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Effects of alendronate-immobilized calcium phosphate coating on bone growth into porous tantalum : a… Hu, Youxin 2007

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EFFECTS OF ALENDRONATE-IMMOBILIZED CALCIUM PHOSPHATE COATING ON BONE GROWTH INTO POROUS TANTALUM —A GAP MODEL ANIMAL STUDY by YOUXIN HU M.A.Sc. in Materials Science & Engineering, Northeast Univ. of Tech. China, 1990 B.A.Sc. in Materials Science & Engineering, Northeast Univ. of Tech. China, 1985 A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF MASTER O F APPLIED SCIENCE in THE FACULTY OF GRADUATE STUDIES (Materials Engineering) THE UNIVERSITY OF BRITISH COLUMBIA April 2007 © Youxin Hu, 2007 ABSTRACT Porous tantalum has been shown to be a promising orthopaedic implant material because of its similarity to bone in both mechanical properties and its three-dimensional porous structure. However, in some circumstances, bone quality or quantity is insufficient to allow adequate bone ingrowth. Alendronate, one of the bisphosphonate families, affects the activities of bone cells and enhances bone formation. In this thesis, we hypothesized that the addition of alendronate could increase the osteoconductivity and bone-ingrowth of porous tantalum and overcome the challenges of bone-implant gaps. To facilitate local delivery of alendronate, a micro-porous calcium phosphate coating was deposited onto the tantalum surface by an electrolytic deposition technique, which was followed by alendronate adsorption. Coating structures and morphologies were confirmed by scanning electron microscopy, X-ray diffraction, and Fourier transform infrared spectroscopy. The presence of alendronate was confirmed by high performance liquid chromatography. To study the effects of alendronate-immobilized calcium phosphate coating on bone reaction to porous implants, an animal gap model, with a fixed gap of 0.6 mm between implants and bone, was developed. Three types of surfaces, which were non-coating (Ta), calcium phosphate coating (Ta-CaP), and alendronate-immobilized calcium phosphate coating (Ta-CaP-AIN), were compared. Two fluorochromes were adopted to track the front of bone formation. After four weeks of healing and following standard histology techniques, the implants were analyzed with backscattered electron microscopy and fluorescent optical microscopy for bone-implant interactions. i i The relative volume increase of gap filling, bone ingrowth and total bone formation were 124 % (2.24-fold), 232% (3.32-fold) and 170% (2.7-fold) respectively in T a - C a P - A L N compared with Ta controls. The contact length of newly formed bone on porous tantalum was increased by 700% (8-fold) in T a - C a P - A L N compared with Ta plugs, suggesting enhanced osteoconductivity of T a - C a P - A L N implants. The bone formation mechanism analysis found that bone growth initiated on both surfaces of the T a - C a P - A L N implants and host bone, while little bone initiation on the Ta implant surface was detected. These significant enhancements of T a - C a P ^ A L N may have direct applications in orthopaedics. For revision arthroplasty with insufficient bone stock, the local delivery of alendronate would enhance biological fixation of the implant and promote the healing of bone defects. i i i TABLE OF CONTENTS A B S T R A C T i i T A B L E O F C O N T E N T S iv L I S T O F T A B L E S v i i L I S T O F F I G U R E S v i i i L I S T O F A B B R E V I A T I O N x i i i A C K N O W L E D G M E N T S xiv C H A P T E R 1: I N T R O D U C T I O N 1 C H A P T E R 2: L I T E R A T U R E R E V I E W 4 2.1 Introduction to artificial hip joints 4 2.2.1 Structure of bone 6 2.2.3 The mechanical properties of bone 9 2.2.4 Bone formation and bone healing process 10 2.3 Implant materials 11 2.4 The challenges in hip replacement 14 2.5 Current material approaches to improve implant-bone fixation 16 2.6 Calcium phosphate 17 2.7 Deposition of hydroxyapatite 19 2.7 Effect of bisphosphonates on bone 22 2.8 Current studies on bisphosphonate-delivery implants 23 2.9 Gap model animal tests 24 2.10 Mechanism of labeling 25 2.11. t-tests 26 C H A P T E R 3: S C O P E A N D O B J E C T I V E S 29 iv C H A P T E R 4: M A T E R I A L S A N D M E T H O D S 31 4.1 Implant preparation 31 4.1.1 Porous tantalum 31 4.1.2 Electrolytic deposition ( E L D ) of calcium phosphate coatings 33 4.1.3 Alendronate immobilization 36 4.1.4 Implant assemblies 37 4.1.5 Sterilization of implants and surgical tools 39 4.2 High performance liquid chromatography ( H P L C ) tests for drug loading 40 4.3 Animal tests 40 4.3.1 Implantation procedures 42 4.3.2 Bone labeling 43 4.3.3 Animal care after surgery 44 4.3.4 Euthanasia, radiograph imaging and sample harvest 45 4.4 Sample analysis 45 4.4.1 Histology sample preparation 45 4.4.2 Histology and histomorphometry analysis 47 C H A P T E R 5: R E S U L T S 51 5.1 The effects of coating parameters on calcium phosphate coating morphology 51 5.2 Calcium phosphate coating and alendronate loading 54 5.2.1 E L D coating of calcium phosphate 54 5.2.2 The alendronate loading amount and estimation of dosage 56 5.3 Crystal structure of E L D calcium phosphate coating 56 5.4 Histology 59 v 5.5 Histomorphometry 61 5.5.1 Bone formation 61 5.5.2 Bone/implant contact length 64 5.6 Bone formation mechanism 66 C H A P T E R 6: D I S C U S S I O N 73 6.1 The effects of calcium phosphate coating 73 6.2 The effects of alendronate-immobilized calcium phosphate coatings 75 6.3 Clinical relevance of the gap model animal tests and alendronate delivery implants ...77 6.4 Key progresses in E L D surface drug delivery and animal studies 79 C H A P T E R 7: C O N C L U S I O N S 81 C H A P T E R 8: R E C O M M E N D A T I O N F O R F U T U R E W O R K 83 R E F E R E N C E S 84 A P P E N D I C E S 92 Appendix A : Student's t table (reprinted from ref 82 courtesy of Deacon J.) 92 Appendix B : Ethical approval 94 v i LIST OF TABLES Table 1.1 Total numbers of hip and knee replacement procedures, Canada [3] 1 Table 4.1 The weight of porous tantalum implant plugs 32 Table 4.2 Animal test plan for bisphosphonate-coated porous Ta bone implants 41 Table 4.3 Information of fluorochomes and their solutions 44 Table 4.4 Labeling administration dosages and schedule 44 Table 4.5 Processes and schedules for fixation, dehydration and infiltration 45 Table 5.1 The statistic results of bone formation 64 Table 5.2 The statistic result on contact length 65 v i i LIST OF FIGURES Fig . 2.1 The illustration of an artificial hip joint, 5 Fig. 2.2 The acetabular shell made of porous tantalum (reprinted from ref. 14, with Fig . 2.3 The 7 hierarchical levels of organization of the bone family of materials. Level 1: Isolated crystals from human bone (left side) and part of an unmineralized and unstained collagen fibril from turkey tendon observed in vitreous ice in the TEM(right side). Level 2: T E M micrograph of a mineralized collagen fibril from turkey tendon. Level 3: T E M micrograph of a thin section of mineralized turkey tendon. Level 4: Four fibril array patterns of organization found in the bone family of materials. Level 5: S E M micrograph of a single osteon from human bone. Level 6: Light micrograph of a fractured section through a fossilized (about 5500 years old) human femur. Level 7: Whole bovine bone (scale: 10 cm), (reprint from ref 16 with permission from Annual Review) 8 Fig . 2.4 Hierarchical structural organization of bone: (a) cortical and cancellous bone; (b) osteons with Haversian systems; (c) lamellae; (d) collagen fiber assemblies of collagen fibrils; (e) bone mineral crystals, collagen molecules, and non-collagenous proteins, (reprinted from ref 17 with permission from Fig. 2.5 Logarithm of the product of calcium and phosphate concentrations plotted against p H values of solutions saturated with respect to various calcium phosphate phases in the ternary system Ca(OH)2-H 3 P0 4 -H20 Calculated for permission from J of Bone and Joint Surg.) .5 Elsevier) 9 v i i i 37°C.(reprinted from ref 50 with permission from American Chemistry Society Publications) 18 Fig. 2.6 S E M images of biomimetic coatings of CaP on porous tantalum (reprinted from ref. 56, with permission from Springer) 20 Fig. 2.7 S E M image of the E L D coating of calcium phosphate 22 Fig . 2.8 Chemical structure of alendronate (reprinted from ref. 66, with permission from Springer) 23 Fig . 4.1 S E M images of porous tantalum cylinder at various magnifications 32 Fig . 4.2 The image of E L D coating cell 34 Fig. 4.3 Mechanical drawings of porous tantalum implant assembly. The left is a 3D image of the implant assembly, and the right is a 2-D drawing with all the dimensions 38 Fig . 4.4 The image of bone cement casting mold 38 Fig . 4.5 Surgical tools and porous tantalum implant 39 Fig . 4.6 The images and schema of the implant location, a) digital image of implant position during surgery; X-ray image shows the implant position after surgery; c) the schematic image 43 Fig. 4.7 Schema for sectioning and observation directions 46 Fig . 4.9 Schema image for 3 grinding sections 46 Fig . 4.9 Histogram for image analysis and the images before and after dye process 49 Fig . 4.10 The measuring images for contact length and the available metal length 49 Fig. 4.11 The schema drawings illustrate the bone ingrowth and gap filling 50 ix Fig . 5.1 S E M images of calcium phosphate coatings, a) without sulfuric acid cleaning, the porous tantalum strut was not coated; b) spallation of calcium phosphate coatings; c) coatings morphologies at central pores without periodical stirring; (d) coating morphology at superficial layer when periodical stirring was not applied 52 Fig . 5.2 S E M images on calcium phosphate coatings at a central pore when periodical stirring was applied 53 Fig . 5.3 S E M images at varying magnifications of Ca-P coating morphology, a) digital image of implant assembly; b) low magnification S E M image shows the coverage of CaP coating; c) and d) the porous morphology of CaP coating 55 Fig. 5.4 S E M image showing coating thickness 55 Fig. 5.5 X R D results for CaP coatings before and after alendronate immobilization 57 Fig. 5.6 The FTIR results for CaP coatings before and after alendronate immobilization 58 Fig . 5.7 B S E and fluorescence images of bone formations on 3 types of implants, a) and b) non coated control; c) and d) calcium phosphate coated control; e) and f) alendronate immobilized calcium phosphate coated implant 60 Fig . 5.8 The mean and standard deviation of volume of gap fill ing, bone ingrowth and total bone formation. T a - C a P - A L N implants demonstrated significantly increased gap filling, bone ingrowth and total bone formation compared with Ta-CaP controls (p < 0.0005, p < 0.0001 and p < 0.0001 x respectively *), and significantly increased gap fill ing, bone ingrowth and total bone formation compared with Ta control (p < 0.0001, p < 0.0001 and p< 0.0001 respectively **). N o significant differences in gap fill ing, bone ingrowth and total bone formation were found between Ta-CaP and Ta. (p = 0.12, p = 0.07, p = 0.07) 63 Fig . 5.9 Bone formation on the surface of porous tantalum expressed as a percentage of the available porous tantalum surface (contact length) of three sections performed through each implant. T a - C a P - A L N demonstrated significantly increased contact length compared with Ta-CaP controls, p < 0.0001 *, and significantly increased contact length compared with Ta controls, p < 0.0001 **. Ta-CaP controls demonstrated significantly increased contact length compared with Ta controls, p < 0.02 + 66 Fig . 5.10 Fluorescence images showing the labels on tooth and ingrowth bone, and the ordinal numbers gave postoperative time sequence, a) 4 color rings indicated the growth direction on tooth. The 2 green rings were calcein and 2 red rings were alizarin; b) labels indicated the bone initiation on implant surface 67 Fig . 5.11 Fluorescence images showing bone initiation, a) non coated porous tantalum control; b) alendronate immobilized calcium phosphate coated implant; c) and d)calcium phosphate coated controls 69 xi Fig. 5.12 Fluorescence images showing bone ingrowth pattern on Ta control, a) low magnification image; b) higher magnification image showing the labeling lines 70 Fig . 5.13 Fluorescence images showing bone ingrowth pattern on Ta-CaP control 70 Fig . 5.14 Fluorescence images showing bone ingrowth pattern on T a - C a P - A L N implant ...70 Fig. 5.15 Fluorescence images showing the bone formation at gap areas. The white lines with double arrows indicated the bone bridging gaps a) Ta control; b) Ta-CaP control; c) and d) T a - C a P - A L N implants 71 xii LIST OF ABBREVIATION A L N Alendronate B M P Bone morphogenic protein B P P Bisphosphonate B S E Backscattered Electron imaging CaP Calcium phosphate D C P D Dicalcium phosphate dihydrate E C D Electrochemical deposition E D S Energy dispersive spectrometry E L D Electrolytic deposition FTIR Fourier transform infrared spectroscopy H A Hydroxyapatite H L P C High performance liquid chromatography O C P Octacalcium phosphate P M M A Polymethylmethacrylate SD Standard deviation S E M Scanning electron microscopy Ta Tantalum Ta-CaP Calcium-phosphate-coated porous tantalum T a - C a P - A L N Alendronate-immobilized calcium-phosphate-coated porous tantalum T C P Tricalcium phosphate T H R Total hip replacement X R D X-ray diffraction x i i i ACKNOWLEDGMENTS I would like to thank my supervisor, Dr. Rizh i Wang for his guidance during my research work, and Dr. Don Garbuz for supervising surgical protocols during animal tests. I also would like to thank my colleagues, K e Duan for his help throughout the entire project; Dr. Winston K i m for his excellent work on implant surgeries; Dr. Tom Oxland, Bas Masri and Clive Duncan for support in surgery designs; and Al l en Tang for his support in biomaterial laboratory and in the Jack Bel l Centre of the Vancouver General Hospital. I would like to extend my thanks to Dr. Helen Burt and M r . John Jackson in the Pharmaceutical Science Department of U B C ; M s . Mary Mager and other staff in the Materials Engineering Department of U B C and the animal care team at Jack Be l l Centre in V G H . Finally, Zimmer Inc. and N S E R C , thank you for your support with the project. xiv CHAPTER 1: INTRODUCTION Osteoporosis and osteoarthritis may cause severe damage to the hip and/or knee joints. Injuries occurring during sports and exercises may also break the femoral necks or the knee patellae. To treat these diseases and traumas, orthopaedists often have to replace the failed hip or knee joints with artificial ones made of metals, polymers or ceramics. Women are more vulnerable to osteoporosis [1]. After menopause a woman can lose up to 20% bone mass in 5 to 10 years [2], and this may cause brittleness of bone. In the year 2004 - 2005, statistics showed that the number o f knee replacements in Canada reached 33.6 thousand, which was an increase of 125% from the year 1994-1995, Tab 1.1. Also in 2005, hip replacements in Canada reached 25 thousand, an increase of 52% and 6% compared with years 1994 -1995 and 2003 - 2004 respectively [3]. Considering that the population of Canada in 2005 was 30 millions, almost two in every thousand Canadians have had either a hip or knee replacement in every year for the past 10 years. Therefore, any improvement in knee or hip replacements w i l l significantly contribute to the Canadian health care system and dramatically benefit Canadian society. Table 1. 1 Total numbers of hip and knee replacement procedures, Canada [3] 1994-1995 2003-2004 2004-2005 10-year % Change 1 - Year % Change Knee Replacements 14,938 29,848 33,590 125% 13% Hip Replacements 16,525 23,669 25,124 52% 6% Total Replacements 31,463 53,517 58,714 87% 10% 1 Currently, poor fixation and unacceptably high revision rates are the major problems with knee and hip joint replacements. Poor early bone/implant integration, stress shielding and aseptic loosening are the major causes of failure of joint replacement implants [4, 5], and the longevity of these joint implants is 15 to 20 years. With the extension of human lifespan to 80 or 85 years in the 21 s t century, a revision joint arthroplasty may be needed. Therefore, enhancing the bone/implant fixation and reducing the revision rates of joint replacements remain the primary topics in orthopaedic research. In addition, the management of bone stock deficiency and the achievement of reliable and durable implant/host-bone fixation remain significant challenges in revision joint replacements. The concept of a drug-implant delivery system in reconstructive orthopaedics was well established [6]. Bisphosphonates have been investigated as a means to promote early bone ingrowth [7 , 8, 9] . Recently, an investigation of a canine model where a third generation bisphosphonate, zoledronic acid, bound to hydroxyapatite-coated porous tantalum implants demonstrated an increase in peri-implant bone growth within the intramedullary canal and a 58% increase in the bone ingrowth within the porous tantalum[10]. Laboratory experimental models should replicate clinical scenarios as closely as possible. In revision hip arthroplasty, host bone deficiencies, gaps and defects between implant and host bone are frequently encountered. Gaps as small as 50 to 500pm failed to be bridged by new bone in a study of human host bone and hydroxyapatite-coated implants [11]. The design of implant materials has to face challenges of the inevitable gaps between bone 2 and implants during arthroplasty surgeries, and the gap model animal tests are necessary for examining new designed implant materials. Although the promising results on bisphosphonate-delivery implants were reported [10], the effects of bisphosphonate-immobilized calcium-phosphate coating on the location, pattern and extent of bone formation in the presence of a gap have not been previously investigated. We conducted a gap model animal study on alendronate-immobilized calcium-phosphate-coated porous tantalum implants. The calcium phosphate coatings we processed via electrolytic deposition were porous at 1 micrometer level, and the porous morphology was good for bisphosphonate immobilization. L o w concentration (less than 10"4 mol/1) of alendronate, which was one o f the bisphosphonate families, was immobilized on the surfaces of E L D calcium phosphate coating. We hypothesized that the alendronate-immobilized calcium phosphate coating could increase the osteoconductivity and bone ingrowth of porous tantalum implants and therefore enhance the peri-prosthetic bone growth. Eventually, better bone/implant fixation could be achieved. The purpose of this study was (1) to develop an experimental animal model which would simulate the clinical revision hip scenario, (2) to determine the effects of alendronate-immobilized calcium phosphate coating on porous tantalum on gap filling and bone ingrowth in a rabbit model, and (3) to develop techniques of improving the bone/implant fixation of porous tantalum joint replacements. 3 CHAPTER 2: LITERATURE REVIEW 2.1 Introduction to artificial hip joints Since the 1960s, when the artificial hip joint with polymer liner was developed and applied by John Charnley [12], many improvements have been made, and the total hip replacement (THR) has become an industry worth billions of dollars. Nowadays, many different materials such as cobalt-chromium alloys, titanium alloys and tantalum are being used to ensure the biocompatibility of T H R , and different rough surface treatments such as beads and meshes are employed to enhance bone attachment to T H R implants. The biocompatibility here is defined as "the ability of a synthetic or natural material for medical implantation to receive the appropriate reaction from recipient environment" [13]. A n artificial hip joint typically includes an acetabular shell for attaching to the pelvis, a polyethylene liner for reducing friction and an artificial femoral head and stem for inserting into the femoral intramedullary canal. One example of these hip joint replacements is a titanium alloy stem with dents of 2 to 3 millimeters and an acetabular shell with porous titanium alloy fiber mesh on top(Fig. 2.1). Recently, porous tantalum, also known as trabecular metal, was adopted for manufacturing the acetabular shell (see Fig . 2.2) [14]. This sponge-like porous tantalum had 400 to 500 micrometer pores that allow bone to grow into them. The outcome of this implant design was the mechanical interlock of the pelvis bone with the acetabular shell, resulting in improved bone / implant fixation. 4 Fig . 2.1 The illustration of an artificial hip joint Fig . 2.2 The acetabular shell made o f porous tantalum (reprinted from ref. 14, with permission from J of Bone and Joint Surg.) 2 .2 Bone Knowledge of bone, such as its physical and chemical properties, structures, and reactions to wounds etc., is essential to developing joint replacements. Bone is the main in vivo environment that joint replacements deal with, and bone is also the target that implants should imitate. Also , bone w i l l undergo a healing process due to the interference caused by implantation. A brief review is provided to illustrate the structure, mechanical properties, and the formation and healing process of bone. 2.2.1 Structure of bone The structural components of bones are calcium minerals, collagen proteins, water, and some non-collagen proteins. The two main components of bone are minerals and collagen proteins, which contribute to more than 90 % of its weight [15,16]. The bone minerals are primarily carbonated apatite plates with dimensions of up to 50 nm by 25 nm by 3 nm. The collagen proteins are mainly type I and fibrous, and assembled from triple-helical collagen molecules that are 300 nm in length by 1.2 nm in diameter (see F ig 2.3 Level 1 and Fig 2.4 Sub-nanostructure). The mineral plates and collagen fibers assemble together to form the mineralized collagen fibrils [17]. The mineralized collagen fibrils are the building blocks of bones (see F ig 2.3 Level 2 and F ig 2.4 Sub-nanostructure). These mineralized fibrils are 50-100 nm in diameter and up to a few millimeters in length. Bundles of collagen fibrils form individual sheets, or bone lamellae, which stack together to make the lamellar structure [16]. The Haversian system, or the secondary osteon (Fig 2.3 Level 5 and Fig 2.4 Microstructrure), is a small unit of bone with central blood canals (Haversian canals, about 50 um in size) surrounded by bone lamellae. When bone becomes old or fractured, bone cells start the remodeling process [18, 19]. Osteoclasts resorb the bone and generate the tunnels. Following osteoclasts, osteoblasts rebuild the bone layer by layer as lamellar bones. The central tunnel also functions as a blood vessel with smaller channels of canaliculi connected to it. The canaliculi also connect to osteocytes and supply them with nutrition. A t the macroscale level (Fig 2.3 Level 6 and Fig 2.4 Macrostructure), there are two types of bones classified on the basis of porosity and unit microstructure. Cortical bone, also known as compact bone, is much denser with a porosity of 5 and 10%. It consists of Haversian systems and has fewer blood vessels and bone cells inside it. Also , cortical bone has higher mechanical properties (Young's modulus: less than 30GPa). In comparison, spongy bone, also known as 'cancellous bone' or 'trabecular bone', has a porosity of 50% to 90%). Trabecular bone has much more blood vessels and bone cells, but its mechanical properties such as strength and modulus are lower (Young's modulus of architectural trabecular bone: 0.8GPa). The pelvis bone is mainly made of the trabecular bone. For a long bone such as a femur, the shaft is mainly made of cortical bone surrounding the marrow. Trabecular bone is found at the two ends (distal and proximal) of the femur next to the joints. The pelvis and the femur are the places where T H R s are located. 7 Level 1 : Major Components Fig. 2.3 The 7 hierarchical levels of organization of the bone family of materials. Level 1: Isolated crystals from human bone (left side) and part of an unmineralized and unstained collagen fibril from turkey tendon observed in vitreous ice in the TEM(right side). Level 2: T E M micrograph of a mineralized collagen fibril from turkey tendon. Level 3: T E M micrograph of a thin section of mineralized turkey tendon. Level 4: Four fibril array patterns of organization found in the bone family of materials. Level 5: S E M micrograph of a single osteon from human bone. Level 6: Light micrograph o f a fractured section through a fossilized (about 5500 years old) human femur. Level 7: Whole bovine bone (scale: 10 cm), (reprint from ref 16 with permission from Annual Review of Materials Science) 8 Cortical bone Cancellous bone Collagen molecule L Lamella Collagen fiber Osteon Haversian canal Bone Crystals 10-500 Jim Microstructure Nanostructure Macrostructure Sub-microstructure Sub-nanostructure Fig. 2.4 Hierarchical structural organization of bone: (a) cortical and cancellous bone; (b) osteons with Haversian systems; (c) lamellae; (d) collagen fiber assemblies of collagen fibrils; (e) bone mineral crystals, collagen molecules, and non-collagenous proteins, (reprinted from ref 17 with permission from Elsevier) 2.2.3 The mechanical properties of bone The mechanical properties [20] of lamellar bone are anisotropic. Along the long bone axis the mechanical properties such as strength and modulus etc. are higher than those transverse to long bong axis. With the tested angles varying from 0 to 90 degree, the mechanical properties reduce by 3 to 5 times. For example, the bending strengths reduce from 350MPa or higher at 0 degree to lOOMPa at 90 degree; flexural modulus decrease from 16GPa to 7GPa; and the work of fracture (the energy to break the material) reduces from 8kJ/m 2 to less than l k J / m 2 [21]. The reason for these differences is because of the fibril structure inside lamellar bone, in which collagen fibrils are parallel to the long bone axis. Also , the body weight applies loadings on the long bone axis, and so they are expected to be stronger in that direction. 9 2.2.4 Bone formation and bone healing process Bone healing is a complex biological process [22]. It needs the coordinated actions of growth factors, osteoclasts, osteoblasts and immune cells, etc. within the bone marrow. Stem cells that are recruited from the surrounding tissues and the circulation wi l l differentiate into osteogenic cells. Multiple factors control these chains of reactions and further affect the different sites in the osteoblast lineage through various processes such as migration, proliferation, differentiation, inhibition, and protein synthesis. Cortical bone and trabecular bone experience different healing processes [23] . Trabecular healing depends on osteoconduction and de novo bone formation. On the other hand, cortical healing relies on the remodeling of bone osteons. The de novo bone formation can be divided into four-steps [24]. First, differentiating osteogenic cells, which are the cells still possessing traveling capacity but w i l l become osteoblasts, produce collagen-free organic matrix that provides nucleation sites for calcium phosphate mineralization. Then, calcium phosphate nucleation happens on the collagen-free organic matrix, and the initiation of collagen fiber assembly follows as the third step. Finally, calcification of the collagen compartment w i l l occur. Cortical healing process can be divided into five major periods including inflammatory, granulation tissue formation, callus formation, lamellar bone deposition and remodeling [25]. And osteoclasts involve in bone healing at two different periods. The population and activation of osteoclasts are extremely high at the early stage of bone healing 10 because the major function of osteoclasts is to resorb bones. They clean up the broken bone and supply minerals and proteins for the formation of new bones. The second period we see the high activation of osteoclasts is in the bone remodeling process [25]. The newly formed bone during fracture healing is woven bone, which does not have the regular bone structure. The woven bone can be regenerated fast and bear certain loads. This helps to quicken the healing of injuries. The woven bone w i l l be remodeled later to the regular structure with osteons, canaliculi and nerves, etc. A t the remodeling stage, osteoclasts resorb the woven bone and create tunnels, and then the osteoblasts regenerate the lamellar bone surrounding those tunnels. The combination of osteoclasts and osteoblasts activities generates osteons and completes the bone healing process. Contact osteogenesis and distance osteogenesis [23, 26] are the.criteria used to determine whether implant surfaces are osteoconductive (defined on page 12). Contact osteogenesis means that the bone formation happens on the surface of or around the implants, and distance osteogenesis is when the bone starts growing at the site of host bone. If the implant surfaces are osteoconductive, they should induce the bone formation towards them, and both contact and distance osteogenesis w i l l occur. Otherwise, the contact osteogenesis w i l l not happen, and only distance osteogenesis w i l l take place. 2.3 Implant materials Material selection is very important in artificial joint replacements. The physical and chemical properties of the chosen materials have to fit in the applied physiological 11 environment and meet the mechanical requirements. Corrosion of implants in an in vivo environment is a great concern [27]. On the another hand, the implanted materials must establish a firm interface with the adjacent tissues such as bone, and satisfy the loading needs of natural bone. The materials suitable for artificial joints include several metals or metal alloys such as stainless steels, titanium, titanium alloys, cobalt-chromium alloy, and tantalum. A s these metals or alloys are chosen mainly to fit the biomechanical requirements, the interaction with biological environment is compromised. These metallic implants are usually bioinert and have very limited interactions with surrounding bone tissues. Achieving a strong implant/bone interface has been a challenge [28, 29]. Concepts such as osteoconductivity and osteoinductivity need to be clarified. Both osteoconductivity and osteoinductivity are indicators of the capabilities of bone formation onto bone grafts and implants. Osteoconductivity [30] defines the ability which the graft (porous implant as well) supports the attachment of new osteoblasts and osteoprogenitor cells. It provides an interconnected structure through which new cells can migrate and new vessels can form. Osteoconductivity is the ability of bone graft (or porous implant) to enable new bone to grow through it. Osteoinductivity [30] refers jto the ability of a graft (or porous implant) to induce non-differentiated stem cells or osteoprogenitor cells to differentiate into osteoblasts. 12 Tantalum is one of the metallic materials that have excellent corrosion resistance [31]. It hardly stimulates adverse biological response. Many studies have demonstrated the excellent biocompatibility of tantalum as in bone implants [31, 32, 33]. Solid tantalum may not be good for bone implants because of its high density (16.7 grams/cm 3). The implants made of solid tantalum are much heavier than bone they replace and bring much higher loads to the bone skeleton. The high modulus (200GPa) of solid Tantalum (in comparison to modulus of cortical bone: less than 30GPa) may also cause a 'stress shielding' problem (see page 15). Porous Tantalum is an ideal implant material that combines its excellent biocompatibility with its mechanical properties similar to those of bone. The clinically available porous tantalum has 75% to 80% porosity, and its elastic modulus is 3 GPa [34]. "The compressive and shear strengths of this porous tantalum are 35 to 40 M P a . " [31, 34] Another advantage of porous tantalum is that bone can grow into porous tantalum and construct an interlock system between bone and porous tantalum implants, and the interlock between bone and implant resulted in better bone/implant fixation [34, 35]. The manufacture of porous tantalum is done through a chemical vapor deposition process. Obtaining the desired porosity in the carbon skeleton is achieved through pyrolysis of polyurethane sponge. The tantalum is deposited on the carbon skeleton by chemical vapor deposition with a tantalum pentachloride precursor [34]. 13 2.4 The challenges in hip replacement Poor implant/bone fixation and an unacceptably high revision rate are the major problems with current joint replacements [4, 5]. Aseptic loosening [36], which is caused by periprosthetic bone resorption (also called osteolysis) triggered by the polyethylene wearing debris entering the joint space [37], is one of the main causes of the implant failure. Another reason is 'stress shielding' which is caused by the metallic materials used for femoral stems. The present implantation technology is adequate for pain relieving and movement restoring to the ailing hip joints, and the hospitalization time is usually 6 days and full recovery takes about 4 to 6 months [38]. However, many challenges remain unsolved. A s the lifespans of human beings continue to increase in the 21 s t century, the number of patients who need T H R w i l l also increase because of the bone problems such as osteoporosis and osteoarthrisis. In 2001, the waiting time for T H R surgery in Canada was from 6 to 23 months depending on which province patients lived in [39]. Patients have to endure the serious pain or sit in wheelchairs waiting for their arthroplastic surgeries. A t the same time, the designed implants should extend their longevity to meet the requirement of long post-surgery l iving time. The existing hip replacements normally can serve patients up to 15-20 years and 20 percents of these hip replacements fail within 20 years. This may cause serious pain to the patients, and new revision surgeries w i l l be needed. The biological processes of periprosthetic bone resporption (leading to aseptic loosening) include phagocytosis on the wear particles by macrophages, ingestion of debris 14 particulates and eventually the bone resorption [37, 40]. When wear particles are present, the instinctive reaction (called phagocytosis) of the immune system is to deploy the macrophages to consume these alien particles. Then, the ingestion of particulates of debris starts, and the cytokines and other mediators-of inflammation are released simultaneously. These factors then lead cytokines, and other biological components, to activate osteoclasts to resorb bone around the bone/implant interface. Other than polyethylene particulates, metal particles and bone cement particulates also contribute to osteolysis in the same way. The wear debris associate with periprosthetic osteolysis poses a long-term threat to implant longevity [41]. Stress shielding [42] is associated with the load sensing capability of bone and poses a challenge to hip replacements. The lower limbs sustain the load from the body weight and the muscles attached to the long bones exert stresses on the skeleton. Bone is a 'smart material' in that it grows more bone for supporting where the load on bone is larger. For the same reason, it w i l l grow less bone where the load on bone is smaller. The femoral stem of a hip implant (for example, the Young's modulus of T i alloy is 115GPa) is much stiffer than the cortical bone (Young's modulus is less than 30 GPa) and w i l l take the greater part of the load due to body weight. A s a consequence, the femoral bone around the stem is "unloaded", and the femoral bone at the tip of femoral stem is heavily loaded. The loading changes by introducing joint implants to the skeleton induce bone density changes throughout the femur. The upper part of the femoral bone w i l l contain less bone tissue and become more susceptible to fracture. The bone at the tip of implant stem becomes thicker and stronger. Unfortunately, the thickening of the skeleton is often painful [43] . The patients with 15 cementless stems of total hip devices often feel pain in the thigh, especially during the first year after the surgery. A major challenge in revision hip arthroplasty is achieving early and durable implant-host bone fixation, especially in the presence of significant bone stock deficiencies. Traditional uncemented implants (porous -coated and titanium fiber mesh) have shown good results clinically in the revision setting. However in cases where less than 50% host bone contact is made the results of these uncemented implants has been less predictable [44]. The message from orthopaedists is that these bone deficiencies, bone/implant gaps and defects are inevitable during the arthoplastic surgeries. Autopsy studies have demonstrated that bone ingrowth occurs in only 30-40% of the surface area of the implant in primary total hip arthroplasty [31, 45]. In revision arthroplasties, bone defects and gaps may significantly compromise the quantity and quality of host bone available for ingrowth. In many cases less than 50% of the host bone is available for ingrowth of the newly implanted shell. To any porous implant material, the reliable and durable bone ingrowth with limited host bone contact is considered a desirable improvement. The improvement would result in an implant that is more versatile in the challenges of revision cases. 2.5 Current material approaches to improve implant-bone fixation To achieve early implant-bone fixation, many new approaches have been launched to enhance the bone growth on the metal implant surface. Hydroxyapatite, bone morphogenic proteins ( B M P ) and other materials are introduced either for their similarities of crystal 16 structure to bone minerals or for their biological functionalities in the bone formation processes. Sumner et al [46, 47] have studied the effects of B M P on post-surgical bone formation in a canine model. They dipped the implants into different types of B M P s , such as rhBMP-2 , before the implant surgeries and examined the bone growth 4 weeks after surgery. The results were positive on enhancements of B M P s on bone formation. Recently, they conducted new animal tests of a gap model with combined addition of B M P - 2 and the transforming growth factor P 2 (rhTGF-(32) on hydroxyapatite-plasma-sprayed titanium implants [48]. Compared with non-treated control samples, significant increases in bone formation towards implants were reported on B M P s treated implants. The downsides of B M P treated implants are their low efficiency due to the quick liberation of B M P , the unexpected calcification of surrounding soft tissue, the challenges of obtaining growth factors, storage and transportation, and their low versatility. 2.6 Calcium phosphate There are several forms [49] of calcium phosphate. Dicalcium phosphate dihydrate ( D C P D ; C a H P 0 4 - 2 H 2 0 ) , tricalcium phosphate (TCP; Ca 3 (P0 4)2), octacalcium phosphate (OCP; Ca 8 H2 (P04 )6 -5H 2 0) and hydroxyapatite ( H A ; Ca,o(P0 4 )6 (OH) 2 ) are commonly used for implants or implant coatings. The solubility of these calcium phosphates in a physiological condition is generally in following order from highest to lowest: D C P D > O C P > T C P > H A . The p H of the solution also affected the solubility of these calcium 17 phosphates. In general, as the p H increases from 3 to 9, the solubility of all four forms of calcium phosphates decreases [50, 51]. Because of the similarity of hydroxyapatite to bone mineral contents, hydroxyapatite has been studied as an implant material for many years. The crystal structure of hydroxyapatite is hexagonal with a=9.4 A and c=6.9 A cell dimensions [52]. A s a type of ceramic, hydroxyapatite is considered brittle. Its fracture toughness (KJC) is less than 0.9 MPa-m , and compressive and tensile strengths are less than 300 M P a and 50 M P a respectively. Based on its low mechanical properties, hydroxyapatite alone can not be the load bearing prosthetic material for bone. It is often used as a coating on metallic implants. 30 4 0 5 0 , 6 0 70 8-0 90 pH Fig . 2.5 Logarithm of the product of calcium and phosphate concentrations plotted against p H values of solutions saturated with respect to.various calcium phosphate phases in the ternary system Ca(OH )2 -H 3 P04-H20 Calculated for 37°C.(reprinted from ref 50 with permission from American Chemistry Society Publications). 18 Phosphate conjugates into several forms such as PO43", HPO4 2 " and H2PO4". In diluted aqueous solution, the existences of phosphate forms depend on the p H of solution. When the solution is strongly basic (high pH), the phosphate mainly exists as phosphate ions (PO4 "); when the solution is mildly basic, the phosphate is mostly hydrogen phosphate ions (HPCV-) ; when the solution is mildly acidic, the phosphate is comprised of dihydrogen phosphate ions (H2PO4"). This knowledge is essential for the deposition of calcium phosphates from an aqueous solution. 2.7 Deposition of hydroxyapatite There are many methods for depositing hydoxyapatite on metal surfaces. These methods include plasma spray, biomimetic deposition, sol-gel deposition and electrolytic deposition and so on. Electrolytic deposition (ELD) , which is also called electro chemical deposition (ECD) , is one of the non line-of-sight coating methods that are suitable for depositing hydroxyapatite on complex 3-D surfaces. Therefore, E L D is the ideal method for depositing hydroxyapatite on porous tantalum. Plasma spray deposition of hydroxyapatite[ 53 ] is the most popular method for orthopaedic implants. The hydroxyapatite powder is heated to a very high temperature (30,000K [54]) by plasma and sprayed on implant substrates. The coatings are lamellar structure [55]. Also , the melted hydroxyapatite may change to other forms of calcium phosphates such as tricalcium phosphate or even amorphous calcium phosphate [55]. There is hardly any chemical bonding between the substrate and hydroxyapatite coating. Thermal stresses are introduced to coatings and to the coating/substrate interface by the high 19 temperature used in heating. Plasma spray is a line-of-sight process which is not suitable for complex 3D structures such as porous tantalum because only the superficial pores can be coated. Biomimetic coating o f calcium phosphate [56] is a process that mimics the natural mineralization mechanisms in bone. During the coating process, substrates are simply immersed in a supersaturated calcification solution and the coatings grow onto substrates in matter of hours to days. Biomimetic coating of calcium phosphate is attractive because the obtained crystals bear more similarities to the composition of bone minerals, its morphology and structure, and the heat treatment of coating can be eliminated. Also , this method is non line-of-sight and is suitable for complex structures. Cu i et al [56] coated porous tantalum using the biomimetic method and obtained a flower-like calcium phosphate coating of calcium phosphate (see Fig.2.2). F ig . 2.6 S E M images of biomimetic coatings of CaP on porous tantalum (reprinted from ref. 56, with permission from Springer). The downside of the biomimetic method is the requirement o f a strict chemical environment and surface treatment. For the H A crystals to nucleate, the surfaces of prosthetic 20 metals need to be chemically etched (e.g. acid/alkaline treatment) or functionalized with negatively charged groups such as sulfonate [57], carboxyl and phosphate [58]. On the native T i surface without chemical modification, as crystallization theory predicts, the synthesizing parameters need to be strictly controlled [59]. Electrolytic deposition of calcium phosphate is another non line-of-sight process. It takes advantage of the electrode reactions in aqueous solution, and deposits the calcium phosphates on the cathode. When the electrical field is applied, the following cathode reactions w i l l happen [60]: 2 H + + 2e -> H 2 t 2 H 2 0 + 2e -> 2 0 H " + H 2 T 0 2 (aq) + 2 H 2 0 + 4e-+ 4 OFT These reactions w i l l raise the p H of the solution surrounding the cathode. A s mentioned above, the solubility o f calcium phosphates ( D C P D , O C P , T C P and H A ) drops while the p H of the solution increases (see page 17 and 18). B y controlling the p H of electrolyte and the concentrations of calcium and phosphate ions, different forms of calcium phosphates may deposit on the cathode where the prosthetic metal substrates locate. Our group [61] has developed an E L D coating technique that can achieve porous structural coatings of hydroxyapatite on titanium and tantalum. The porous structures of hydroxyapatite with pore size of 1 micrometer provide surfaces for bisphosphonate immobilization. The combination of the E L D coatings of hydroxyapatite and the 21 immobilization of bisphosphonates is one of the promising surface treatments for porous tantalum orthopaedic implants. Fig. 2.7 S E M image of the E L D coating of calcium phosphate. 2.7 Effect of bisphosphonates on bone Bisphosphonates are a family of drugs that have been clinically used to treat osteoporosis and to prevent bone fracture [62, 63]. The effects of bisphosphonates on bone are the decrease in the activities of osteoclasts and the increase of the activities o f osteoblasts. Both osteoclasts and osteoblasts, which resorb bone or deposit minerals on tissue respectively, are bone-related cells. The two effects of bisphosphonates mentioned above, are the enhancement of bone formation and an accelerated healing process after arthroplasty surgery [64,65] . Fig.2.8 is the molecular structure of alendronate which is one of the bisphosphonate families. 22 C H 2 — N H j C H 2 C H 2 H g O a P - C - P O g H j O H Fig. 2.8 Chemical structure of alendronate (reprinted from ref. 66, with permission from The American Society for Pharmacology and Experimental Therapeutics) Another chemical characteristic of bisphosphonates is their affinity with calcium ions. Bisphosphonates can chelate with calcium and form bisphosphonate and calcium bonds [59]. This characteristic enables bisphosphonates to be immobilized on hydoxyapatite and create an in vivo device that slowly releases bisphosphonates. 2.8 Current studies on bisphosphonate-delivery implants The effect of zoledronate, which is one type of bisphosphonates, on bone ingrowth into porous tantalum implants was examined in a canine model [67]. Zoledronate in saline was administered in a single post-operative intravenous injection at a dose of O.lmg/kg. At 6 weeks post-surgery, bone ingrowth increased by 85 % compared with controls. Subsequently, an investigation where zoledronate bound to hydroxyapatite coated (by plasma spray) porous tantalum was performed by the same group [68]. The study demonstrated an increase in peri-implant bone within the intramedullary (inside bone marrow) canal (32.2 % versus 13.8 % in controls) and intra-implant bone formation (19.8 % versus 12.5 %). However, the authors acknowledged the effect of the normal reparative stimulus to reaming the canal, resulting in greater peri-implant than intra-implant bone formation. Also , the plasma spray of 23 hydroxyapatite is a line-of-sight process in which only the superficial layer of porous tantalum was coated. Some researchers designed bisphosphonate-delivery methods on different implants materials targeting different implantation locations, and they reported positive effects on enhancing bone growth. Peter et al [69] soaked plasma spray hydroxyapatite coated titanium alloy in zoledronate for T H R , and Binderman et al [70] applied alendronate on alveolar surfaces before they implanted the dental replacements. 2.9 Gap model animal tests Traditional uncemented implants (porous coated and titanium fiber mesh) have shown good results clinically in the revision setting [71]. However in cases where less than 50% host bone contact is made the results of these uncemented implants have been less predictable [44]. Porous tantalum is a relatively new porous biomaterial, which may represent an improvement from conventional uncemented materials. Its properties of a high coefficient of friction (average 0.86±0.11 and 0.98±0.17 against surface-flattened cortical bone and cancellous bone respectively; average 0.58±0.06 when cancellous bone against surface-flattened cortical bone [72]), interconnected pores and high porosity by volume (80%) [31, 45] make it an ideal material in the revision setting with compromised host bone. The experimental animal tests should replicate the real case scenario during the arthroplastic surgeries. Orthopaedists cannot prepare a patient's pelvis to exactly match the artificial acetabular cup during arthroplastic surgery, especially in the case of bone 24 deficiencies caused by trauma, bone cancer and former damages revealed in revision. The gaps between bone and the acetabular cup are inevitable. A newly developed implant not only has to enhance the grown bone-implant attachment, but also has to overcome the gaps between host bone and implant. To simulate the real case, a gap model animal test is necessary. Studies [48,73] showed that gaps of 0.5 to 3 mm inhibited bone ingrowth, and the inhibition increased with the size of the gap. The gap model animal tests, should consider the bone size of the tested animal and the implant size to decide upon the proper gap size. 2.10 Mechanism of labeling The fluorochromes can chelate with the calcium atoms in hydroxyapatite and attach themselves to newly mineralizing front of bone [74]. Only the newly mineralizing surfaces can be labeled probably because of the difference of crystal sizes of hydroxyapatite between initial stages of mineralization and older mineralization sites [75]. The crystal size of hydroxyapatite is smaller at initial stages of mineralization than at older sites. The commonly used fluorochromes for bone labeling are calcein, alizarin complexone and tetracycline etc. [74]. On retrieved bone samples, these fluorochromes reveal different colors under a fluorescent microscope with the proper light sources and filters. A t different postoperative time-points, different fluorochromes can be administrated, and the mineralization process of newly formed bone is completely monitored [76]. However, 25 the injection dosages have to be carefully controlled because of the potential poisoning of these fluorescent chemicals. Pautke et al [77] developed a practical labeling protocol on mice bones. They adopted 8 different fluorochromes and sequentially injected them into mice bodies with a 3 day interval. Eventually, they labeled 7 time-points during 4 weeks of bone growth. They claimed that they could observe 6 different fluorescent bands by means of spectral analysis after 6 months, and 4 bands by the naked eye with proper fdters. Frosch et al [78] applied 4 different types of fluorochromes, which are xylenol, calcein, alizarin complexone and tetracycline, to study the enhancement of autologous osteoblast on ingrowth inside drilled 400-600 micrometer holes on titanium substrates. Specimen rabbits were injected with fluorochrome every 5 days postoperatively. The color labeling provided the 'wood pattern' rings of different post surgical time. 2.11. t-tests Statistics, as a mathematic tool, is often used for medical and clinical studies [79]. Researchers and medical doctors constantly need to draw conclusions on the effects of a pharmaceutical product or a medical procedure based on the symptoms of a small number of patients to those of the whole patient population. Individual patients differ from one another, and so it is essential to use proper methods in evaluating received data. This is the sample representation problem [79]. Another problem is the comparison among the data groups [79]. Based on the measured data, are there any differences among the data groups and are these 26 differences significant? t-test is one of the statistical methods comparing two groups of data at a time to answer these questions. There are several different types of t-tests, such as equal size or unequal size, dependent or independent, based on the two groups of data [80]. The t-test is equal or unequal size depending on whether the two compared groups have equal data points or not. The t-test is dependent or independent regarding whether there is any relation between two compared groups or not. If one sample is tested twice or two samples are matched or 'paired', the t-test is dependent; otherwise it is independent. Before t-test, we assume that the data points follow the 'normal distribution', and the means and the data distribution curves of two compared groups are not different (null hypothesis) [80]. The null hypothesis is either sustained or rejected based on the t-test result and the selective confidence level. The confidence level is the degree of certainty that a statistical prediction is accurate, and it is normally chosen to be 95 or 99%. p-value is "a measure of probability that a difference between groups during an experiment happened by chance.... The lower the p-value, the more likely it is that the difference between groups was caused by treatment."[81] For instance, i f the confidence level is chosen to be 95%, p-value, as the result oft-test, has to be bigger than 0.05 (1 minus the confidence level 0.95) to sustain the null hypothesis. Otherwise the null hypothesis is rejected, and the difference between two compared groups is caused by treatment. 27 t-test starts with the calculations of the arithmetic averages (X ) and the standard deviation (a) o f the data points of the two compared groups (X 1 and o\ for group 1; X 2 and a 2 for group 2). N represents the number of data in a group(Ni for groupl and N 2 for group 2). The degree of freedom is "a measure of the number of independent pieces of information on which the precision of a parameter estimate is based" (courtsey of Wikipedia). For the independent t-test with unequal sizes, the degree of freedom is the number of data minus 1. Therefore the degree of freedom in total is N i+N 2 -2 . The t-value can be calculated as following (assume the X x is bigger than X 2 )[82]: Once the t-value is calculated, it can be compared with the tabulated t-value at the degree of freedom and chosen confidence level (tabulated probability p equals 1 minus confidence level) from Student's t table (see Appendix A ) [82]. If the calculated t is bigger than the tabulated t, the difference of the two compared groups is significant. In this case, the calculated t can be compared with higher confidence level. Eventually, the p-value can be decided. If the calculate t is smaller than the tabulated t, p-value of t-test is considered bigger than the tabulated probability p, and the difference between the 2 compared groups is not significant. 28 CHAPTER 3 : SCOPE AND OBJECTIVES The scope of this research is to design a proper gap model of animal tests and to conduct one cycle of animal testing to study the effects of a newly developed alendronate-delivery-coating treatment on the osteoconductivity and bone ingrowth of porous tantalum implants. The experiments include implant processing such as electrolytic deposition of calcium phosphate and alendronate immobilization. Animal tests include surgeries and postoperative animal care, data collection and analysis include histology, histomorphometry and statistic analysis. The project requires collaboration with the Faculty of Pharmaceutical Sciences, the Department of Orthopaedics and the animal centre at the Vancouver General Hospital. The newly developed coating treatment takes advantage of the effects of alendronate on bone cells to reduce bone resorption and enhance bone formation, and combines with the electrolytic deposition calcium phosphate coating as a delivery vehicle. ELD coating of calcium phosphate is a non line-of-sight process suitable for complex 3 D structures such as porous tantalum, and the porous morphology of calcium phosphate coatings provides surfaces for alendronate adsorption. The final product is a combined device of orthopaedic implants with a local alendronate, delivery function, which is expected to improve the bone/implant fixation. The design of animal tests has to simulate the real clinical scenario and to separate the effects of alendronate from any other potential factors. The gap model design is employed 29 because the bone/implant gaps are inevitable during revision arthroplasty surgeries. The calcium phosphate coatings on porous tantalums are used as second controls. The fluorescence labeling method is employed to track the new bone formation and therefore also allows for studying the mechanism of bone formation. The objectives of this research are: 1. To modify an existing technique of electrolytic deposition of calcium phosphate to be suitable for porous tantalum substrates. 2. To immobilize an effective amount (dosage of 2 5 0 x l O " 9 M / L to 10"4 M / L per day) of alendronate onto the calcium-phosphate-coated porous tantalum implants. 3. To design a proper gap model of animal tests that can replicate the real case scenario of arthroplasty and revision surgeries. 4. To conduct one cycle of animal tests to evaluate the effects of alendronate delivery coating treatment on osteoconductivity and osteoinductivity of porous tantalum implants. 5. To study the bone formation mechanisms on different surface treatments of porous tantalum implants for the further understanding and improvement of implant surface treatments. 30 CHAPTER 4: MATERIALS AND METHODS Bone deficiencies, bone/implant gaps and defects are inevitable during revision arthoplastic surgeries [44]. The experimental designs of animal tests should replicate the clinical scenario. The intention of the gap model animal tests was to investigate the bone reaction towards different surface treatments of implant materials in the presence of a gap between the implant and bone. A s a result, the better implant surfaces in terms of overcoming gaps and enhancing bone/implant fixation could be determined. This animal study of alendronate-immobilized calcium-phosphate-coated porous tantalum implants included implant preparation, sterilization o f implants and tools, implant surgeries, post surgery care of animals, sample harvest and sample analysis. The implant preparation included pre-coating surface cleaning, electrolytic deposition of calcium phosphate, alendronate immobilization, cap mold design and bone cement cap manufacture, and cap assembling. The sample harvest and sample analysis step included euthanasia, femur dissecting and sectioning, bone histological processing and embedding, and histology and histomorphometry analysis. 4.1 Implant preparation 4.1.1 Porous tantalum The porous tantalum implant (Trabecular Metal ™) cylinders, which were 3.18 mm in diameter and 8 mm in length, were provided by Zimmer Inc. (Warsaw, Indiana, U S A ) . The pore sizes of these tantalum plugs were about 400 to 500 micrometers (shown in Fig. 31 4.1), and the porosity o f these porous tantalum was about 75%. The density of solid tantalum was 16.65 g/cm 3 . Weighted with a 1/100,000 g balance (Mettler AE240) , the weights o f five porous tantalum implants were listed in Table 4.1, and the average weight was 0.237 (SD 0.014) g. Fig. 4.1 S E M images of porous tantalum cylinder at various magnifications. Table 4.1 The weight of porous tantalum implant plugs No. 1 2 3 4 5 average Weight (g) 0.225 0.237 0.261 0.228 0.235 0.237 (SD 0.014) The surface area is a parameter, which affects current density, for porous tantalum plugs as substrates for electrolytic coating of calcium phosphate. The estimation of surface 32 area was based on the assumption that the porous tantalum plug was built with a single tantalum tube wrapped in cylindrical space of the 3.18mm in diameter and 8mm in length. The dimensions of this tube are 0.15 mm for outer diameter and 0.03 mm for inner diameter, as measured on B S E images. Based on the average weight and the density of tantalum we could calculate the volume of solid tantalum, which was then converted to the tube surface area. The surface area of an implant plug thus estimated as 395 (SD 23.7) mm . 4.1.2 Electrolytic deposition (ELD) of calcium phosphate coatings 4.1.2.1 Surface cleaning of tantalum specimens Surface contamination such as grease and non-uniform oxidation of metal substrates etc could affect coating kinetics, uniformity and morphology. Specimens needed to be cleaned properly before coating. The porous tantalum cylinders were ultrasonically bathed in acetone for 3 minutes and immersed in high concentration sulfuric acid (95 wt% H2SO4) (Fisher Science Inc) for 1 minute. The high concentration sulfuric acid can dehydrate and destroy the organic molecules. Both steps were designed to eliminate any organic contamination. The porous tantalum cylinders were then rinsed and cleaned in ultrasonic bath of distilled water for 3 minutes, and then rinsed 5 times in distilled water to remove sulfuric acid (20ml or more water for more than 5 min each time). The residue of sulfuric acid from surface cleaning process should not effect bone cell reaction to implants. In animal or human bodies, there are about 0.3 m M / L sulphate in their serum [83], and these sulphate anions come from food, water and air. After rinsing, the 33 residue o f sulphate anions from implants should be much less than the sulphate anions in blood stream. "Excess sulphate in the blood is rapidly eliminated by urinary excretion" [84]. 4.1.2.2 ELD coating cell fixture A customized electrolytic cell fixture for electrolytic deposition was designed and manufactured in biomaterial laboratory (shown in Fig 4.2). The fixed interval between 2 electrodes, which were cathode for implant plugs and anode for 25><25mm platinum plate respectively, is 9 mm. It was measured from the surface o f the platinum plate to the closer parallel tangent surface of the porous tantalum cylinder. The anode clip was made o f stainless steel which was not immersed in the coating electrolyte, and the cathode rod which fully contacted with the porous tantalum cylinder was made o f tantalum. The porous tantalum cylinders were completely immersed in the electrolyte during E L D deposition. Fig . 4 .2 The image and schematic illustration o f E L D coating cell 34 4.1.2.3 ELD coating parameters In order to study calcium phosphate coating parameters, the crystal structure of the coating, and the morphology and thickness of calcium phosphate coatings, 100-micrometer thick tantalum foils of 20x20mm and porous tantalum blocks of 3x3x8mm were used. Only half of the square tantalum foil was immersed in electrolytic solution, therefore the surface area of this electrode was 400mm 2 which was close to the surface area of porous tantalum cylinders. The dimensions of porous tantalum blocks were chosen to have a closer geometric shape to the porous tantalum cylinders. The modification of coating parameters was conducted based on the E L D coating technique developed by our group [61]. The parameters were selected to avoid the formation of hydrogen bubbles which harms the adhesion of coatings, and the formation of D C P D s which affects the uniformity of coating structure and morphology. Other criteria, such as the cracking o f coatings, the thickness of coatings and the uniformity of thicknesses of coatings at both inner and outer pores, which is particularly important for porous tantalum, are evaluated as well . The goal is to achieve coatings that are crack-free and have homogeneous micro-porous structure with a uniform thickness of 3 to 5 micrometers of calcium phosphate throughout the porous tantalum cylinder. Some of the coating parameters, such as the concentrations of calcium and phosphate ions in the coating electrolyte, p H of the solution and the applied voltage, and so on, have already been optimized [61] and were used without being changed in this study. Periodical stirring was introduced to increase the thickness uniformity across at central pores. The effect 35 of periodical stirring is examined by comparing coatings with and without this process. Coating time was chosen to achieve the maximum thickness of coatings with minimum cracking. The modified processes of E L D of calcium phosphate are as follows. The E L D of calcium phosphate was carried out in an aqueous solution that contained calcium and phosphate ions which was prepared by adding proper amount of calcium nitrate (Ca(N03)2-4H20), ammonium dihydrogen phosphate (NH4H2PO4) and sodium chloride (NaCl) (Fisher Scientific). The concentrations of calcium and phosphate ions were 5.25 m M and 10.5 m M respectively, and the concentration of Sodium chloride was 150 m M . The p H of the solution was then adjusted to 5.25±.07 by adding I N sodium hydroxide (NaOH) under the monitoring of a calibrated p H meter (Thermo Orion, Beverly, M A . U S A ) . The applied D C voltage was 2.6 V , and the duration of coating was 3 hrs. During the coating period, the coating solution was periodically stirred every 30 minute which should enhance the ion supply to the centre pores of implant plugs. 4.1.3 Alendronate immobilization Bisphosphonates contain two phosphonate groups attached to a single carbon atom (C), forming a P-C-P structure. Alendronate (monosodium 4-amino-l-hydroxybutylidene-1,1-diphosphonate trihydrate) was the bisphosphonate used in the study. It was chosen because of its efficacy, availability, and clinical record. The immobilization of alendronate (Sigma Aldrich) was carried out by immersing calcium phosphate coated implants in 3.2 ml 10"4 M sodium alendronate in 0.01 M Phosphate 36 Buffer Saline (PBS) solutions at 37°C for 7 days. Proportion of besphosphate ions from solution were adsorbed by the calcium ions on the Ca-P coating surfaces. Then, the soaked implants were rinsed with distilled water five times (20ml or more water for more than 20 min each time). The rinsing steps ensured that most of the non-adsorbed alendronate ions were removed from implants. 4.1.4 Implant assemblies The gap model was achieved by installing polymethymethacrylate ( P M M A ) bone cement (Simplex ® P, Stryker Howmedica Osteonics, Limerick, Ireland) caps at both ends of the implant plugs. The caps act as spacers between bone and the porous tantalum plugs. A s a result, the desired gap between the implant and the surrounding trabecular bone was maintained when the caps were press-fitted to the pre-drilled holes in the femoral bone. The final gap is a 0.6mm thick cylindrical wall with the inner diameter of 3.18mm and an outer diameter of 4.37mm. The volume of the gap zone is about 35 m m 3 . The details of the dimensions of the final implant assemblies are illustrated in Fig. 4.3. Bone cement was used as the material for caps because it is clinically used in orthopaedic surgeries. Also , the flexible mixing ratio and adjustable curing time make it an ideal casting material. 37 Bone cement Spacing Cup Porous Tantalum Implant Cyfinder Designed Gap 04.37mm - f — 03.18mm|-9mm Fig . 4.3 Mechanical drawings o f porous tantalum implant assembly. The left is a 3D drawing of the implant assembly, and the right is a 2-D drawing with al l the dimensions Four customer-designed molds (Fig.4.4) were machined and used to cast bone cement caps. The ratio for bone cement was 0.7 grams of powder per 0.5ml o f monomer liquid, which allows 5 min of viscous state for handling. The dimensions of the cap were 4.37mm in outer diameter, 3.18 mm in inner diameter, 2mm in height and 1.5mm in depth. During assembly, the implant plugs were carefully handled to avoid any damage to the coating surfaces, and a drop of bone cement (same ratio as above) was added into each cap as a glue to firmly attach the 2 caps to each end o f the implant cylinder (Fig. 4.3). Fig . 4.4 The image o f bone cement casting mold 38 4.1.5 Sterilization of implants and surgical tools To minimize any damage to the rabbit's femoral bone, such as cracking during surgical drilling, a set of twist drill bits of serial diametric sizes was employed. They were 5/64" (1.98mm) brad tip, 1/8" (3.18mm) and 11/64" (4.37mm). These drill bits (Lee Valley Co. Canada) were ultrasonically cleaned in acetone and ethanol to remove organic contaminations. Then they were immersed in 70% isopropyl for pre-sterilization before being sealed in polyethylene bags (Food Saver). The implant assemblies were re-sealed in polymer packaging bags from Zimmer (see Fig. 4.5). Fig. 4.5 Surgical tools and porous tantalum implant Beta-ray radiation sterilization was carried out at Iotron Tech. Inc., Canada following the ISO13485:2003 international standard for medical devices. The applied radiation dosage was 25 to 27 KGys. Other tools such as the hand drill and surgical forceps were sterilized by autoclave at Jack Bell centre, Vancouver General Hospital. 39 4.2 High performance liquid chromatography (HPLC) tests for drug loading Five alendronate-immobilized calcium-phosphate-coated porous tantalum implants were tested for total amounts of drug loading. They were processed in the same way as the implants for animal tests. These 5 pilot samples were immersed in 1ml 1M hydrochloric acid, and the alendronate-containing calcium salt on the samples was dissolved in the acid solution. The tantalum samples were examined with the Nikon Eclipse optical microscope and the coating was confirmed to be completely dissolved. Seven alendronate solutions with known concentrations, which were 100, 50, 25, 12.5, 6.3, 3.1 and 1.6 ug /ml respectively, were also prepared as standard samples. These samples generated a linear calibration equation between alendronate concentration and a peak area of alendronate (IT = 0.9998). During HPLC (high performance liquid chromatography) analysis, solutions reacted with 5 fold (V:V) of 2 mg/ml fluorescamine in acetonitrile to form a conjugate. The acetronitrile was extracted off with dichlormethane. The aqueous phase containing the conjugate was analyzed with a Waters chromatography station, with Nova Pac CI8 column and 470 scanning fluorescence detector. The conjugate was detected by excitation at 395 nm and emission at 480 nm. 4.3 Animal tests Alendronate-immobilized and calcium-phosphate-coated porous tantalum implants were tested in comparison with uncoated porous tantalum as control samples. To confirm the 40 effects of alendronate on bone formation, calcium-phosphate-coated porous tantalum was adopted as another set of control samples as well . The detailed plan is shown in Table 4.2. Eighteen New Zealand White female rabbits of 3.5-5.0 kg (28 to 34 weeks old) were employed for the animal tests. They were randomly divided into 3 groups for the 3 types of implant assemblies which are non-treated porous tantalum (Ta) from Zimmer as control samples (6 rabbits), calcium-phosphate-coated porous tantalum (Ta-CaP) as second type of control samples (6rabbits), and alendronate-immobilized calcium-phosphate-coated porous tantalum (Ta-CaP-ALN) (6 rabbits). The animal tests were conducted at the Jack Bel l centre (animal centre), Vancouver General Hospital. The study was approved by the Animal Research Ethics Review Board at the University of British Columbia (Protocol number: A04-0275). Table 4.2 Animal test plan for bisphosphonate-coated porous Ta bone implants Ta Ta-CaP Ta-CaP-ALN (Porous Ta control) (Porous Ta with calcium Phosphate coating) (Ta-CaP soaked in alendronate) Rabbits 6 6 6 Number of implants 12 12 12 Fluorescence labeled samples 12 12 12 Histological sections 36 36 36 S E M / fluorescene 36 36 36 microscopy Image analysis 36 36 36 41 4.3.1 Implantation procedures The implantation location of the distal femur was selected because the trabecular bone inside the femoral distal end is similar to the bone at the pelvis. The rabbit was shaved and then anaesthetized by intramuscular injection of a sleeping dose of 3.15 to 3.5 mg ketamine/xylazine. The anaesthesia was maintained by inhalation of 2% to 2.5% isoflurane. Then, pain killer (0.3mg Buprenorphine, Temgesic) was administrated subcutaneously. The skin of the rabbit, where the incision would take place, was cleaned with 0.5% chlorhexadine and then sprayed with 0.5% dexidin. These pre-surgery procedures were conducted by the animal care staff at the Jack Bel l Center. The surgical procedure on the rabbits was performed by Dr. Winston K i m (a surgeon) in the animal operating room. A 3 cm incision was made on the distal lateral aspect of the femur, and the vastus lateralis was split along its fibers to expose the underlying bone. One 4.37 mm hole was made bilaterally through sequential drilling (1.95 mm, 3.18 mm and 4.37 mm) under saline irrigation. The orientation of the hole was perpendicular to the distal femoral condyle, and the depth was confirmed with a depth gauge. The porous implant with P M M A bone cement caps was gently inserted into the hole (see F ig 4.6). The wound was then irrigated and closed in a uniform manner. Two implant plugs with the same surface were implanted in each rabbit, to avoid the confounding influence of the presence or absence of different implant coatings. The anaesthetic, operative procedure and animal care were performed in compliance with University and federal guidelines. 42 *orous Tantalum Condy! Cement Cap / G a p Trabecular Bone Fig. 4.6 The images and schematic illustration of the implant location, a) digital image of implant position during surgery; b) X-ray image shows the implant position after surgery; c) the schematic illustration 4.3.2 Bone labeling Bone growth front was labeled with fluorescence dyes to track new bone formation with time. Two commonly used fluorochromes from Sigma-Aldrich, Alizarin complexone (absorption: 530-560 nm, emission: 624-645 nm) and calcein (absorption: 494 nm, emission: 517 nm), are dissolved in distilled water, and sodium bicarbonate was added to reach the concentration of 1.4% wt. The pH of the solutions was adjusted by adding hydrochloric acid if necessary. Within a Type II A bio-safety cabinet, these solutions were filtration sterilized with 0.2 micron filters (Millex GS, Carrigtwohill, Co. Cork, Ireland) and packaged in 10ml sealed syringes. The detail information about these two fluorochromes is listed in Table 4.3. Table 4.3 Information of fluorochomes and their solutions Al izar in compleone Calcein Sigma Catalog No . A3882 C0875 Absorption 530-560 nm 494 nm Emission 624-645 nm 517 nm Solution Concentration 30mg/ml 1 Omg/ml Solution p H 7.4 7.4 The labeling chemicals were administrated subcutaneously. The labeling should provide information about the direction of bone growth, growth speed and bone growth mechanism. Calcein and Al izar in complexone were chosen because their emission wave lengths are within visual light range which means no special light sources such as ultra violet are needed. Also , these 2 chemicals are relatively mild to rabbits. A s shown in Table 4.4, the dosages of calcein and alizarin were 10 and 30milligrams per kilogram of rabbit's body weight respectively. Table 4.4 Labeling administration dosages and schedule Dosage concentration Solution volume Day after surgery Calcein lOmg / K g 1 Omg/ml 1 m l / K g 7 t h day Calcein lOmg/Kg 1 Omg/ml 1 m l / K g 14 t h day Al izar in 30mg/Kg 3 Omg/ml 1 m l / K g 21 s t day Al izar in 30mg/Kg 3 Omg/ml 1 m l / K g 28 t h day . 4.3.3 Animal care after surgery Antibiotic medicine, Notrile, was administrated after the surgeries. Rabbits were fed with carrots, hay and vitamins, and the weights and body temperature of these rabbit patients were monitored regularly. Force feeding may be applied to those 'depressed' rabbits. Further administration of antibiotics may be applied to those rabbits with higher body temperature. 44 4.3.4 Euthanasia, radiograph imaging and sample harvest After 29 days of healing process, the rabbits were intravenously injected with Pentobarbital 2mg/Kg of body weight to stop the heart beat in less than 2 seconds. After the animal deceased, X-ray images were taken to reveal the implant positions. Afterward, the lateral legs were dissected, and then the femoral bones with the implants were removed for further histology analysis. 4.4 Sample analysis 4.4.1 Histology sample preparation The femur was cut parallel to the radius direction of the condyle arc (see Fig. 4.7). The harvested samples were fixed, dehydrated, infiltrated and then embedded in epoxy resin ( S P U R R , Canemco, St. Laurent, Q C , Canada) following a standard histological procedure (see Table 4.5) [85]. Table 4.5 Processes and schedules for fixation, dehydration and infiltration Processes Chemicals Temperature Duration Fixation 10% buffered formalin 20°C 2days Dehydration 70% Ethanol/Water 20°C 3 days Dehydration 90% Ethanol/Water 20°C 3 days Dehydration 100% Ethanol 20°C 3 days Dehydration 100% Ethanol 20°C 3 days Defat 100% Acetone 20°C 3 days Infiltration 50% SPURR/Acetone 20°C 3 days Infiltration 80% SPURR/Acetone 4°C 3 days Infiltration 100% S P U R R 4°C 3 days Infiltration 100% S P U R R 4°C 3 days 45 Fig. 4.7 Schematic image for sectioning and observation directions In order to examine bone formation at the site o f the implants and to reveal the 3-D structure, each embedded sample was longitudinally cut with a diamond saw (Buehler, Lake Bluff, IL U S A ) , and ground and polished (Buehler) into three sections at 200 um, 850 um and 1500 urn deep from the tangent surface o f the implant. Fig.4.8 gives the schematic illustration of the three examining sections. Section 3 Section 1 Section 2 Fig . 4.8 Schematic illustration for three examining sections 46 After samples were embedded in S P U R R (Canemco, St. Laurent, QC, Canada), the examining surfaces were ground to just touch the implant cylinders, which meant the grinding surfaces were the tangent surfaces of the porous Ta implant cylinders. The surfaces were considered the origin of the coordinate. Then the 2 samples from same rabbit were re-embedded together in Epothin (Buehler) resin in a standard (1.25 inch in diameter) mold. After curing, the bottoms of the Epothin blocks were ground parallel to the sample surface with the grinding fixture at an accuracy of ±20pm. The final surfaces were vibration polished with 0.05 micrometer silica slurry. 4.4.2 Histology and histomorphometry analysis 4.4.2.1 Fluorescence microscopy A fluorescence microscope (Eclipse E 600, Nikon, Tokyo, Japan) with a F ITC & T R I T C dual exciter and dual emitter filter block, (exciter wavelengths: 475 - 495 nm and 540 - 575 nm, emitter wavelengths 500 - 535 nm and 580 - 620 nm), was used to analyze the fluorescence labels and new bone formation. The fluorescent imaging was performed before sputtering of Au /Pd alloy for S E M . Under the fluorescence microscope, bone formation front in the first 2 weeks was marked by two green calcein labels, while red alizarin Complexone lines marked bone formation in the 3rd and 4th week. 4.4.2.2 BSE imaging The back scattered electron (BSE) images were taken under a Hitachi-S3000N S E M (Hitachi Scientific Instruments, Tokyo, Japan). Polished samples were sputtered with Au/Pd. B S E mode under high vacuum pressure was used. The beam current was set to 65mA, 47 voltage was 20 k V , and the aperture was removed. A l l these setups guarantee bone, epoxy and Ta to have differentiable gray scales which meant the histogram would show 3 peaks. 4.4.2.3 Image analysis The image analysis was carried out on the low magnification B S E images using image analysis software (Clemex Vis ion P E 3.5, Longueuil, Q C , Canada). The magnification calibration of B S E images with image analysis software was 3.93 pm per pixel. The quadrilateral defined by the 4 corners of caps was considered the sample area. Gray level discrimination was used to identify epoxy, new bone formation and implant. Figure 4.9 shows how Clemex processes the images. Analysis on the 3 sections of B S E micrographs generated quantitative information on total gap volume, total volume available for bone ingrowths, and total volume available for new bone formation. The volume of bone gap filling, which was new bone formation in the gap region between the tantalum implant and host bone, and bone ingrowth into porous tantalum which was bone formation within the pores of the tantalum plugs, were measured. The total bone formation (the sum of the gap-filling and ingrowth) was then calculated. Total available tantalum length (see Fig . 4.10 (b): black lines) and the contact length (see Fig. 4.10 (a): blue lines) of newly formed bone on the porous tantalum surface of the three longitudinal sections of all implants were also measured. F ig 4.11 gave the schematic illustration of the defined gap-filling and bone ingrowths in this study. 48 < 3 i 0 CQ s I 128 192 255 Fig. 4 .9 Histogram for image analysis and the images before and after the dying process 49 Fig . 4.11 The schematic drawings illustrate the bone ingrowth and gap fil l ing 4.4.2.4 Statistical data analysis Differences in the mean and standard deviation were investigated by a two-tailed t-test, using a statistical software program (SPSS Version 14.0, Chicago, IL, U S A ) . A n independent sample t-test was used to compare the volume of gap fil l ing, bone ingrowth, total bone formation, and the new bone formation contact length of the different implants. We chose 95% for the confidence level and considered a p-value o f < 0.05 to be significant. 50 CHAPTER 5: RESULTS 5.1 The effects of coating parameters on calcium phosphate coating morphology Many coating parameters such as p H and the temperature of the electrolytes, applied voltage, calcium and phosphate ion concentrations and coating duration affect the final coating crystal structure and coating morphology. Duan [61] has studied these parameters on flat titanium substrates and has optimized the coatings to a micro-porous structure good for drug delivery. To apply this E L D coating technique to porous tantalum substrates for the animal test, the coating process needs to be standardized to achieve repeatable coatings on all the alendronate-immobilized implants and the calcium-phosphate-coated controls. Some coating parameters need to be modified to fit the porous substrates of tantalum. The porous structure of calcium phosphate coatings can be obtained on porous tantalum substrates by using the same concentrations of calcium and phosphate ions, solution p H and applied voltage as those found on titanium plates. The remaining two major concerns on calcium phosphate coatings were cracking and spallation of coatings and the uneven coating thicknesses at superficial pores and central pores. Porous tantalum plugs were coated with calcium phosphate without sulfuric acid cleaning. The calcium phosphate did not uniformly cover the tantalum struts. On some samples, uncoated spots surrounded by calcium phosphate coating could be seen (see Fig. 5.1 (a)). Wi th the sulfuric acid cleaning, these phenomena dramatically reduced. 51 S p a l l a t i o n a n d c r a c k i n g o f the p o r o u s c a l c i u m p h o s p h a t e c o a t i n g are r e l a t e d to the t h i c k n e s s o f the c o a t i n g i f o t h e r f a c t o r s , s u c h as c o a t i n g p o r o s i t y , r e m a i n c o n s t a n t . C o a t i n g d u r a t i o n w a s c h o s e n to c o n t r o l the c o a t i n g t h i c k n e s s i n o u r e x p e r i m e n t s . O b s e r v a t i o n s u s i n g a n o p t i c a l m i c r o s c o p e w e r e a d o p t e d to e x a m i n e c o a t i n g s u r f a c e s . T h e d u r a t i o n o f the E L D c o a t i n g p r o c e s s w a s d e t e r m i n e d ; h e n c e f e w c r a c k s o r c r a c k - f r e e c o a t i n g s w e r e a c h i e v e d . F i g . 5.1 (b) s h o w s s p a l l a t i o n h a p p e n i n g o n a p o r o u s t a n t a l u m s a m p l e . T h e c o a t i n g d u r a t i o n f o r p o r o u s t a n t a l u m i m p l a n t s w a s s h o r t e r t h a n that for th i s s a m p l e . F i g . 5.1 S E M i m a g e s o f c a l c i u m p h o s p h a t e c o a t i n g s , a) w i t h o u t s u l f u r i c a c i d c l e a n i n g , the p o r o u s t a n t a l u m strut w a s n o t c o a t e d ; b ) s p a l l a t i o n o f c a l c i u m p h o s p h a t e c o a t i n g s ; c ) c o a t i n g s m o r p h o l o g i e s at c e n t r a l p o r e s w i t h o u t p e r i o d i c a l s t i r r i n g ; (d ) c o a t i n g m o r p h o l o g y at s u p e r f i c i a l l a y e r w h e n p e r i o d i c a l s t i r r i n g w a s not a p p l i e d . 52 Several porous tantalum plugs were coated with calcium phosphate without periodical stirring. Fig. 5.1 (c) and (d) show coatings at the central pores and at the superficial pores on one sample respectively. At the central pores (Fig. 5.1 (c)) the coating was very thin (less than 500 nm), and the sharp edges o f the metal grains could be seen. Also , the pore size of calcium phosphate coatings was very small. A t the surface (Fig. 5.1 (d)) of the porous tantalum cylinder, larger pores of calcium phosphate coatings were formed, and the metal grain edges were blunt or buried under calcium phosphate coating. The estimated thickness was more than 3 micrometers. The causes o f these uneven coating thicknesses in Figure 5.1 might be the insufficient supplies of calcium and phosphate ions to the central pores because all ion supplies defusing from the surrounding solution were consumed by the superficial layer. Periodical stirring was introduced to the coating process to manually deliver the calcium and phosphate ions into the central pores of porous tantalum as the ion supplies for coating. Fig. 5.2 shows the improved coating at a central pore with the periodical stirring. Fig. 5.2 S E M images on calcium phosphate coatings at a central pore when periodical stirring was applied. 53 5.2 Calcium phosphate coating and alendronate loading 5.2.1 ELD coating of calcium phosphate The digital image and S E M images under different magnifications (Figure 5.3) show the morphology of calcium phosphate coatings on porous tantalum implants. The digital image (Fig. 5.3(a)) shows the complete implant assembly. The low magnification image (Fig. 5.3(b)) shows that the coverage of the calcium phosphate coating over the complex 3D tantalum structure is almost 100%. Higher magnification images (Fig. 5.3(c), (d)) illustrate the coating morphology which is a thin layer of calcium phosphate with a micro-scale porous structure (about 1 um in pore size). A tantalum foil sample was employed to conduct a coating thickness measurement. The coating parameters for this foil substrate were the same as those for porous tantalum cylinders. After coating, the foil was bent 180° to crack the calcium phosphate coatings. The coating thickness was 5.8 micrometers as measured on an S E M image (Fig. 5.4). 54 Fig. 5.3 S E M images at varying magnifications of Ca-P coating morphology, a) digital image of implant assembly; b) low magnification S E M image shows the coverage of CaP coating; c) and d) the porous morphology of CaP coating. 5.2.2 The alendronate loading amount and estimation of dosage The H P L C experimental results of alendronate loading amount on the 5 pilot porous tantalum plugs were calculated, and the drug dosages were estimated accordingly. The total drug amount loaded by immobilization was at microgram level, and the average drug amount on these 5 porous tantalum plugs was 1.37 (SD 1.0) micrograms. The molecular weight of alendronate was 325.12 grams per mol. The volume of gap was calculated as -35 mm , and the volume of pores was calculated as -48 mm 3 , based on the dimensions and measured porosity of porous tantalum cylinders. The total volume of sample zone was -83 mm . Assuming that all drugs dissolved in one day, which was the extreme case that was unlikely to happen, the highest possible drug dosage was -4.7 x l 0 " 5 M / L . This dosage was still in the effective zone of alendronate concentrations which were from 250x10" 9 to 10"4 M / L per day according to Bergstrom[86]. In our animal tests, there should not be any concern of overdose of alendronate. If all the drugs and calcium phosphate slowly dissolved in 20 days, the average daily dosage should be ~2><10"6 M / L which was within the effective concentration zone. 5.3 Crystal structure of ELD calcium phosphate coating X R D results showed the E L D coatings to be a mixture of H A and O C P (Fig. 5.5), as evidenced by the diffraction (shoulder) of O C P at 29 = 4.72° and diffraction of H A at -50 (49.46 (2 1 3) plane). The standard H A or O C P has higher intensity at 33 degrees than that at 26 degrees. The X R D spectra of E L D coatings (and alendronate-immobilized coatings) have higher intensity at 26 degrees ((002) planes), indicating a preferred orientation of the calcium 56 phosphate crystals in E L D coatings. Comparing the two X R D spectra, it can be concluded that alendronate immobilization did not change the crystal structure of the calcium phosphate coating. cz CD HA/OCP (002) Af te r B e f o r e 10 20 30 2e (degree) 40 50 Fig . 5.5 X R D results for CaP coatings before and after alendronate immobilization F T I R spectra of the coatings before and after alendronate immobilization are shown in Fig.5.6. The PO4 3" peaks at 561 and 600 cm"1 are close to those of octacalcium phosphate (OCP), which is one type of calcium phosphate crystal structure with PO4 3" peaks at 561 and 599 cm" [87 ]. A s a comparison, the PO4 peaks of hydroxyapatite (HA) around 600 cm" are 566.7 and 603 cm"1 [88], and the interval between these two peaks is smaller than the current result. After alendronate immobilization, FTIR spectrum shows slight changes at 1080 and -875 cm" 1. FTIR spectrum of alendronate calcium salt shows the P - 0 stretching modes of phosphonate at 1090, 1000 and 964 cm"1 [89] . The changes in F T I R spectrum after 57 immobilization might be because of the chelation of alendronate and calcium ions on the calcium phosphate coating surfaces. The energy dispersive X-ray spectrometry (EDS) tests on E L D calcium phosphate coatings were carried out on tantalum foil substrates to compare the coatings of calcium phosphate before and after alendronate immobilization. The atomic ratio of calcium and phosphorus can be used as supporting evidence of the crystal structure because different calcium phosphate crystals have different Ca/P ratio. On each type of coating, five different areas located at the centre and the 4 corners were tested, and the averages of CaP atomic ratios were calculated to be 1.54 and 1.53 before and after alendronate immobilization respectively. The adsorption of alendronate onto the coating thus did not significantly change the Ca/P ratio of the calcium phosphate coating. 1400 1200 1000 800 600 400 Wavenumber ( c m 1 ) Fig . 5.6 The FTIR results for CaP coatings before and after alendronate immobilization 58 In conclusion, the crystal structure of the calcium phosphate coatings was not changed during alendronate immobilization based on the results of X R D , FTIR and E D S . The coatings before and after alendronate immobilization are predominantly O C P . However, the existence of H A can not be excluded. The crystals in the E L D coating are oriented. 5.4 Histology The microscopic analysis of peri-implant bone growth was carried out with a fluorescence optical microscope and a scanning electron microscope. The bone growth patterns, including bone formation, growth directions and the bone/implant interlocking etc, were studied. The comparisons of these histological results between control groups and the drug-loaded group were conducted. Figure 5.7 shows the B S E images (left column) and fluorescence optical images (right column) of bone growth on the 3 types of implants: Ta control (Fig.5.7 (a), (b)), Ta-CaP control (Fig.5.7 (c), (d)) and T a - C a P - A L N implant (Fig.5.7 (e), (f)). In the B S E images, it could be obviously seen that peri-implant bone growth and ingrowth into the implant were the highest on the T a - C a P - A L N sample (Fig.5.7 (e)). Between the 2 types of control samples, the Ta-CaP control had more bone growth in defined sample area than the Ta control did. 59 c 2 mm d 4 2 m m f / • ft * 2 mm Fig . 5.7 B S E and fluorescence images of bone formations on 3 types of implants, a) and b) non-coated control; c) and d) calcium phosphate coated control; e) and f) alendronate-immobilized calcium-phosphate-coated implant. The fluorescence optical images in Figure 5.7 show the whole bone sections where the sectioning planes were made along the axis o f implant cylinders. The shining green colour was the emission light of calcein under the excitation light of 495 nm wavelength. 60 Again, we could observe significant increases of bone growth on T a - C a P - A L N implant comparing with those on the 2 types of controls. Also;" the pore size of the newly grownperi-implant bone was smaller in comparison with two other control samples. Furthermore, the t green colour indicated that these new bones were formed in the first 2 weeks post-surgery. The depth of ingrowth into porous tantalum implants on T a - C a P - A L N sample was higher than those on the two types of controls in which ingrown bone only reached the superficial layer of pores. 5.5 Histomorphometry The ingrowth was the new bone that grew into porous tantalum plugs; the gap filling was the new bone growth in the designed gap area, and the total bone growth was the sum of the ingrowth and the gap filling. The information from 3 sections was integrated into 3-D volume percentages. For the bone/implant contacts, the data were presented separately. 5.5.1 Bone formation 5.5.1.1 Gap filling The mean extent of gap fillings was 13.0 % (SD 5.0 %) for Ta controls, 15.8 % (SD 3.3 %) for T a - C a P - A L N controls, and 29.1 % (SD 10.0 %) for T a - C a P - A L N implants (shown in F ig 5.8). The relative increases in volume of gap fillings on T a - C a P - A L N implants were 124% (2.24 fold) and 84% (1.84 fold) compared with Ta controls and Ta-CaP controls respectively. 61 The t-tests, comparing T a - C a P - A L N implants with Ta controls, indicated the difference of gap fillings between these two groups was significant (p-value was less than 0.0001) (see Table 5.2). Between T a - C a P - A L N implants and Ta-CaP controls, the difference of gap fillings was significant as well (p-value was less than 0.0005). On the other hand, the calcium phosphate coating slightly increased the bone formation volume of gap filling, but the p-value of t-test was 0.12 which meant that the difference of gap fillings between these two groups were not significant. 5.5.1.2 Bone ingrowth The mean volume of bone ingrowth was 6.0 % (SD 2.7 %) for Ta controls; 8.3 % (SD 2.9 %) for Ta-CaP controls, and 19.9 % (SD 5.9 %) for the T a - C a P - A L N (see Fig.5.8). The increases in bone ingrowths on T a - C a P - A L N samples were 232% (3.32 fold) and 140% (2.40 fold) compared with Ta controls and Ta-CaP samples respectively. The p-value of the t-test in comparison of T a - C a P - A L N implants with Ta controls resulted in less than 0.0001 which meant the difference of ingrowths between these two groups was significant (see Table 5.1). The difference of ingrowths was also significant between T a - C a P - A L N implants and Ta-CaP controls (p-value was less than 0.0001). In comparison to Ta controls, bone ingrowth was slightly increased on Ta-CaP samples, but the p-value of the t-test was 0.07 which meant the increase of bone ingrowth on Ta-CaP samples was not significant. 62 5 .5 .1 .3 Total bone formation The volume of total bone formation, which was the sum of gap filling and bone ingrowth, was 9.7 % (SD 3.5 %) for Ta controls, 12.4 % (SD 3.2 %) for Ta-CaP controls and 26.2 % (8.3 %) for T a - C a P - A L N implants. The increase of total bone formation on Ta-CaP-A L N implants was 170 % (2.7 fold) compared with that on Ta controls. The increase was significant because the p-value of t-test was less than 0.0001. The total bone formation on T a - C a P - A L N implants showed a 111 % (2.11 fold) increase in comparison to Ta-CaP samples, and it was also statistically significant (p < 0.0001). 6 0 5 0 -4 0 -| 3 0 o > 2 0 1 0 0 ** * — Gap Filling Ta W& Ta-CaP I I Ta-CaP-ALN ** * ** * Bone Ingrowth Total Bone Formation Fig. 5.8 The mean and standard deviation of volume of gap filling, bone ingrowth and total bone formation. T a - C a P - A L N implants demonstrated significantly increased gap filling, bone ingrowth and total bone formation compared with Ta-CaP controls (p < 0.0005, p < 0.0001 and p < 0.0001 respectively *), and significantly increased gap filling, bone ingrowth and total bone formation compared with Ta control (p < 0.0001, p < 0.0001 and p< 0.0001 respectively **). No significant differences in gap filling, bone ingrowth and total bone formation were found between Ta-CaP and Ta. (p = 0.12, p = 0.07, and p = 0.07). 63 Table 5.1 The p-values of t-tests on bone formation, gap filling and ingrowth Compared groups Bone formation, v% Gap filling v% Ingrowth v% Ta & Ta-CaP 0.07 0.12 0.07 Ta & T a - C a P - A L N < 0.0001 < 0.0001 < 0.0001 Ta-CaP & T a - C a P - A L N < 0.0001 O.0005 < 0.0001 5.5.2 Bone/implant contact length The direct contact between newly grown bone and the implant is a very important factor that affects the mechanical fixation of implants. This is especially critical when there is a gap existing between the implant and its surrounding host bone, the contacts between newly formed bone and the implant directly dictate the implant fixation. In the 2D sections, the contacts between bone and implant only encounter new growth bone because the gap model design avoids any press-fit contacts of host bone and implants. Also , the bone/implant contact interfaces in 3D structure becomes the border lines between bone phase and implant metal phase in the 2D images. The length of these contact border lines directly represents how well the new bone contacts the implants. Therefore, the bone/implant metal contact length can be used as a criterion to judge the bone/implant contact and the implant fixation. The mean percentages of bone-implant contact length over total available tantalum length of the three sections were measured (Fig 5.9). In section 1 (200 micrometer deep from tangent surface of implant), the percentage of contact length was 8.6% (SD 3.9%) for the Ta controls; 13.2% (SD 4.9%) for Ta-CaP controls and 52.3% (SD 18.3%) for T a - C a P - A L N implants. In section 2 (850 micrometer deep), the percentage contact length was 3.8% (SD 64 2.1%) for Ta controls;..8.1% (SD 2.8%) for Ta-CaP controls and 39.4% (SD 14.3%) for Ta-C a P - A L N implants. In section 3 (the central surface of implant), the percentage contact length was 4.6% (SD 1.4%) for Ta controls; 9.0% (SD 3.0%) for Ta-CaP controls and 40.0% (SD 17.6%) for T a - C a P - A L N implants. The percentage of contact length of T a - C a P - A L N implants increased by an average of 700% (8-fold) and 342% (4.42 fold) compared with those of Ta and Ta-CaP controls respectively. The average increase of contact for Ta-CaP controls was 87% (1.87-fold) compared with that of Ta controls. The t-test results show that the enhancements of bone/implant contacts at the site of T a - C a P - A L N samples were significant compared with Ta and Ta-CaP controls. The p-values were less than 0.0001 in both cases. Also , the increase of contact at the site of Ta-CaP controls was significant compared with Ta controls because the p-value oft-test was 0.02. Table 5.2 The p-values of t-tests on bone/implant contacts of 3 sections Compared groups Section 1 Section 2 Section 3 Ta & Ta-CaP 0.02 0.0004 0.0002 Ta & T a - C a P - A L N O.0001 O.0001 O.0001 Ta-CaP & T a - C a P - A L N O.0001 O.0001 O.0001 65 C? 100 CO •£ 80 c OD O i <D CL £ 60 CT) c d) _l o 40 CO c o O 20 I 0 c o ** * Section 1 ** * Section 2 Ta T a - C a P T a - C a P - A L N * Section 3 Fig . 5.9 Bone formation on the surface of porous tantalum expressed as a percentage of the available porous tantalum surface (contact length) o f three sections performed through each implant. T a - C a P - A L N demonstrated significantly increased contact length compared with Ta-CaP controls, p < 0.0001 *, and significantly increased contact length compared with Ta controls, p < 0.0001 **. Ta-CaP controls demonstrated significantly increased contact length compared with Ta controls, p < 0.02 +. 5.6 Bone formation mechanism Fluorescence labeling provides the information on when and where the bone formation happens and the directions of bone growth. Base on the fluorescent images, some facts regarding the bone formation mechanism were observed and the comparison among 2 control groups and T a - C a P - A L N implant group was conducted. Fig. 5.10 shows how the fluorescence labeling works. F ig . 5.10 (a), which is the transverse section of a tooth, shows the fluorescent rings. The growth direction can be identified by these rings with reference to the administration sequence o f fluorochromes (see Table 4.4 on page 45). F ig . 5.10 (b) shows the bone initiation on a Ta-CaP control surface. 66 A l l samples (108 sections from 36 implants) revealed the new bone formation initiated at the surfaces of the host bone and bridged over the gap towards metal implants during the first 2 weeks after surgeries. This is called the distance osteogenesis [26], and the distance osteogenesis w i l l occur regardless of the types o f implants because the host bone is the best places for new bone formation. F ig . 5.10 Fluorescence images showing the labels on tooth and ingrowth bone, and the ordinal numbers gave postoperative time sequence, a) 4 color rings indicated the growth direction on tooth. The 2 green rings were calcein and 2 red rings were alizarin; b) labels indicated the bone initiation on implant surface. The major difference between T a - C a P - A L N implants and 2 control groups was the interaction o f newly formed bone with the implant surfaces. The pattern and site o f bone initiation were strongly modulated with alendronate-immobilized calcium phosphate coating. In the Ta controls, bone formation occurred inside the tantalum pores with little attachment to the surface o f porous tantalum. Fig. 5.11(a) shows the bone initiation (the green lines) 67 occurred around the tantalum strut (not touching) and expanded (the red lines) afterward. In contrast, early bone formation on T a - C a P - A L N plugs occurred predominantly on the surface of the tantalum struts. In Fig 5.11(b), the new bone initiation (green lines) spread on the implant surface and grew outward. Ta-CaP controls had a similar pattern of bone formation to that of the Ta controls, and bone initiation on implant surfaces was occasionally observed on the Ta-CaP controls. F ig 5.11 (c) shows similar bone initiation on a Ta-CaP control to that on a Ta control. Fig. 5.11 (d) demonstrates the bone initiation on the implant surface of a Ta-CaP control. But the initiation of bone on Ta-CaP controls was slightly different than that on the T a - C a P - A L N implants. It only covered a portion of implant surface and expanded into pore center. The bone ingrowth patterns were different on T a - C a P - A L N implants from those on the two types of controls because of the change in bone formation initiation. On Ta controls (see Fig . 5.12), early bone formation occurred in the center of the porous area and expanded outwards onto the porous tantalum struts, as was tracked by the fluorescence labels (green lines were for first and second postoperative weeks, and red lines were for third and fourth weeks). The bone ingrowth went through the channel of tantalum pores and barely touched the struts. On Ta-CaP controls (see F ig . 5.13), the similar ingrowth pattern to that of Ta controls could be seen, and also a branch of bone ingrowth anchored on implant surface (Fig 5.13 (b) left bottom corner). On T a - C a P - A L N implants (see Fig . 5.14), the bone ingrowth spread on the implant surface and followed by bone expansion towards the center of the porous channels. 68 Fig . 5.11 Fluorescence images showing bone initiation, a) non-coated porous tantalum control; b) alendronate-immobilized calcium-phosphate-coated implant; c) and d)calcium-phosphate-coated controls. 69 Fig . 5.12 Fluorescence images showing bone ingrowth pattern on Ta control, a) low magnification image; b) higher magnification image showing the labeling lines. Fig . 5.13 Fluorescence images showing bone ingrowth pattern on Ta-CaP control. F ig . 5.14 Fluorescence images showing bone ingrowth pattern on T a - C a P - A L N implant. 70 The changes in bone formation mechanism also affected the patterns o f gap filling bone growth. On the Ta controls (see F ig . 5.15 (a)), new bone formation crossed over the gap area but did not anchor on the implant surface; on the Ta-CaP controls (see Fig . 5.15 (b)), there were two small portions o f new bone anchored on the implant surface; on T a - C a P - A L N implants (see Fig . 5.15 (c) and (d)), new bone fully contacted the implant surface. a C a p — • ! la b t iap ~~1 Ta ' * 300 u m 1 ^ ^ ^ ^ 1 V 300 um Fig . 5.15 Fluorescence images showing the bone formation at gap areas. The white lines with double arrows indicated the bone bridging gaps a) Ta control; b) Ta-CaP control; c) and d) T a - C a P - A L N implants. It is very important for bone/implant fixation that the new bone crossing gaps anchored on implant surfaces. Without the full bone/implant contact, there would be a possibility of micro-motion between implant and host bone. Without the bone formation 71 initiation on implant surface, the bone/implant integration would be postponed because the bone initiation on implant surfaces ensured a certain level of bone/implant integration in first 2 weeks postoperatively on T a - C a P - A L N implants compared with bone still barely touching the implant surfaces 4 weeks after the surgeries on the Ta controls. 72 CHAPTER 6: DISCUSSION 6.1 The effects of calcium phosphate coating Hydroxyapatite is a bioactive ceramic. Many reports on hydroxyapatite enhancing bone formation were published [90, 91,92]. Barrere et al [93] claimed that biomimatic coating of O C P increased the osteoinductivity (defined on page 12) of porous tantalum when implant locations were inside the back muscles goats, and he could not fully claim the increase of osteoinductivity by biomimatic coated O C P when the porous tantalums were implanted in distal femur. Davies JE [23] believed the B M P might adsorb on porous-structure hydroxyapatite and therefore enhance osteoinduction. However, with a.gap existing between implants and host bone, our animal tests did not reveal significant increase of peri-prosthetic bone formation. Although there were 21.5% and 38.3% increases of gap filling and ingrowth respectively, the p-values of t-tests were 0.12 and 0.07 respectively. Therefore, the osteoinductivity of porous tantalum implants was not significantly increased by calcium phosphate coating process when challenging bone/implant gaps and defects. Despite the weak effects on the osteoinductivity of implants, the calcium phosphate coating significantly increased the new bone-formation/implant contacts. Our testing results showed an average of 87% (1.87-fold) increase of contact lengths on Ta-CaP implants compared with those on Ta controls. With the help of fluorescence images, the initiation of new bone formation on the strut surfaces, which was also called contact osteogenesis, was occasionally observed on Ta-CaP implants compared with little bone initiation on implant 73 surfaces of Ta control. These results provided the direct evidences and explanations to why bone/implant contacts were increased by calcium phosphate coating process. The initiation of new bone formation on calcium phosphate coating surfaces indicated better attachment of osteoblasts and osteogeneritor cells. In another words, the calcium phosphate coating process increased the osteoconductivity (defined on page 12) of porous tantalum implants. Theoretically, the increase of bone /implant contact may contribute to the eventual implant fixation. Summer et al [46] reported that the hydroxyapatite-tricalcium phosphate coating did not contribute to fixation, bone ingrowth, and bone formation based on their mechanical testing and histomorphometry results of a gap model animal test on B M P treated implants. They analyzed the fixation strength, interface stiffness and energy to failure and found no difference with or without hydroxyapatite-tricalcium phosphate coatings. They also analyzed the bone volume, trabecular numbers and trabecular plate thickness etc on both calcium phosphate coated samples and the controls. They found no difference either. Their results supported our findings. In conclusion, in the presence of a gap between bone and the implant, our E L D calcium phosphate coating increases the bioactivity of porous tantalum implants in terms of osteoconductivity. The testing results show evidence of better osteoconductivity and contact osteogenesis on calcium phosphate coating surfaces compared with non-coated tantalum surfaces. However, the calcium phosphate coating treatment did not significantly change the 74 osteoinductivity of porous tantalum, and therefore the peri-prosthetic new bone formation was not increased. A s a result, the bone/implant fixation might not be obviously improved by calcium phosphate coating process. 6.2 The effects of alendronate-immobilized calcium phosphate coatings Our gap model study demonstrated a significant enhancement of bone ingrowth (19.9 % versus 6.0%), bone gap filling (29.1% versus 13.0%) and total bone formation (26.2 % versus 9.7%) associated with T a - C a P - A L N samples compared with Ta controls. A t the same time, an alendronate surface treatment significantly enhanced bone ingrowth (140 %; 2.40-fold), gap fil l ing (84%; 1.84-fold) and total bone formation (111%; 2.11-fold) in comparison with Ta-CaP controls. Possible explanations of the enhancement of bone formation towards implants could be: the effects of alendronate on bone cells increased the activities of osteoblasts work and decreased the activities of osteoclasts; or the alendronate-immobilized calcium phosphate coating increased osteoinductivity of porous tantalum which meant more stem cells differentiated to osteoblasts. Davies JE [23] believed the osteoblasts would be trapped in bone and became osteocytes after the mineralization of the bone matrix. For further bone growth, new osteoblasts needed to be recruited. In order to clearly understand the mechanism behind the enhancement of new bone formation, further studies need to be conducted. The histomorphometry analysis also showed that the alendronate-immobilized calcium phosphate coating significantly increased bone/implant contacts compared with Ta-75 CaP controls (342%; 4.42-fold) and Ta controls (700%; 8-fold). The fluorescence imaging found that the initiation of bone formation happened on the implant surface throughout the T a - C a P - A L N implants. The osteoconductivity of these implants was therefore dramatically increased compared with both Ta-CaP and Ta controls. The significant increase of contact osteogenesis on alendronate-immobilized calcium phosphate coating indicated that it was the best bioactive implant surface out of the three types of implants. The X R D , FTIR and S E M imaging analysis of coatings revealed no changes on morphology and crystal structure of the E L D calcium phosphate coating after alendronate immobilization. This eliminated any other potential factors, such as a change in the crystal structure of calcium phosphate from O C P to H A , that happened and affected the bioactivity of the coating surfaces. Therefore, the dramatic enhancements of bone formation and osteoconductivity of porous tantalum implants were essentially due to the presence of alendronate. The results of our study are also consistent with recent in vivo tests reported in the literature. The effect of zoledronate on bone ingrowth into porous tantalum implants was studied in a canine model by a single post-operative intravenous injection [67]. Subsequently, zoledronate was bound to hydroxyapatite-coated porous tantalum, and an increase of 134% in bone ingrowth was reported by the. same group [68]. In our gap model with a less potent bisphosphonate (alendronate), however, we found a greater relative increase in bone ingrowth compared with the zoledronate study (232% versus 134%) [68]. 76 V The differences in our findings may be related to several factors including differences in animal models, the type of bisphosphonate used, length of implantation, location of implantation and the design of the implants. Another key difference between our study and previous studies is the unique micro-porous calcium phosphate coating made by E L D technique. Such a coating may not significantly affect bone formation itself, but it acts as a very effective carrier for bisphosphonate delivery. Because of the effects of alendronate-immobilized calcium phosphate coating on osteoconductivity of porous tantalum implants, bone reaction to implants was completely changed, and the initiation of bone formation happened on the surface of the alendronate-immobilized calcium phosphate coatings. In addition, alendronate-immobilized calcium phosphate coating increased new bone formation towards implants. A s a result, more bone formed bridging the gap between host bone and implant and anchored onto the implant surface, and eventually better bone/implant fixation could be achieved. 6.3 Clinical relevance of the gap model animal tests and alendronate delivery implants It has been known that biological fixation of porous implants occurs by micro-interlocking with ingrown tissue. Autopsy studies have demonstrated that bone ingrowth occurs in only 30-40% of the surface area of the implant in primary total hip arthroplasty [94]. In revision hip arthroplasties, bone defects and gaps may significantly compromise the quantity and quality of host bone available for ingrowth. In many cases less than 50% host 77 bone is available for bone ingrowth of the newly implanted shell. Therefore, any solutions to achieve a reliable and durable bone ingrowth to porous materials would potentially lead to novel implants for challenging revision cases. Gaps and defects that occur between the surface of the implant and the host bone in an acetabular reconstruction are usually filled with bone graft or bone graft substitutes in current clinical practice. However, bone ingrowth into porous implants is poor in the presence of acetabular defects, regardless of the defect filling with bone graft. In an earlier study, acetabular defects were created in a canine model which were left unfilled or were filled with either autograft bone or a 50:50 mixture of autogenous bone graft and hydroxyapatite/tricalcium phosphate ( H A / T C P ) ceramic [95]. A t 6 weeks post operation, grafting improved defect healing but had inferior bone ingrowth compared with empty defects. In our study, a rabbit implant model incorporating a gap between porous tantalum and bone was developed to assess the effect of alendronate coated porous tantalum compared against the controls. The model sought to replicate the clinical scenario in total hip revision arthroplasty. Implantation of the plugs in the metaphyseal region also closely replicates the clinical scenario in that the implants were placed in cancellous bone. The positive results of our study indicate that the surface treatment technique could provide an effective solution to the clinical challenges discussed above. Another application is the spinal fixation where fast and more bone ingrowth to the implants are critical to the function. 78 6.4 Key progresses in ELD surface drug delivery and animal studies A unique technique in this study is the electrolytic deposition and the resulting micro-porous calcium phosphate coating. The calcium phosphate coating itself does not seem to have significant impact on bone growth. This relative uniform calcium phosphate coating, however, provides an ideal surface carrier for bisphosphonate local delivery. The electrolytic deposition used in this study has obvious advantages over plasma spray and biomimetic coating for porous Ta implants. In plasma spray, only a half thickness of the porous cells on Ta implant surface can be coated with calcium phosphate because o f the line-of-sight limitation. A s a result, bone ingrowth into deeper porous area can not be improved. Biomimetic coating technique is a relatively slow process. It requires careful surface pre-treatment and strict control over the solution conditions. The calcium phosphate coatings tend to form spherulites. Electrolytic deposition with periodical stirring could make a relative uniform and porous coating both inside and on the surface of the porous Ta implants. Another unique feature of our animal tests is the introduction of fluorochrome injections to the bone formation mechanism studies. The fluorochromes mark the front of bone formation, and many people have used a series of fluorochromes to reveal bone growth direction and growth speeds etc. We employed this fluorescence technique to examine the mechanism of bone formation on different types of surfaces of porous tantalum implants and visualize the initiation of bone formation on the surface of alendronate-immobilized calcium phosphate coating. 79 Alendronate-immobilized calcium phosphate coating provides an advanced drug delivery method that can directly send alendronate to symptom location. It dramatically reduces the drug dose without compromising any healing effects. Also , alendronate, as one of bisphosphonates, has many advantages over many other materials such as hydroxyapatite and B M P . Other methods of augmenting the rate and extent of bone ingrowth into porous implants have been described in the literature, including the use of B M P [46, 96], Bone-like Carbonated Apatitic ( B C A ) coating [93], bone matrix gel [97] and non invasive low intensity ultrasound [98]. The advantage of bisphosphonates over B M P is their relatively low cost. Other advantages of bisphosphonates are their established clinical record, their effects on reducing bone resorption thereby decreasing stress shielding [99] and wear particle-induced osteolysis [41]. Improvements to the design and biomaterial properties of any drug-implant delivery system can only be made by understanding the principles underlying the mechanism of action of the drug- delivery system. The mechanism of enhanced bone ingrowth and osteoconductivity of T a - C a P - A L N provides further insight into the underlying action of this drug delivery system. Enhancement of bone ingrowth, gap filling and increased osteoconductivity of T a - C a P - A L N could have very significant advantages in revision joint arthroplasties associated with host bone defects, providing early and enhanced biological fixation in complex cases. 80 CHAPTER 7: CONCLUSIONS A gap model animal test was designed and successfully conducted to study the effects of alendronate-immobilized calcium phosphate coatings on bone formation towards porous tantalum implant. The gap model imitated the clinical scenario. This and other unique features, such as the porous E L D coatings of calcium phosphate, alendronate immobilization and fluorescence labeling etc, distinguished this animal test from any others and represented significant improvements in implant/bone interaction study. Following are the main conclusions drawn from this animal test: 1. Alendronate-immobilized calcium phosphate coating enhanced the bone formaton towards porous tantalum implants. The treatment could significantly accelerate early stage (4 weeks) bone growth (by 170%) as compared with porous Ta by itself. The increase was 111% when compared with Ta-CaP controls. Although Ta-CaP implant increased total bone formation by 28%, the improvement was not statistically significant. 2. Both calcium phosphate coatings before and after alendronate immobilization significantly increased the osteoconductivity of porous tantalum implants. Alendronate immobilized calcium phosphate coating achieved much greater enhancement. Alendronate immobilization treatment demonstrated dramatic and significant increase in bone/Tantalum contact. A s compared with porous tantalum itself, T a - C a P - A L N implant increased bone coverage on tantalum by 770%. The increase was 344% when compared with Ta-CaP controls. Calcium phosphate 81 coating increased the bone/implant contact length by 87% compared with Ta controls. 3. The observation of fluorescence labeling images showed that the treatment of alendronate-immobilized ELD-coating of calcium phosphate changed the new bone formation mechanism. New bone formation initiated on the T a - C a P - A L N surfaces, and this provided the explanation why new bone/implant contacts increased on Ta-C a P - A L N implants. 4. Early bone/implant integration may be achieved on T a - C a P - A L N implants. The fluorescence images showed that new bone crossing gaps anchored on the implant surface in the first two weeks postoperatively. 5. Alendronate-immobilized calcium phosphate coating on porous tantalum has advantages in overcoming the gaps between the bone and the implant, and the surface design can be potentially employed to solve some of the clinical issues associated with revision surgery. 82 CHAPTER 8: RECOMMENDATION FOR FUTURE WORK This animal test provided very promising results on the enhancements of osteoconduction and bone ingrowth on porous tantalum by alendronate-immobilized calcium phosphate coating. It is a technique that has potential clinical applications. Before any clinical trial could be discussed, the following experiments need to be conducted for further confirmation of the results. 1. Systematic study on the effect of alendronate loading amounts to optimize the drug dosage and to achieve the best effects on bone/implant fixation. 2. 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Cook SD et al. The effect of demineralized bone matrix gel on bone ingrowth and fixation of porous implants. J Arthroplasty 17:402-408, 2002. 98. Tanzer M et al. Enhancement of bone growth into porous intramedullary implants using non-invasive low intensity ultrasound. J Arthroplasty 2002; 19:195-199. 90 99. Wilkinson J M et al. Effect of pamidronate in preventing local bone loss after total hip arthroplasty: a randomized, double-blind, controlled trial. J Bone Miner Res 2001; 16: 556-64. 91 APPENDICES Appendix A: Student's t table (reprinted from ref 82 courtesy of Deacon J.) Degrees of Freedom Probability, p 0.1 0.05 0.01 0.001 1 6.31 12.71 63.66 636.62 2 2.92 4.30 9.93 31.60 3 2.35 3.18 5.84 12.92 4 2.13 2.78 4.60 8.61 5 2.02 2.57 4.03 6.87 6 1.94 2.45 3.71 5.96 7 1.89 2.37 3.50 5.41 8 1.86 2.31 3.36 5.04 9 1.83 2.26 3.25 4.78 10 1.81 2.23 3.17 4.59 11 1.80 2.20 3.11 4.44 12 1.78 2.18 3.06 4.32 13 1.77 2.16 3.01 4.22 14 1.76 2.14 2.98 4.14 15 1.75 2.13 2.95 4.07 16 1.75 2.12 2.92 4.02 17 1.74 2.11 2.90 3.97 18 1.73 2.10 2.88 3.92 19 1.73 2.09 2.86 3.88 20 1.72 2.09 2.85 3.85 21 1.72 2.08 2.83 3.82 22 1.72 2.07 2.82 3.79 23 1.71 2.07 2.82 3.77 24 1.71 2.06 2.80 3.75 25 1.71 2.06 2.79 3.73 26 1.71 2.06 2.78 3.71 92 27 1.70 2.05 2.77 3.69 28 1.70 2.05 2.76 3.67 29 1.70 2.05 2.76 3.66 30 1.70 • 2.04 2.75 3.65 60 1.67 2.00 2.66 3.46 120 1.66 1.98 2.62 3.37 00 1.65 1.96 2.58 3.29 93 

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