact ratio could then be expressed as a percentage, comparing bone contact length vs total length.

97 Results from W2602

98 Total Length Sec 1 Sec 2 Sec 3 Sec 4 Sec 5 Sec 6 Sec 7 Sec 8 Sec 9 Sec 10 Sec 11 Total BIC TOTAL BIC Sample (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) (px) (%)

L1 13547 0.00 684.51 1233.13 1078.46 1003.46 882.08 474.81 0.00 0.00 0.00 5356.45 39.54

W2602 L2 14423 0.00 277.83 704.09 1085.37 840.72 1031.56 797.50 864.74 666.81 624.69 6893.31 47.79 Canc. R1 12538 406.17 677.79 974.70 1226.22 783.73 970.14 890.75 1130.22 445.20 7504.91 59.86

R2 13528 0.00 0.00 0.00 1052.42 1048.04 703.58 1013.45 0.00 0.00 0.00 3817.49 28.22 L3L 8833 1204.01 922.28 1116.61 1151.44 933.53 1289.78 6617.66 74.92 L3M 8584 1266.53 938.62 997.98 1224.93 1111.62 689.92 6229.59 72.57 L4L 8487 22.11 781.63 1115.25 567.51 365.85 734.95 3587.30 42.27 L4M 7155 1130.76 674.70 321.54 729.92 1118.53 457.61 4433.06 61.96 L5L 9193 1451.86 640.78 1275.77 1229.04 977.53 1350.43 6925.41 75.33

W2602 L5M 9547 1081.37 1409.99 1368.97 830.46 1276.51 858.86 6826.16 71.50 Cort. R3L 7730 760.07 780.58 351.07 568.42 670.51 460.59 3591.24 46.46 R3M 5676 705.61 195.91 863.80 930.25 2695.56 47.49 R4L 8692 874.03 869.22 983.98 1120.42 1099.53 698.96 5646.15 64.96 R4M 8178 1007.61 934.92 143.98 588.06 818.69 927.50 4420.76 54.06 R5L 7045 230.76 631.68 409.85 1112.29 1095.71 3480.29 49.40 R5M 7306 765.61 830.50 728.97 637.43 1196.07 410.31 4568.90 62.54 Results from W2603

99 Total Length Sec 1 Sec 2 Sec 3 Sec 4 Sec 5 Sec 6 Sec 7 Sec 8 Sec 9 Sec 10 Sec 11 Total BIC TOTAL BIC Sample (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) (px) (%)

L1 9900 412.38 814.07 557.19 1120.55 0.00 68.50 0.00 2972.70 30.03

W2603 L2 14230 351.79 684.45 1144.38 758.24 1020.19 1030.92 981.26 795.66 887.88 718.56 8373.33 58.84 Canc. R1 13884 671.18 923.64 495.83 67.60 420.19 1098.27 1090.79 909.54 1098.65 443.62 7219.30 52.00 R2 13077 0.00 0.00 635.50 665.80 732.23 791.84 512.45 0.00 0.00 0.00 3337.81 25.52 L3L 9072 1188.10 435.29 446.97 952.27 1024.36 1357.51 5404.50 59.57 L3M 9104 901.69 899.88 953.56 1159.54 1140.05 389.40 5444.11 59.80 L4L 8900 980.49 643.08 340.75 915.77 762.06 874.99 4517.13 50.75 L4M 6762 27.10 0.00 1279.69 628.16 248.20 2183.15 32.29 L5L 8969 796.39 1557.47 1351.76 1204.52 1351.96 112.46 6374.55 71.07

W2603 L5M 9307 1028.60 835.67 657.29 1165.18 505.77 1077.49 5270.00 56.62 Cort. R3L 7751 714.29 820.04 979.57 650.39 1087.54 380.00 4631.84 59.76 R3M 8908 479.40 420.43 0.00 1229.11 961.70 926.08 4016.72 45.09 R4L 9430 1150.08 1199.52 602.47 1224.57 734.08 1249.72 6160.45 65.33 R4M 9442 1419.19 758.29 1017.17 1078.84 1153.63 866.80 6293.92 66.66 R5L 7992 347.46 706.55 279.36 768.00 1428.39 1172.45 4702.22 58.84 R5M 8576 773.57 683.10 1239.17 505.88 1306.79 1012.47 5520.98 64.38 Results from W2604

100 Total Length Sec 1 Sec 2 Sec 3 Sec 4 Sec 5 Sec 6 Sec 7 Sec 8 Sec 9 Sec 10 Sec 11 Total BIC TOTAL BIC Sample (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) (px) (%)

L1 13683 1060.09 760.49 1002.47 1038.58 586.50 213.46 113.40 587.30 905.79 904.51 7172.59 52.42

W2604 L2 12630 696.78 889.27 847.08 913.66 527.44 911.05 554.30 885.66 0.00 6225.23 49.29 Canc. R1 13929 1312.03 995.99 749.11 1160.34 1004.30 1043.93 984.19 329.69 0.00 0.00 7579.59 54.42

R2 13207 0.00 0.00 0.00 316.16 662.76 557.98 1091.75 0.00 0.00 347.42 2976.07 22.53 L3L 7636 1387.20 1197.50 1118.81 1247.09 1586.96 6537.55 85.61 L3M 3169 939.84 321.59 999.58 209.51 2470.52 77.96 L4L 6012 795.96 1155.60 440.32 1003.33 1080.57 367.49 4843.27 80.56 L4M 5223 720.22 1083.65 220.53 147.46 683.72 156.40 3011.98 57.67 L5L 8500 571.00 1057.92 1325.84 444.54 1377.29 1438.97 1177.50 596.10 7989.16 93.99

W2604 L5M 9490 1120.08 1136.28 894.60 425.37 1071.89 1088.50 1140.76 6877.48 72.47 Cort. R3L 14434 1082.20 1024.00 956.08 1090.84 930.90 480.07 598.05 623.22 1002.29 764.69 819.00 9371.34 64.93 R3M 4404 794.08 859.70 973.25 534.76 3161.79 71.79 R4L 7684 1244.33 1282.50 1337.51 714.41 536.54 1278.72 598.08 6992.07 91.00 R4M 3829 1195.20 407.67 315.94 689.45 2608.26 68.12 R5L 13912 1172.35 1136.52 1102.83 823.67 748.29 1238.11 1202.58 1035.18 774.33 740.24 9974.10 71.69 R5M 5287 370.19 648.79 642.94 507.58 854.07 3023.56 57.19 Results from W2605

101 Total Length Sec 1 Sec 2 Sec 3 Sec 4 Sec 5 Sec 6 Sec 7 Sec 8 Sec 9 Sec 10 Sec 11 Total BIC TOTAL BIC Sample (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) BIC (px) (px) (%)

L1 14378 0.00 518.76 454.19 699.15 333.93 0.00 200.58 884.41 969.95 596.90 0.00 4657.87 32.40

W2605 L2 14368 788.41 1017.36 1088.60 980.03 904.56 543.83 1238.40 1136.43 752.99 1113.19 9563.79 66.56 Canc. R1 4148 430.59 778.25 853.93 2062.77 49.73

R2 11789 0.00 208.10 1152.75 1113.03 639.84 882.76 875.10 1066.26 443.03 6380.88 54.13 L3L 11169 1102.83 1350.77 1303.64 1313.05 1198.23 1020.29 1241.08 774.30 9304.18 83.30 L3M 2970 1100.77 897.60 288.78 2287.15 77.01 L4L 10798 788.18 975.10 723.16 1138.77 900.03 984.00 920.81 859.86 7289.92 67.51 L4M 5702 790.74 620.25 894.43 1177.34 714.17 4196.93 73.60 L5L 5176 1215.05 1125.55 1064.25 1084.76 4489.60 86.74

W2605 L5M 10217 1082.21 1034.10 897.95 656.03 1081.20 1022.13 1171.52 952.63 7897.78 77.30 Cort. R3L 5492 1193.20 903.76 216.25 648.87 1185.02 4147.10 75.51 R3M 6487 810.82 1155.75 832.41 893.91 755.55 462.22 4910.66 75.70 R4L 5056 1020.74 1258.93 372.00 603.00 1079.23 4333.89 85.72 R4M 6266 1150.91 1101.48 1068.41 1032.87 579.45 4933.11 78.73 R5L 6918 778.24 1114.16 991.76 1176.76 967.10 983.35 6011.37 86.89 R5M 6403 1128.90 724.82 700.85 698.40 1243.32 579.31 5075.60 79.27