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Calcium phosphate coatings on coronary stents by electrochemical deposition 2006

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C A L C I U M PHOSPHATE COATINGS ON C O R O N A R Y STENTS BY E L E C T R O C H E M I C A L DEPOSITION by M A N U S P U I - H U N G TSUI B . A . S c , Department of Metals and Materials Engineering, The University of British Columbia, Vancouver, B . C . , Canada, 2004 A THESIS S U B M I T T E D I N P A R T I A L F U L F I L L M E N T O F T H E R E Q U I R E M E N T F O R T H E D E G R E E O F M A S T E R O F A P P L I E D S C I E N C E in T H E F A C U L T Y OF G R A D U A T E S T U D I E S (Materials Engineering) T H E U N I V E R S I T Y OF B R I T I S H C O L U M B I A October 2006 © Manus Pui-Hung Tsui, 2006 A B S T R A C T Calcium phosphate ceramic coatings, especially hydroxyapatite (HA) , have attracted much attention in the orthopedics and dentistry due to their excellent biocompatibility and bioactivity. Among the different methods of calcium phosphate coatings processing, electrochemical deposition (ECD) is a relatively low cost and flexible process technology. In this study, electrochemical deposition was used to deposit uniform calcium phosphate coatings on 316L stainless steel coronary stents. The influence of the E C D process parameters (deposition time, current density, electrolyte temperature, p H , and Ca/P ratio) on the resulting deposition morphology was investigated. Scanning electron microscopy (SEM) and X - R a y diffractometry ( X R D ) were used to analyze the coatings. The results demonstrated that both dicalcium phosphate dihydrate ( C a H P 0 4 - 2 H 2 0 , D C P D ) and hydroxyapatite (Ca 1 0 (PO 4 ) (OH) 2 , H A ) were present in the uniform -0.5 um thick as-deposited coating. However, a,post- - treatment process, including a 0.1N NaOH( a q ) phase conversion at 75°C and a 500°C heat treatment produced a pure phase H A coating. The final deposit revealed a highly porous surface morphology which could be useful for drug encapsulation. Wi th the application of the substrate surface modification and the post-treatment processes, sufficient coating adhesion was achieved as demonstrated by the in-vitro stent deployment tests without visible damage to the coating. Commercial in-vitro 40 mil l ion cycles fatigue tests demonstrated that the coatings exhibit good adhesion to the stent substrate, with no coating cracking or delamination. It was confirmed that the E C D - H A coating process for coronary stents is reliable and reproducible. i i T A B L E OF CONTENTS A B S T R A C T i i T A B L E OF C O N T E N T S i i i LIST OF T A B L E S v i LIST OF F I G U R E S v i i A C K N O W L E D G M E N T S x i i i 1 I N T R O D U C T I O N 1 1.1 Biomaterial Coatings for Stents 2 1.2 Calcium Phosphates and Hydroxyapatite 3 1.3 Electrochemical Deposition of H A Coatings 5 1.4 Motivation and Focus of the Present Study 6 2 L I T E R A T U R E R E V I E W 8 2.1 Coronary Heart Disease (CHD) 8 2.2 Coronary Artery Stent 12 2.2.1 Coated Stents 16 2.2.2 Biodegradable Stents 17 2.2.3 Radioactive Stents 17 2.2.4 Drug Eluting Stents 18 2.3 Bioceramics 19 2.3.1 Calcium Phosphate Bioceramics 22 2.3.1.1 Bioresorption and Biodegradation 23 2.3.1.2 Mechanical Properties 25 2.3.2 Hydroxyapatite Bioceramics 27 i i i 2.3.2.1 Thermal Stability 28 2.3.2.2 Mechanical Properties 29 2.4 Biomaterial Coatings 30 2.4.1 Hydroxyapatite Coatings 31 2.4.2 Processing of Hydroxyapatite Coatings 32 2.5 Electrochemical Deposition 34 2.5.1 Electrochemical Deposition of Hydroxyapatite Coatings 36 2.5.2 Microstructure and Phase of E C D Calcium Phosphate Coatings 40 2.5.3 Adhesion of H A Coatings 43 3 S C O P E A N D O B J E C T I V E S 46 3.1 Scope of the Investigation 46 3.2 Objectives 47 4 E X P E R I M E N T A L M E T H O D O L O G Y 49 4.1 Sample Preparation 50 4.1.1 Substrate Surface Modification 51 4.2 Electrochemical Deposition 52 4.2.1 Electrochemical Deposition Process Parameters Investigation 53 4.2.2 Electrochemical Deposition Optimization 55 4.2.3 Phase Conversion Process 55 4.3 Microstructural and Phase Characterizations 56 4.4 In-vitro Evaluations 56 4.4.1 Crimping and Expansion Test 57 4.4.2 Dissolution Test 58 iv 4.4.3 Fatigue Test 58 5 R E S U L T S A N D D I S C U S S I O N 61 5.1 E C D of Calcium Phosphate Coatings - Process Parameters Investigation 61 5.1.1 Current Density 61 5.1.2 Deposition Time 64 5.1.3 Ca/P Ratio 68 5.1.4 Temperature 70 5.1.5 The Influence of Electrolyte p H 72 5.2 E C D of Calcium Phosphate Coatings on Coronary Stents 74 5.2.1 Deposition Process Optimization 74 5.2.2 In-vitro Crimping and Expansion Tests on E C D Coated Stents 81 5.2.3 Substrate Surface Modification for Improvement o f Coating Adhesion 84 5.2.4 Phase Composition of E C D Calcium Phosphates Coatings 90 5.2.5 In-vitro Fatigue Test 98 5.2.6 Reproducibility and Consistency of E C D - H A Process 108 5.2.6.1 Errors in E C D - H A Characteristics and Process Parameters Measurement 109 6 C O N C L U S I O N S '. I l l 7 R E C O M M E N D A T I O N S F O R F U T U R E W O R K 114 R E F E R E N C E S 116 A P P E N D I X A - P R E L I M I N A R Y R E S U L T S O N E C D C O - D E P O S I T I O N O F O R G A N O - C E R A M I C C O A T I N G S 124 A P P E N D I X B D E T A I L E D R E C O R D S OF R E P R O D U C I B I L I T Y S T U D Y 128 v LIST OF T A B L E S Table 2.3-1. Various calcium phosphate bioceramics 23 Table 2.3-2. Mechanical and Physical Properties of Calcium Phosphates 26 Table 2.4-1. Purposes of biomaterial coatings 30 Table 2.4-2. Summary of techniques used for deposition of hydroxyapatite coatings..... 34 Table 4.1-1. Nominal chemical composition of 316L stainless steel 50 Table 5.1-1. The concentration of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 for the preparation of different Ca/P ratio electrolytes 69 Table 5.2-1. Optimum parameters for E C D of calcium phosphate coatings* 75 Table 5.2-2. Summary of E C D coatings dissolution test data. Dissolution tests were conducted with 10 m L of phosphate buffer saline (PBS) at 37°C (pH = 7.4) with rotation speed at 20 rpm 97 Table 5.2-3. Summary of five batches of E C D coating average weight and yield rate. 109 Table B 1. Preparation record for E C D coating batch ECD-RP-001 128 Table B 2. Preparation record for E C D coating batch ECD-RP-002 129 Table B 3. Preparation record for E C D coating batch ECD-RP-003 ..130 Table B 4. Preparation record for E C D .coating batch ECD-RP-004 131 Table B 5. Preparation record for E C D coating batch ECD-RP-005 132 vi LIST OF FIGURES Figure 2.1-1. Fat and cholesterol accumulated oh the inside of coronary arteries 9 Figure 2.1-2. Schematic of percutaneous transluminal coronary angioplasty technique. A guide-wire is placed across the blocked section of the artery and a balloon is positioned beside the blockage. The balloon is then inflated compress the blockage against the artery wall 11 Figure 2.1-3. Schematic of percutaneous transluminal coronary angioplasty technique with the use of coronary artery stent. Coronary artery stent is used as mechanical scaffold to provide support to the vascular wall during and after the P T C A procedures 11 Figure 2.3-1. Logarithm of the product of calcium and phosphate concentrations plotted against p H values of solution saturated with respect to various calcium phosphate phases in the ternary system Ca(OH)2-H3P04-H20. Calculated for 37°C 25 Figure 2.5-1. Cathodic polarization curve of T i substrate in a Ca(N03)2 4 H 2 O and N H 4 H 2 P O 4 electrolyte 39 Figure 2.5-2. X R D analysis of coatings deposited at current density of 1, 5, 10, 15, and 20 mA/cm2 for 30 min 43 Figure 4 .1-1. Schematic diagram of M I V I 700 Series Coronary Stent 51 Figure 4 .2-1. Schematic diagram of electrochemical deposition setup 53 Figure 5.1-1. S E M images of E C D coating deposited at various current densities: (a) 15 m A / c m 2 , (b) 10 m A / c m 2 , (c) 5 m A / c m 2 , (d) 3 m A / c m 2 , (e) 1 m A / c m 2 [Left: xlOO; Right: x 1,500] 63 v i i Figure 5.1-2. Weight gain of E C D coated specimens versus current density with 5 minutes of deposition 64 Figure 5.1-3. S E M images of E C D coating deposited with various deposition time: (a) 15 min, (b) 10 min, (c) 5 min, (d) 3 min, (e) 1 min [Left: xlOO; Right: x l500] at 1 m A / c m 2 65 Figure 5.1-4. High magnification S E M image of E C D coating with deposition of 1 min [x20,000] at I m A / c m 2 67 Figure 5.1-5. Weight gain of E C D coated specimens versus deposition time deposited at 1 m A / c m 2 67 Figure 5.1-6. Ca/P ratio of resulting deposit with the use of various Ca/P ratio electrolytes. Ca/P ratio was derived from E D X spectra 69 Figure 5.1-7. S E M images of resulting deposits from various Ca/P ratio electrolytes; a) 2.92, b) 2.63, c) 1.95, and d) 0.49. [x 15,000] 70 Figure 5.1-8. S E M images of resulting CaP deposits conducted in electrolyte with temperature a) 25 °C b) 45°C, c) 75°C. [x 15,000] 71 Figure 5.1-9. Influence of electrolyte temperature on measured supply voltage, for constant current source (I = 13.77 mA) 72 Figure 5.1-10. S E M Image of E C D conducted with p H 3.0 electrolyte at 45°C. [x5,000] 73 Figure 5.1-11. S E M images of E C D coatings deposited with various electrolyte p H : a) 4.0, b) 4.5, and c) 5.5. [x3,000] 73 Figure 5.2-1. S E M images of bare metal stent with various magnification: a) [xlOO], b) [x800],c) [x 1,500] 76 v i i i Figure 5.2-2. S E M images of E C D coating deposited on coronary stent with optimum parameters for 1 minute, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] 77 Figure 5.2-3. S E M images of E C D coating deposited on coronary stent with optimum parameters for 2 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] 77 Figure 5.2-4. S E M images of E C D coating deposited on coronary stent with optimum parameters for 3 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] 78 Figure 5.2-5. S E M images of E C D coating deposited on coronary stent with optimum parameters for 5 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] 78 Figure 5.2-6. E D X surface analysis of the E C D coating deposited on coronary stent with optimum parameters for 1 minute 79 Figure 5.2-7. X - R a y diffraction of the E C D coating deposited with optimum parameters (Table 5.2-1), for 1 minute, showing a mixed phase of D C P D and H A . . . 80 Figure 5.2-8. Cross-section S E M image of E C D coating deposited on stent. Estimated coating thickness was approximately 0.5 urn 81 Figure 5.2-9. (a) S E M images of an expanded bare metal stent [xlOO], (b) high magnification revealing a significantly deformed surface [x3,000] 82 Figure 5.2-10. S E M images of expansion test result from an E C D coated stent specimen deposited with optimum parameter for 1 minute deposition, (a) Expanded area [x50] (b) Expanded area [x300] (c) Compressive stress area showing ix coating delamination [x800] (d) Tensile stress area showing coating delamination [x800] 83 Figure 5.2-11. Compressive spallation by buckling showing localized interfacial decohesion 83 Figure 5.2-12. Tensile stress in brittle f i lm causing through-thickness cracking and interfacial delamination 84 Figure 5.2-13. S E M image of a stent surface after surface modification [x20,000]. The surface showed a nano-size roughness 85 Figure 5.2-14. E D X surface analysis of a surface modified stent 86 Figure 5.2-15. Expansion test result from a bare metal stent specimen after surface modification, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [xlO,000]. The Na4Cr04 inter-compound can be seen with cracks 86 Figure 5.2-16. S E M images of E C D coating deposited with optimum deposition parameters on surface modified stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 87 Figure 5.2-17. S E M images of expansion test result from an E C D coating deposited with optimum deposition parameters on surface modified stent. Expansion performed with a 3.0 mm diameter catheter, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 88 Figure 5.2-18. S E M images of expansion test result from an E C D coating deposited with optimum deposition parameters on surface modified stent. Expansion performed with a 3.5 mm diameter catheter, (a) [xlOO], (b) [x800], (c) [x3,000],and (d) [x 10,000] 90 Figure 5.2-19. X - R a y diffraction patterns of as-deposited E C D coating and the resulting coating after 12, 24, 48, and 72 hours of NaOH( a q ) phase conversion at 75°C 92 Figure 5.2-20. S E M images of a 12 hours phase conversion E C D coating deposited with optimum deposition parameters, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 93 Figure 5.2-21. S E M images of a 72 hours phase conversion E C D coating deposited with optimum deposition parameters, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 93 Figure 5.2-22. X - R a y diffraction patterns of 12 hours NaOH( a q ) phase converted E C D coating and after 300°C, 500°C, and 750°C for 20 minutes of heat treatment of the coating 95 Figure 5.2-23. S E M images of E C D coating after phase conversion process [x 10,000]. (a) 12 hours N a O H ( a q ) treatment (b) 12 hours N a O H ( a q ) treatment + 500°C heat treatment (c) 12 hours NaOH( a q ) treatment + 750°C heat treatment.. 97 Figure 5.2-24. S E M images of expansion test result from an as deposited E C D coating upon NaOH( a q ) treatment + 500°C heat treatment. Expansion performed with a 3.5 mm diameter catheter, (a) [x50], (b) [x300], (c) Showing the compressive stress area [x 1,500], and (d) Showing the tensile stress area [x 1,500] 98 Figure 5.2-25. Summary of average % O D strain for the six fatigue tested stent specimens 99 x i Figure 5.2-26. S E M images of explanted fatigue tested specimen IP from vessel #1 proximal end (a) [xlOO], (b) [x800], (c) [x3,000] (d) E D X analysis from surface area scan 101 Figure 5.2-27. S E M images of explanted fatigue tested specimen 2 from vessel #2 proximal end (a) [xlOO], (b) [x800], (c) [x3,000] (d) E D X analysis from surface area scan 102 Figure 5.2-28. S E M images of explanted fatigue tested specimen 2 from vessel #3 proximal end (a) [xlOO], (b) [x800], (c) [x3,000] (d) E D X analysis from surface area scan 103 Figure 5.2-29. The six main categories of debris found on filter after the fatigue test.. 106 Figure 5.2-30. E D X analysis of the six main categories of debris found after the fatigue test 107 Figure A 1. S E M images of an E C D coating co-deposited with 0.1 wt% of P V A on coronary stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 125 Figure A 2. S E M images of an expanded stent coated with E C D co-deposited with 0.1 wt% P V A . (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [xl0,000]. . 126 Figure A 3. S E M images of an E C D coating co-deposited with 0.8wt% P V A on coronary stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 126 Figure A 4. S E M images of an expanded stent coated with E C D co-deposited with 0.8wt% P V A . (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [xl0,000]. . 127 x i i A C K N O W L E D G M E N T S I wish to express my sincere gratitude to the many people who have kindly provided me with great contribution to the completion of this thesis. I owe particular thanks to my supervisor, Dr Tom Troczynski for his advice and guidance during this project. I like to extend my appreciation to my fellow graduate students, faculty, and staff in the Department of Materials Engineering for all their support. Also , I thank M I V Therapeutics who has supported me and this research project in many ways. Special thanks are owed to my family and friends, whose have supported and encouraged me throughout my years of education. x i i i 1 INTRODUCTION Coronary artery disease ( C A D ) is the leading cause of death in North America. In 2004, approximately 54% of all cardiovascular deaths are due to coronary artery disease1. C A D occurs when fat (cholesterol) deposit blocks the arteries, reducing the oxygen supply to the heart muscle. Angioplasty is a technique of opening a narrowed blood vessel without having to resort to a major bypass surgery. Metallic stents have been used successfully by cardiologists to battle the ravages of C A D . A stent is a stainless steel or Co-Cr alloy wire mesh tube designed to keep arteries open after the angioplasty procedure. Before implantation, stent is put over a balloon catheter and collapsed into a smaller diameter, then positioned to the site of blockage. When the balloon is inflated, the stent expands, locks in place and forms a scaffold. Stent stays in the artery permanently, keeps it open, improves blood flow to the heart muscle and. relieves symptoms such as chest pain. Although metallic stent is an effective solution in providing structural support, it does not eliminate the recurrence of blockage (restenosis) in the artery in all cases. The bare metal stent may trigger a natural inflammatory response resulting in proliferation of smooth muscle cells 2 , thus, narrowing or re-closing of the artery, which often requires a repeat operation within a year. Unlike restenosis, which is fairly common, stent thrombosis is a rare but much more dangerous complication after coronary stent placement. A thrombus is dangerous because it can block a blood vessel partially or entirely, cutting off blood flow to the area supplied by that vessel. A thrombus w i l l produce different symptoms depending on where it forms. A coronary thrombus wi l l 1 produce chest pain and may result in artery blockage leading to a heart attack . A thrombus in the brain w i l l result in a temporary ischemic attack (TIA) or even a stroke. Restenosis, thrombosis, and inflammation are the problems associated with metallic stent implantation in coronary arteries. Applying biocompatible coatings on the stents, in particular drug eluting coatings, has been taken as a new approach to solve these problems 3 ' 4 . The new generation of coated coronary stents should be less restenotic, thrombogenic, be more acceptable by body environment, and should be capable of locally deliver drug to the surrounding tissue. 1.1 Biomaterial Coatings for Stents A biomaterial is any synthetic material used to replace part of a l iving system or to function in intimate contact with living tissues6. The ability of such material, device or system to perform without a clinically significant host response in a specific application is defined as biocompatibility. The performance of a biomedical device is greatly depended on the material properties, design of the device, surgical techniques, health of patient, and the surface properties. Because of the complex interface of biomaterial- biological medium, surface properties is often the key to determine success of the device. The interfacial phenomena at the biomaterial-biological medium are complex and involve water adsorption, ion exchange, cell and bacterial adhesion, corrosion, and decomposition of the biomaterial. Nowadays, medical devices and implants manufacturing commonly involves a biomaterial coating process and, as a result, 2 provides an opportunity to tailor the surface characteristics of the biomaterial to a specific application without detrimentally affecting its bulk properties. Ultimately, biomaterial coatings combine the advantages of bulk material and coating material to create the best desired biocompatible environment. The state-of-the-art stents are coated with polymers carrying drugs. Unfortunately, there is accumulating evidence that after the drugs are leached from the coating, the polymer triggers a negative response from the surrounding tissue, referred to as "late restenosis"5. Consequently, there is a need for more biocompatible coatings on stents, which would not cause restenosis or thrombosis at any stage o f post-implantation, and with or without drugs. It is hypothesized that calcium phosphate ceramic may constitute such coating, providing the motivation for the present work. 1.2 Ca lc ium Phosphates and Hydroxyapati te Calcium phosphates have been used since 1969 for manufacturing various forms of implants, as well as for solid or porous coating on other implants. Calcium phosphate can crystallize into various forms depending on the calcium to phosphorus ratio, presence of water, impurities, and temperature6. Calcium phosphate ceramics, particularly hydroxyapatite (HA) , have received much attention and clinically applied in orthopedics and dental industries due to their excellent biocompatibility 7" 9. H A is recognized as osteoconductive and is able to accelerate bone ingrowth and attachment to the surface of implant during the early stages after implantation 8 ' 9 . In addition, it can induce tissue ingrowth and formation of chemical bonding to achieve better implant fixation. H A 3 interacts with the biological environment through a complex dissolution, precipitation, and ion exchange process 1 1 ' 1 2 . Following the introduction of H A to the biological environment, a partial dissolution of the surface is initiated causing the release of Ca , HPO4 2 " and PO43", increasing the supersaturation of the micro-environment with respect to the stable (HA) phase. This saturation leads back to the precipitation hydroxyl- carbonate apatite ( H C A ) layer. The polycrystalline H C A phase is equivalent in composition and structure to the mineral phase of bone 1 0 ' H A coating can prevent the fibrous tissue encapsulation of implants and reduce the potentially harmful release of metallic ion from the implant into the biological environment 1 2. Therefore, H A coatings on metallic substrate were applied to achieve better biocompatibility. Hydroxyapatite Caio(P0 4 )(OH)2 is chemically similar to the mineral component of bones and hard tissues in mammals. Hydroxyapatite crystallizes into hexagonal crystal structure and has a unit cell dimension a = 9.432 A and c = 6.881 A . The Ca/P ratio of H A is 10:6 and the density is 3.219 g/cm 3 [ 6 \ The mechanical properties of synthetic calcium phosphate vary widely as a function of their crystallinity and porosity. Limitations of H A application relate to its brittleness (Kic < 0.9 MPa-m ) and low strength (generally, compressive strength <300 M P a , tensile < 50 MPa) [ 1 3 ] . Therefore, H A coatings on metallic implants were often combined to achieve high biocompatibility and high reliability. Although most data for H A performance comes from the field of orthopedics, there is accumulating evidence that H A performs equally well in blood circulation environment 1 4. 4 1.3 Electrochemical Deposition of HA Coatings Different deposition technologies of H A have been investigated in the past including; plasma spray 1 5" 1 7, sol-gel 1 8 , pulse laser deposition 1 9 , sputtering 2 0, electrophoretic deposition 2 1, and electrochemical deposition 2 2" 2 4 . Electrochemical deposition (ECD) of calcium phosphate coatings at room (or near-room) temperature has been investigated since the 1990s. E C D is particularly attractive for coating internal surfaces of porous metallic substrates or highly complex surface. The resulting properties of H A can be controlled in E C D by regulating the electrochemical potential, current density, electrolyte concentration, and temperature of the process, as these conditions directly determine the deposition rate, morphology, and chemical composition of the H A coatings 2 5. In the cathodic electrochemical deposition method, cathodic reactions are used to generate OH" groups and increase p H at the electrode. Metal ions or complexes, which are stable in the bulk solution at low pH, are hydrolyzed by electro generated base at the electrode surface to form colloidal part icles 2 3 ' 2 6 . A l l applications of H A coatings reported so far are related to solid implants which do not undergo any deformation during the implantation procedure. Coronary stents are very different in this respect as there is a significant strain of stent during expansion, as large as 15% . Due to brittleness of ceramic, there exists a challenge for the E C D hydroxyapatite coating to survive under the stent expansion during the angioplasty procedure. While it is acceptable to micro-fracture the coating during stent expansion, it is not acceptable to separate the coating from stent during implantation. Debris of 5 separated coating have the potential to clog up blood supply causing fatal affect. Therefore, one of the crucial factors for E C D - H A fi lm processing is the preparation of the metallic substrate such that the E C D coating can achieve high adhesion strength, sufficient to maintain coating/stent integrity during implantation of the stent. Bio-resorption or dissolution is a major concern of in the biological environment. In terms of the H A coating, variables such as phase composition, crystallinity, Ca/P ratio, microstructure, thickness, porosity, and surface morphology are crucial. It is believed * 28 30 that amorphous coatings have a higher bio-resorption rate than crystalline coatings " . Thus, phase composition and crystallinity of E C D - H A coatings are important parameters for stability evaluation, and it is important to characterize these for the coatings on coronary artery stents. 1.4 Mot iva t ion and Focus of the Present Study There is a need for a novel process for non-polymeric, biocompatible coatings for cardiovascular stents. Such E C D process is developed in the present work for deposition of calcium phosphate coatings on stents. The E C D parameters, such as deposition temperature, current density, heat treatment and different solvent systems, were investigated in the past to create thick film coatings (>10 Lim). However, very few studies have applied these processes into implantable medical device or combine these finding to assess the mechanical behavior, and none was reported to concern H A coatings on cardiovascular stents. In the present study, we are exploring E C D process parameters, such as current density and time of deposition, to produce reliable thin film (<500nm) 6 calcium phosphate coatings at low temperatures (40-70°C) on cardiovascular stents. A substrate surface pre-treatment is employed prior to the electrochemical deposition to achieve high level of coating adhesion. The coatings phase composition and crystallinity are investigated and optimized through introduction of a post-deposition phase conversion process. Mechanical behavior of the coatings is characterized qualitatively through observation of the coatings behavior on expanded stents. H A was applied on coronary stents followed by a simulated stent implantation procedure to ensure high reliability of coating, i.e. no separation, delamination or visible damage. Finally, performance of the H A - E C D coated stents was evaluated in-vivo by the collaborating industry ( M I V Therapeutics, Vancouver, B.C. ) using porcine models. 7 2 L I T E R A T U R E REVIEW 2.1 Coronary Hear t Disease ( C H D ) Coronary heart disease (CHD) , also called coronary artery disease ( C A D ) , is the result of the atheromatous plaque accumulation, which composed of cholesterol, fatty compounds, and a blood-clotting material (fibrin), within the walls of the arteries that supply blood to the myocardium- the middle layer of the heart 3 1 ' 3 2 . Figure 2.1-1 illustrates accumulated atheromatous plaques on the inside of arteries. While the atheromatous plaques initially expand into the arteries walls, they w i l l finally expand into the lumen of the vessel, affecting the blood flow through the arteries. When atheromatous plaques obstruct less than 70 percent of the diameter of the vessel, symptoms of obstructive C A D are rarely observed. However, as the plaques obstruct more than 70 percent of the diameter of the vessel, symptoms of obstructive C A D w i l l starts to develop . A t this stage of the obstruction, the symptoms of ischemic heart disease are often first noted during increased workload of the heart. A s the obstruction develops, there may be near complete blockage of the lumen of the coronary artery, restricting the flow of oxygen rich blood to the myocardium. Individuals with this degree of coronary heart disease w i l l typically suffer from myocardial infarctions- heart attacks, and may have signs and symptoms of chronic coronary ischemia, including symptoms of angina at rest and flash pulmonary edema 1. 8 Figure 2.1-1. Fat and cholesterol accumulated on the inside of coronary arteries. Coronary artery disease is the leading cause of death in North America. In 2004 approximately 54% of all cardiovascular deaths were due to coronary artery disease1. Due to the varying severity levels of C A D , treatments for the condition range from minor lifestyle alterations to major invasive surgery. There are three main objectives in treating C A D : (1) to prevent the development of symptoms or to reverse atherosclerosis; (2) to relieve symptoms; and (3) to lower an individual's risk of suffering from a heart attack and sudden death. C A D treatments are done in the following three stages: (1) Initial Treatment: occurs immediately after an individual is diagnosed with C A D or is recognized as being at risk. This treatment stage is usually based on making key lifestyle changes, such as quitting smoking, eating healthier foods, and exercising. (2) Ongoing Treatment: continue monitoring an individual following initial treatment, such as monitoring of blood pressure, weight, and cholesterol levels. This is used to 9 determine i f the lifestyle changes have produced favorable results and to examine the patient's continued risk for developing C A D . (3) Coronary Artery Bypass Surgery/Angioplasty Procedures: when an individual is determined to be at high risk for a heart attack, two revascularization procedures are available: coronary artery bypass graft ( C A B G ) , which is also called a "bypass", and Percutaneous Transluminal Coronary Angioplasty ( P T C A ) . Since both therapies have established favorable results, the decisions are based primarily on several key factors, such as is the severity of the narrowing, number of arteries being affected, the location of the narrowing, and individuals' factors, such as age and general health. Percutaneous transluminal coronary angioplasty is a technique of opening a narrowed or closed vessel without having to resort a major bypass surgery. P T C A has become an increasingly used and successful treatment by cardiologists to battle the ravages of C A D . Figure 2.1-2 demonstrates the P T C A technique with a balloon catheter. The balloon catheter is carefully positioned to the blockage site under fluoroscopy by the interventionist, and then it is inflated to compress the plaque against the artery wall . The use of a coronary artery stent together with balloon catheter in P T C A has been significantly increased during the past 10 years. Figure 2.1-3 illustrates the use of a coronary artery stent during and after P T C A to provide support for the artery wall . In 1996, stents were used in 50% of the 20,000 P T C A procedures in United Kingdom 3 4 . 10 Figure 2.1-2. Schematic of percutaneous transluminal coronary angioplasty technique. A guide-wire is placed across the blocked section of the artery and a balloon is positioned beside the blockage. The balloon is then inflated compress the blockage against the artery wall . Figure 2.1-3. Schematic of percutaneous transluminal coronary angioplasty technique with the use of coronary artery stent. Coronary artery stent is used as mechanical scaffold to provide support to the vascular wall during and after the P T C A procedures. 11 2.2 Coronary Artery Stent Coronary artery stents are metallic implantable tubular medical devices used as mechanical scaffolds to provide support to the vascular wall during and after the revascularization procedures of coronary arties. Stent is commonly mounted on a balloon catheter as a single unit, with the use of fluoroscopic screening and a radiopaque marker, the unit is carefully positioned across the site of lesion. Inflation of the balloon catheter results in expansion and deployment of the stent circumferentially to the surface of the coronary artery. In 1964, Dotter et al were the first to describe a wire tubular implant as "stent" which was used in a non-surgical treatment for femoral arteries of animal 3 5 . In 1983, Dotter et al and Gragg et al implanted the first "stent-like device", a wire'spring of nitinol, into the canine coronary arteries with a catheter3 6. In 1985, Palmaz et al reported the introduction of coronary artery stent together with a non-deployed balloon in the site of lesion . Two years after, in 1987, Roussau et al tested a flexible, self expanding stainless steel stent. Forty-seven stents were implanted into 28 dogs, and 21 of these devices are implanted in the arteries. It was reported that 35% o f the animals exhibited partial or total thrombotic occlusion . It was in 1986 that the first implantation of coronary artery stent in human was performed and reported by Jacques P u e l 3 6 ' 3 1 . In 1991, Schatz announced the results from a multicenter study of 229 successful deliveries of intracoronary stent placements in 230 lesions in 213 patients 3 7. It was reported that the subacute thrombosis rate reached 14% and 40% of restenosis was determined in the following 6 months. The high level of thrombosis rate had convinced cardiologists that 12 stent, as a foreign object, when implanted wi l l cause high thrombogenicity. In the early 1990s, the Belgian-Netherlands Stent study ( B E N E S T E N T - 1 ) 3 8 and Stent Restenosis Study ( S T R E S S ) 3 9 were carried out to compare the conventional coronary balloon angioplasty with stent implantation. The results have shown beneficial effects of coronary stenting following balloon angioplasty, in comparison to angioplasty procedure alone. It was reported that among the 407 patients in the studies, the incidence of restenosis was significantly lower after stent implantation ( B E N E S T E N T 22%, STRESS 32%) when compared with after balloon angioplasty alone ( B E N E S T E N T 32%, S T R E S S 4 2 % ) 3 8 ' 3 9 . The beneficial effect of coronary stenting has been attributed to the larger 37 • acute lumen dimensions and to the prevention of constrictive remodeling . It was until 1994, that the first F D A approval of stent device was issued to Palmaz-Schatz stent. Many pioneer researches have underlined major problems with the use of stents. Subacute stent thrombosis is undoubtedly one of the problems despite the fact that very aggressive anticoagulation regiment was used in several studies. Restenosis, defined as the decrease of the vessel lumen diameter by at least 50% at the site of angioplasty procedure, is the other problem. Nevertheless, the combination of the technical advances and previous research data helped to define the ideal characters of today's stent as fo l lows : 3 6 ' 3 7 • Flexible • L o w unconstrained profile • Radio-opaque • Non-Thrombogenic 13 • Biocompatible • Reliable • High radial strength • Circumferential coverage • L o w surface area Alloys routinely used for the manufacture of stents include 316L stainless steel, nitinol, cobalt-chromium alloy, and tantalum. It is obvious that the mechanical properties of these alloy materials can dramatically influence the stent properties and stent design possibilities. In medical grade 316L stainless steel, L is used to indicate the low carbon content (0.03%). This alloy is composed of iron (60% to 65%), chromium (17% to 18%) and nickel (12% to 14%). Stainless steel provides good mechanical, chemical, and physical properties but its biocompatibility remains an issue. 316L stainless steel made stents often require plastic deformation b y balloon catheter to deploy in the artery. Nitinol or N i T i alloy consists of - 5 5 % Nickel and - 4 5 % titanium. Ni t inol is gaining popularity in the design of stent due to its shape memory capabilities. Under a mechanical load, nitinol can deform reversibly up to 10% of strain by the creation of a i n stress-induced phase transformation . Once the applied load is removed, the stress- induced phase becomes unstable and the material recovers to the original shape. However, concerns regarding nickel release from Nit inol have limited its applications ' . Tantalum is another option for materials in stent manufacture, though questions have been raised because of its exaggerated radiopacity 3 7 ' 4 0 . Cobalt-chromium alloys based stents are under development by leading stent manufacturing companies such as: Guidant 14 and Medtronic. Cobalt-chromium alloys allow thinner strut design on stent without compromising radial strength and radiopacity 3 7 ' 4 0 . Although coronary artery stenting results in a larger acute lumen and prevents elastic retraction of vessel, the presence of a metallic stent stimulates the restenosis mechanisms more significantly than simple balloon angioplasty owing to intimal hyperplasia, i.e. the biological response of an injured vessel. The complex restenosis mechanism following the stent implantation is due to the f o l l o w i n g 3 8 ' 4 1 : • The inner elastic membrane induced by injuries. • The secretion of mitotic substances and growth factors. • The prolonged and continuous stress created by the stent. • The chronic irritation targeted by the presence of a foreign object. In 2002, the worldwide market of coronary and peripheral intervention was about U.S . $5 bil l ion dollars and growing 5% per annual. O f all percutaneous coronary intervention, 75% involve stent implantation and as much as 50% for peripheral intervention, approximately 1,000,000 stents are being implanted each year worldwide 1 . Nowadays, multidisciplinary studies are being conducted to enhance the overall stent performance. The current stenting technology includes: coated stents, biodegradable stents, radioactive stents, and drug-eluting stent that release pharmaceutical agents locally. 15 2.2.1 Coated Stents Various coatings have been researched for coronary stents, including: diamond- like carbon 4 2 ' 4 3 , hydrogen-rich amorphous silicon carbide 4 4 ' 4 5 , amorphous titanium oxide 4 6 , and gold 4 7 . Some of these inert or tissue friendly coatings do not cause platelet accumulation and inflammation. Comparing to uncoated stents, these coated stents demonstrated a lower thrombogenicity in the biological environment and improved overall biocompatibility. In an in-vitro study, Gutensohn et al have reported that diamond like carbon can not only reduce thrombogenicity but also reduces the release of metallic ions from the stainless steel to the surrounding tissue 4 3. The inertness of gold was investigated by Kastrati et al with the expectation of increasing the stent biocompatibility and radiopacity. However, a randomized trial has suggested the gold coated stents exhibit an increased possibility of restenosis when compared with stainless steel stents4 7. It was further reported that the undesired result may be related to a coating defect rather than the gold itself. This demonstrates the importance of the coating processing along with biocompatible concerns, which together ultimately determines the outcome of the procedure. Heparin coating on stent is another approach to reduce thrombogenicity. Heparin is a complex organic acid found in lung and liver tissue, having the ability to prevent blood clotting. Although heparin can not break down already formed clot, it allows the AO body to dissolve away the clot . In animal model, heparin coatings have been proven to reduce thrombus formation, however, long term clinical data is still uncertain. One of the 16 commercialized heparin stents, B X Velocity by Johnson & Johnson, immobilizes heparin on the stent surface and remains free to inhibit thrombus formation 4 9 . 2.2.2 Biodegradable Stents One of the primary reasons to use coronary artery stents is for the scaffolding effect. While this is observed to be beneficial in the short term, the presence of a metallic stent may trigger chronic inflammation and in turn stimulate restenosis. To avoid the latter problem, a biodegradable stent has been proposed. A biodegradable stent exhibits similar mechanical functionality of a metallic stent but w i l l then slowly degrade as the artery become stable. Stackle et al were the first to develop biodegradable stents with poly-L-lactase 5 0. In 2000, Tamai et al have reported the results of the first 6-month clinical trial of biodegradable poly-l-lactic coronary stents in humans. In these results, no thrombosis or death was observed for the first 30 days, and an acceptable restenosis rate was reported to be 10.6% 5 1 . 2.2.3 Radioactive Stents It was demonstrated that a low dose of radiation can effectively inhibit or decelerate the proliferation of vegetative c e l l s 3 7 , 4 9 ' 5 2 . It is believed that this phenomenon can lead to a reduction of intimal hyperplasia and as a result decreases the chance of restenosis. Herlerin et al have reported the use of radioactive stent in an animal study 5 2. The radioactive stents used are made of steel bombarded with Co , M g , or Fe ions in a cyclotron, and thus emitted y and p radiation. These stents can emit various radiations 17 doses between 15-23 uCurie. The animal study has shown that while both the low and high doses radiation reduces intimal hyperplasia, the intermediate dose caused an increase of hyperplasia . This demonstrates the complexity between radiation and vessel response. While the therapeutic and toxic limits still have not been determined accurately, the rate of restenosis at the edges of radiated area remains a major problem 4 9 . In general, there still exist unsolved and unknown problem with the use of radiation in form of short therapy and radioactive stents. 2.2.4 Drug Eluting Stents Drug eluting stents (DES) that can efficiently deliver biologically active agents locally to the vascular wall have been under intensive research in recent years. Early results with drug eluting stents were often unsatisfactory due to uncertain require dosage, uncertain delivery duration, poor drug efficacy, and proinflammatory polymeric coating 3 7. More recent studies on D E S with the use of anti-proliferation drugs have generated excellent results. Sirolimus and Paclitaxel are the two most commonly used anti-proliferation drugs. Sirolimus is a macrolide antibiotic that possesses immunosuppressant activity, it is also known to be an inhibitor of smooth muscle cells (SMC) proliferation and migration 5 3 . Paclitaxel is an approved drug used as cancer chemotherapeutic agents, it inhibits cell proliferation and migration by assembling the tubulin dimmers into the non-functional microtubules, causes alteration of the cell cytoskeleton structure5 4. Other anti-proliferation drugs under investigation includes: everolimus, tacrolimus, estradiol, dexamethasone, and angiopeptin. 18 Two of the commercialized polymer-based D E S are C Y P H E R ™ stent with sirolimus by Cordis® and E X P R E S S ™ stent with paclitaxel by Boston Scientific®. Both of these D E S use polymer as a vehicle to carry the drug eluting into surrounding tissue through diffusion and convection 5 5. While concerns have been raised with polymer coating, non-polymeric stents A C H I E V E ™ coated with paclitaxel by Cook® is under investigation. However, clinical studies have been inconclusive. The question between the potential long term degradation and inflammatory effect of polymer coating and control release rate of non-polymeric coating are still in debate. Although there can be no doubt that the development and implementation of drug-eluting stent in past clinical studies has been a major milestone that w i l l evolve the approach to the management of coronary artery disease, yet, it is also certain that more data are required to fully understand the scope of its benefits and to identify its limitations. In spite of the different drugs being investigated and their different therapeutic effects, the technique of coating, and drug delivery platform are still the key component leading to the success of the device. 2.3 Bioceramics The American Society for Testing and Materials ( A S T M ) defines ceramics as "...essentially inorganic, nonmetallic substances..." 5 6 . Ceramic bonding is generally a hybrid of ionic and covalent bonds. The strong bonding forces of ceramic materials result in properties such as high elastic modulus and hardness, high melting points, low thermal expansion, and good chemical resistance6. However, the bonding nature and minimum number of slip system of ceramics make it difficult to shear or deform 19 plastically, therefore lowering the fracture toughness. Ceramic materials are hard and brittle. Unlike polymers and metals, ceramics are non-ductile and are very susceptible to notches or microcracks. It is not easy to process flawless ceramics, pores or micro-cracks often cause stress concentration to initiate cracks. Because of this, it is very difficult to precisely determine the tensile strength of ceramics, moreover, it is the reason why ceramics have low tensile strength compared to its high compressive strength. Modern advance techniques for ceramic processing have led them to become advance engineering materials. More recently, the biocompatibility and bio-inertness of ceramics have been realized, and these "bioceramics" are being widely used in orthopedics and dental industries. Bioceramics can be manufactured in various forms such as: micro-spheres, thin film coatings, porous network blocks, composites with a polymer component, and well polished surfaces . Bioceramics used within the body can be categorized into three classifications: bio-inert, bio-resorbable and bio-active 6. Bio-inert ceramics are inert in the physiological processes, meaning that they maintain their physically and mechanical properties unchanged, and have almost no influence on the surrounding living tissue. Examples of bio-inert ceramics are alumina, zirconia, silicone nitride, and carbon. Most of these bioceramics are used in structural support devices such as: bone plates, bone spacers, bone screws, and femoral heads 5 7. Alumina has been used in orthopedics and dental surgery for almost 20 years, it properties include high hardness, inertness, low CO coefficient of friction and wear resistance . These have made alumina as the ideal candidate for joint replacement. Medical grade zirconia has been developed for use in 20 total joint prostheses because of its high fracture toughness and tensile strength. Such improved properties of zirconia have enable manufacture of femoral head into smaller diameter than present generation of alumina. Pyrolytic carbon is commonly used in coating of artificial heart valves for the last 30 years 5 9. Properties that make this material suitable for this application include good strength, wear resistance, durability, and most importantly, thromboresistance. Bio-resorbable ceramics, as its name implies, w i l l dissolve or degrade upon implantation into the body. The degradation rate of bio-resorbable ceramics varies among different materials. Most of the bio-resorbable ceramics are in forms of calcium phosphate. Due to the resorbable nature of most calcium phosphate, they have been often used as potential bone defect fillers. In such application, the calcium phosphate filler would f i l l the void and gradually dissolve away, being replaced by host tissues. Two critical issues with the development of bio-resorbable ceramics are: maintenance of stability and strength during the degradation period, and matching resorption rate to the natural repair rates of the body tissues. Tricalcium phosphate is one example; it has been used as temporary bone substitute6 0, where it has to maintain sufficient strength at first to allow bone ingrowth, then eventually degrade and be replaced by endogenous bone. Other uses of bio-resorbable ceramics include drug deliver devices and ocular implants. Tricalcium phosphate cysteine composites loaded with erythromycin or penicillin were developed by Morris et al. for the treatment of bone infection. It was reported that TCP/cysteine composite releases the antibiotics over a period of 3 weeks at the site of 21 infection. The authors suggested that antibiotics released from T C P amino acid composites effectively utilized in the treatment of bone infection 6 1 . Bio-active ceramics are surface reactive and bond to adjacent tissue some time after implantation 1 2. In many cases, the interfacial bond strength is equivalent to or greater than the cohesive strength of the implant material. Bioactive glass and hydroxyapatite are examples of bio-active ceramics. Unlike the other bio-resorbable calcium phosphates, hydroxyapatite does not break down in physiological conditions. Hydroxyapatite is thermodynamically stable at physiological p H and actively takes part in bone bonding, forming strong chemical bonds with surrounding bone. Although the mechanical properties of bio-active ceramics have been found to be unsuitable for load- bearing applications such as in orthopedics, they are commonly used as surface coating on metallic implants to provide bonding with tissue, while the metallic component bears the l o a d 1 7 ' 2 6 . 2.3.1 Calcium Phosphate Bioceramics Calcium phosphate (CaP) has been used in medicine and dentistry for more than 20 years and it has.been used as artificial bone since 1970s . It has been synthesized and used for manufacturing various forms of implants and also coating for implan t s 7 " 9 ' 1 7 ' 2 6 ' 6 3 . Depending on the Ca/P ratio, water content, impurities, and temperature, calcium phosphate can be crystallized into different forms 6. There are several calcium phosphate ceramics that are considered biocompatible. O f these, most are bio-resorbable and wi l l dissolve when exposed to physiological environments. Some of these materials include: 22 amorphous calcium phosphate ACP, dicalcium phosphate DCP ( C a H P C ^ H i O ) , octacalcium phosphate O C P (Ca8H2(P04)65H20), tetracalcium phosphate TTCP (Ca4P209), alpha-tricalcium phosphate a-TCP (Ca3(P04)2), beta-tricalcium phosphate (3- TCP (Ca3(P04)2), and hydroxyapatite H A (Cai 0(PO 4)6(OH) 2). These forms of CaP are summarized in Table 2.3-1 6 4 . Table 2.3-1. Various calcium phosphate bioceramics Chemical Name Abbreviation Chemical Formula Ca/P Ratio Amorphous calcium phosphate ACP - - Dicalcium phosphate DCP CaHP0 4 1.00 Dicalcium phosphate dihydrate DCPD CaHP0 4 2H 2 0 1.00 Octacalcium phosphate OCP Ca 8 H 2 (P0 4 ) 6 5H 2 0 1.33 Alpha-tricalcium phosphate a-TCP a-Ca 3 (P0 4 ) 2 1.50 Beta-tricalcium phosphate P-TCP B-Ca 3(P0 4) 2 1.50 Hydroxyapatite H A Ca 1 0 (PO 4 ) 6 (OH) 2 1.67 Tetracalcium phosphate TTCP C a 4 P 2 0 9 2.00 2.3.1.1 Bioresorption and Biodegradation The dissolution products of CaP bioceramics can be readily assimilated by the human body. Bioresorption and biodegradation is generally controlled by Physiochemical dissolution- degradation depending on the local p H and the solubility of the biomaterial, Physical disintegration- degradation due to disintegration into small particles, and Biological factors- degradation cause by biological responses leading to local pH decrease, such as inflammation. A l l calcium phosphates degrade differently, the 23 degree of biodegradation in calcium phosphate ceramics is controlled by the aforementioned three factors, and also dependent on their properties such as surface area, density, porosity, composition, Ca/P ratio, crystal structure, and crys ta l l in i ty 6 5 ' 6 6 . While various variables w i l l have an effect on the biodegradation of calcium phosphate, the general order of solubility near-neutral p H environment is as follows (from highest to lowest): A C P > D C P > T T C P > O C P > a -TCP >p-TCP > H A Stability of various forms of CaP as a function of p H is illustrated in Figure 2.3-1 6 7 . A s seen in Figure 2.3-1 hydroxyapatite is relatively insoluble compare to other calcium phosphate phases, it is the only stable phase above p H 4.2. Below p H 4.2, dicalcium phosphate dihydrate (DCPD) is the stable phase. It is often observed that unstable phases of calcium phosphate wi l l dissolve and re-precipitate into stable phase at a given p H . The p H of normal physiological environment is 7.2, however, this may decrease to as low as 5.5 in the region of tissue injuries or inflammation, and slowly return to 7.2 over time. 24 0.0 30 40 50 ^ 60 70 80 90 pH Figure 2.3-1. Logarithm of the product of calcium and phosphate concentrations plotted against p H values of solution saturated with respect to various calcium phosphate phases in the ternary system Ca(OH)2-H 3 P04-H 2 0. Calculated for 37°C. © American Chemical Society, 1980, adapted by permission1 2.3.1.2 Mechanica l Properties The biomedical applications of calcium phosphates depend greatly on their mechanical properties; however, the poor mechanical properties of ceramic often limit its application. Table 2.3-2 summarizes the mechanical and physical properties of calcium ' Reprinted from Journal of Cystallinization of Calcium Phosphates, 102, P Koutsoukos, Z Amjad, M B Tomson, G H Nancollas, Crystallization of Calcium Phosphates. A Constant Composition Study, 1553 - 1557, Copyright (1980), with permission from American Chemical Society. 25 phosphates6. The wide range of variations in properties is due to the variation in structure, phase composition and manufacturing process. Tensile and compressive strength and fatigue resistance of H A (like most ceramics) are highly dependent on the total volume of porosity. For example, Jarcho et al reported compressive strength of H A as high as 917 M P a 6 8 . The H A specimen has been heat treated for 1 hour at 1100°C, while maintaining very fine grain size of 0.3 urn. In another study, a low compressive strength value o f 138 M P a was reported by Rao et al, in which the H A specimen was heat treated at 900°C for 0.5 hour 6 9 , and likely produced high porosity fraction in the material. Furthermore, the presence of porosity in form of either micropores (<lum) or macropores (>10um) can affect both the compressive and tensile strength o f the material 1 0 . Table 2.3-2. Mechanical and Physical Properties of Calcium Phosphates P r o p e r t y V a l u e Elastic Modulus (GPa) 4 . 0 - 1 1 7 Compressive Strength (MPa) 7 0 - 3 0 0 Bending Strength (MPa) 147 Hardness (Vickers, GPa) 1 .1-6.0 Poisson's Ratio 0.27 Bulk Density (g/cm J) 1.90-3.27 Theoretical Density (g/cm 3) 3.16 Facture Strength (MPa) 12 .8 -60 .4 26 2.3.2 Hydroxyapati te Bioceramics Hydroxyapatite (HA) is chemically similar to the mineral component of bones and teeth, it is the most used phase among the various calcium phosphate bioceramic. H A is classified as the bio-active material, it exhibits good biocompatibility, bioactivity, and osteoconductivity 9' 6 4 . It has been intensely researched and used in orthopedic, dental, and maxillofacial applications. Hydroxyapatite crystallizes into hexagonal crystal structure and has a unit cell dimension a = 9.432 A and c = 6.88lA. The Ca/P ratio of H A is 10:6 and the theoretical density is 3.219 g/cm . The mechanical properties of hydroxyapatite vary widely as many other calcium phosphates. Depending of the final heat treatment conditions, Ca/P ratio, and the presence of water and impurities, calcium phosphate can transform into H A or P -TCP. In many final products both H A and p-TCP phase coexist. A bioceramic consisted of a mixture of H A and P -TCP was first described 7ft as biphasic calcium phosphate (BCP) by Nery et al in 1986 . It was further researched and reported that the bioactivity of B C P maybe controlled by manipulating the H A / p - T C P ratios. Hydroxyapatite exhibits excellent biocompatibility, it has the ability to integrate into bone structure and support bone ingrowth, and can form direct chemical bond with hard t i ssue 7 ' 8 ' 6 4 . It was reported that upon 4 to 8 weeks implantation, new lamellar bone and cancellous bone forms into the pore of hydroxyapatite implant . Hydroxyapatite can be employed as bone substitute in forms of powders, porous blocks, or beads. Bone substitute or bone filler was most often used in the case of bone cancer and bone augmentation, where a large section of bone must be removed or reconstructed. Because 27 of the osteoconductive nature, H A bone filler can provide a scaffold and induce rapid filling of natural bone, reducing the healing time. More recently, the effect of nanocrystals of H A on microvascular endothelial cell was studied. Pezzatini et al have exposed microvascular endothelial cells to stoichiometric H A nanocrystals and reported H A sustained endothelial survival without any cytotoxic effect 1 4. Endothelial cell is a layer that lines the cavities of the heart and the blood vessels. The reaction of vascular endothelium cell to implanted biomaterial is of great importance because they interact in each step of tissue integration. It was concluded that H A nanocrystals exhibit high biocompatibility for microvascular endothelium, and do not acquire a proinflammatory or thrombogenic phenotype. Furthermore, endothelium was observed to be functioning toward angiogenesis - the formation of new blood vessels. 2.3.2.1 Thermal Stability Hydroxyapatite is a hydrated calcium phosphate material. The stability of calcium phosphate depends on the temperature of preparation and the partial pressure of water. H A was found to be the stable phase up to 1360°C at 500mmHg partial pressure of water. And , in the absence of water content, p -TCP is found to be stable phase 7 1. A t dry environment and at elevated temperature above 900°C, H A w i l l decompose and form other calcium phosphate as follows: Caio(P0 4)6(OH) 2 <=> 2 p- C a 3 ( P 0 4 ) 2 + C a ^ O o + H 2 0 Ca io (P0 4 ) 6 (OH) 2 <=> 3 p- C a 3 ( P 0 4 ) 2 + CaO + H 2 0 28 A t dry environment and at temperature above 1350°C, P -TCP w i l l form. Both H A and P- T C P are very tissue compatible and o s t e o i n d u c t i v e 6 , 7 1 . Therefore, heat treating H A under vacuum condition can induce decomposition at lower temperature, promoting the formation of water vapor. Whereas, heat treating in a high water content environment can counteract this effect and delay decomposition. Thus, it is important to control the atmospheric water content in order to prepare the desired final calcium phosphate phase. 2.3.2.2 Mechanica l Properties Hasimah et al have conducted a study on the effect of heat treatment temperature on the mechanical properties of hydroxyapatite 7 2, it was reported that upon heat treating H A from 1100°C to 1300°C, the compressive fracture strength increases from 12.87 M P a to 60.36 M P a and the porosity decreased from 37.7% to 5.5%. It was concluded that the increase in fracture strength was due to the increased bulk density as pore has a tendency to initiate crack. Similar study was also performed by Muralithran et a l 7 3 . It was observed that when H A was heat treated from 1000°C to 1450°C, the relative density increased from 77.3% to 98.5% and the average grain size was 2.03 um and 12.26 um respectively. The authors suggested a definite positive correlation between hardness and grain size, however, the hardness was found to have decreased when a critical grain size is reached. 29 2.4 Biomater ia l Coatings Manufacturing of medical devices often incorporates a surface coating process. The design of medical devices is generally based on the bulk properties of material, it is often difficult to have a material that is both mechanically and biologically suitable. Since the biomaterial-biological medium interface is complex and usually involves reactions such as adsorption of water, protein-blood clot, inflammation, ion exchange, cell adhesion, corrosion, and decomposition, the surface properties o f medical device plays a key role in determining its success. A biomaterial coating can tailor the surface characteristic of a medical device for a specific application without detrimentally affect the bulk properties of the biomaterial. Table 2.4-1 summarizes the purposes of biomaterial coatings. The ultimate objective of biomaterial coatings on medical devices is to combine the mechanical advantages of metallic implant and the biocompatible properties of coating to achieve the best desired biocompatible environment. Table 2.4-1. Purposes of biomaterial coatings. Wettability Permeability Bio-stability Chemical inertness Adhesion Biocompatibility Drug Delivery Electrical Characteristics Optical Properties Frictional Properties Biomaterial coatings can be categorized into two classes: passive surface coatings and active surface coatings. In general, passive coatings are used to alter selected properties of the implant surface without delivery of therapeutic or other agents. 30 Whereas, active surface coatings are aimed to deliver bioactive compounds or drugs to influence the biological response in order to achieve the desired function. In coronary stent applications, passive coating has often been applied as a thromboresistant coating; phosphorylcholine (PC) is one of the examples. In a six month implantation study, Grenadier et al have reported stents coated with phosphorylcholine appears to be safe and efficacious in the treatment of complex coronary lesions 7 4. Drug eluting coatings are the one of the latest active coating developments in the medical industry, where formulated therapeutic agents are embedded into the coating reservoirs or matrix to enable site specific delivery. The drugs are gradually and locally released from the coating and simultaneously being absorbed by the surrounding tissues 4 0. 2.4.1 Hydroxyapati te Coatings Due to the highly bioactive and biocompatible nature of H A , a medical device coated with H A is less likely to be recognized as foreign object by the body. Clinical data have shown that i f an implant is coated with H A , it is possible to allow for positive material connections to establish between inorganic material and vital bone, and hence achieve long term osteointegration. In the case of osseous applications, hydroxyapatite coating can prevent fibrous tissue encapsulation on metallic implant, at the same time improves the integration of bone into the implant, and thus provides a strong bond Q between them . In 1991, Furlong et al have performed a histological section of H A coated stem, which shown good osteointegration and formation of new vital bone. There was no evidence of an inflammatory reaction or of fibrous tissue formation 7 5 . Further 31 study on H A coated porous metallic implant demonstrated a clear acceleration of new bone ingrowth 7 5 . It is known that various metal ions can affect the functionality of osteoblast cell , and H A coating was also observed to be able reduce the release of metallic ions from the implant into the physiological surrounding. Sousa et al have reported that a f i lm of 50pm H A coating can act as an effective barrier to metal ion release . In their study, T i e A U V alloy was coated with 50pm H A by plasma spraying. N o titanium, aluminum or vanadium was detected in by electrothermal atomic absorption spectroscopy. Similar study with 305 stainless steel coated with 50pm H A also found no release of metallic ion 77 . This barrier effect o f H A was claimed to be a result of metal phosphate formation or incorporation of metal ions in the H A structure. 2.4.2 Processing of Hydroxyapati te Coatings In the past, many different H A deposition methods have been reported, for example: plasma spray 1 5" 1 7, sol-gel 1 8 , pulse laser deposition 1 9, chemical vapor deposition, sputtering 2 0, electrophoretic deposition 2 1, and electrochemical deposit ion 2 2 ' 2 4 . Among these methods, pulse laser deposition, sputtering, and chemical vapor deposition all involve a two step process. The first step involves synthesizing the desired coating material in bulk form, and in second step a pellet of coating material is irradiated by a high power source. Therefore these methods consume high amount of energy, and often require an ultra-high vacuum system, increasing the capital cost intensively. Furthermore, the coating composition frequently varies from the target material due to 32 the different sputtering speed. These methods are also not suitable for deposition on complex surface. Most of al l , the high temperature of deposition may lead to detrimental effects on the of target substrate. Plasma spray deposition of H A has become the most popular and commercial method for coating H A on orthopedic devices, however, technical issues still remain. One of the issues is the formation of other calcium phosphate phases other than H A resulting from the extremely high temperatures (> 10,000 K ) used in the plasma spray process. Ji et al have reported a micro structural study of plasma sprayed H A on titanium alloy and found that other than crystalline H A , amorphous calcium phosphate with Ca/P ratios of 0.6 - 1 . 0 are present, also a calcium titanate phase was detected at the coating- 78 substrate interface . Another issue is the poor adhesion of H A by plasma spray. Radin et al have found delamination of H A coating from titanium substrate in an in-vitro stability study 7 9. Table 2.4-2 summarizes the general advantages and disadvantages of the techniques used for the deposition of hydroxyapatite. 33 Table 2.4-2. Summary of techniques used for deposition of hydroxyapatite coatings Technique Thickness Range Advantages Disadvantages Plasma Spray 10 - lOOOum • Rapid deposition rate • Dense under well control • spray very high melting point materials • Wide range of applications • Line of sight process • High temperature induced • High cost • May decrease fatigue life of substrate • High capital cost Sol Gel <l[im • High purity • Low cost process • Low heat treatment process temp. ( 2 0 0 - 6 0 0 ° C ) • Expensive raw materials • May require control atmosphere Pulsed Laser Deposition 0.5-2mm • Cost effective • Rapid deposition rate • Uniform coating thickness on flat substrate • Line of sight process • Cannot coat complex or porous substrate • High capital cost Sputtering 0.02-lmm • Coating same composition as the source material • Rapid deposition rate • Uniform coating thickness on flat substrate • Line of sight process • Cannot coat complex or porous substrate Electro-deposition 0.1-50um • Uniform coating thickness • Rapid deposition rate • Can coat complex substrate " Low cost • Sometime require high temperature heat treatment • Difficult to produce thick crack-free coatings 2.5 Electrochemical Deposition The electrochemical deposition (ECD) is done by passing an electric current between two electrodes separated by an electrolyte, and the deposition process takes place at the electrode-electrolyte interface. The electrochemical synthesis can be a reduction or an oxidation reaction. B y manipulating the applied potential, the E C D reaction can be continuously varied. There are a number of features that differentiate E C D from other deposition technologies: (1) The reaction of electrochemical process takes place within the electric double layer of the electrode, which has a high potential gradient. Electric double layer is the layer which forms at the solid/liquid interfaces as a 34 result of a net charge on the solid surface (usually negative) causing a localized layer of neutralizing counter-ions (usually positive) from the solution phase to form near the solid on surface . (2) The deposition process occurs in a solid-liquid environment, which facilitates the growth of conformal coating on complex shaped substrates. (3) Process temperature is often low (<70°C) since it is limited by the boiling point of applicable electrolyte. (4) The electrolyte bath composition can be varied to control the coating composition. (5) Equipments are simple, inexpensive, and readily available. However, there are some disadvantages in the E C D process. The deposition product is often poorly ordered, making structural characterization difficult and sometimes contain amorphous 80 impurties . Furthermore, the deposition process can only be done with a conductive substrate. Since coronary stent is a three-dimension mesh tube, the ability of E C D to deposit conformal films on complex non-planar devices is advantageous. Coronary stents can be coated relatively quickly (in minutes) and at relatively low (40°C) temperature. Capital cost of deposition instruments are relatively low compared to other technologies. F i l m thickness can be tailored over a wide range from ~0.1 um up to millimeter range while still maintaining its usefulness for a desired appl icat ion 2 3 ' 2 6 . The thickness and chemical composition of the E C D coatings can be well controlled through adequate selection of the process parameters, such as substrate composition and electrolyte composition, electrolyte concentration, p H , and temperature, and mode of applied potential (constant or periodic). The two most important parameters in determining the course of E C D 35 reactions are current density and applied potential, either one of these parameters can be controlled as a function of time to control the rate of coating deposition. 2.5.1 Electrochemical Deposition of Hydroxyapatite Coatings In E C D , metal ions or complexes are hydrolyzed to form deposits on cathodic 23 25 26 substrates ' ' . In the electrochemical deposition process, the p H in the bulk solution is relatively low (typically 4.0<pH<6.0). The hydrolysis reaction and various cathodic reactions wi l l consume water and produce OH", resulting in an increase of p H near the cathode. In general, acidic solutions containing calcium and phosphate ions were used for the electrolyte for E C D - H A . Redepenning et al have reported the use of an aqueous solution saturated with Ca(H 2P04)2 as an eletrolyte 8 1. Shirkhanzadeh prepared an electrolyte for bioactive calcium phosphate coating by dissolution of hydroxyapatite into N a C l solution and further adjusted the p H by H Q 8 2 . A solution saturated with CaHP04-2HiO and added K N O 3 and N a H C 0 3 was prepared as electrolyte by Royer el 83 at . Electrochemical deposition of hydroxyapatite ( H A ) has been conducted in the mixed aqueous solution of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 2 3 ' 8 4 ' 8 5 . According to a study by Shirkhanzadeh et al, the p H value within the diffusion layer of the cathode during E C D wi l l be increased by the following two reactions: 2 H 2 0 + 2e- -» H 2 + 2 0 H " (1) 2 H + + 2e" H 2 (2) 36 The concentration of phosphate HPO4 " is determined by the following reactions: H+ + P 0 4 3 " o H P 0 4 2 " (3) H P 0 4 2 " + OH" H 2 0 + P 0 4 3 " (4) With the extra OH" being produced, solubility limit of H A is reached (as illustrated in Figure 2.3-1) and consequently H A is deposited on the cathodic surface by the following reaction: 10Ca 2 + + 6 P 0 4 3 " + 2 0 H " C a 1 0 ( P O 4 ) 6 ( O H ) 2 (5) Although this has been accepted as the general reaction sequence leading to H A deposition, details of cathodic reactions have been studied via cathodic polarization study to suggest further process controls. In a cathodic reactions study done by Yen et al, hydroxyapatite was coated via electrochemical method on pure titanium 2 3 . Yen et al reported that the cathodic reaction changes for different applied voltage: (1) Applied Voltage at [-0.1 to-0.3 V] 0 2 + 2 H 2 0 + 4e" -> 4 0 H " H 2 P 0 4 " + H 2 0 + 2e" -» H 2 P 0 3 " + 2 0 H " (2) Applied Voltage at [-0.3 to -1. 1 V] 2 H + + 2e" -» H 2 H 2 P 0 4 " + e" H P 0 4 2 " + lA H 2 (3) Applied Voltage at [-1.1 to-1.5 V] 2 H P 0 4 2 " + 2e" -» 2 P 0 4 3 " + H 2 (4) Applied Voltage at [-1.5 to -3.0 V] 2 H 2 0 + 2e"-> H 2 + 2 0 H " 37 It was found that at an applied voltage of-0.7 V , C a ( H P 0 3 H ) 2 - H 2 0 and C a ( H 2 P 0 4 ) 2 H 2 0 were formed. A t -1.25V, C a H P 0 4 - 2 H 2 0 were found to be the major phase and H A were found to be the minor phase. And , at -1.55V, H A became the major phase of 23 2 deposition . A purer H A was claimed at a lower deposition current density (2mA/cm ) after being annealed at 300°C for 4 hours. Kuo et al have also performed the cathodic polarization test, and suggested three stages of the reaction 2 6: (1) Reduction of oxygen at [-0.4 V] 0 2 + 2 H 2 0 + 4e" -¥ 4 0 H " (2) Reduction of H2P04' and 2HP042' at [-0.4 to -1.6 V] 2 H P 0 4 2 - + 2e" -» 2 P 0 4 3 " + H 2 2 H 2 P 0 4 " + 2e" -» 2 H P 0 4 2 " + H 2 (3) Reduction of water at [-1.6 to -3.0 V] 2 H 2 0 + 2e"-* H 2 + 2 0 H " It can be seen that both Y e n et al and Kuo et al have suggested very similar cathodic reactions. Figure 2.5-1 shows the cathodic polarization curve of titanium substrate in the electrolyte compose of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 . It was found that the greater amount of hydroxyls produced from reduction of water at voltage -1.6 to 3.0 V promotes the formation of a purer H A 2 6 . 38 "I'TITJ 1 r""l"i"llTI| 1 l - T T T T I T J 1 r ""I - IT IT 11 1 r ' T T I ' T T T J 1 | - T T T T I T | 1 f T ' T T T T F I ' ' 1 , 1 ' • • • » t l l l t I I I I t i l l . • 1 1 I 1 111 1 I I I • F i l l • • • • t i ••! • . • i • • 1.1 10"7 10s 1x10"5 1x10"4 10"3 10"2 10"1 I/Area (A/cm2) Figure 2.5-1. Cathodic polarization curve of T i substrate in a Ca(NC>3)2 4H2O and N H 4 . H 2 P 0 4 electrolyte. © Journal of Materials Science and Engineering C, 2002, adapted by permission2. The electrochemical deposition of H A has also been conducted in an electrolyte composed of calcium acetate, acetic acid, and sodium phosphate 8 4. Manso et al have subsequently dried the coated sample at 100°C and 900°C, X-ray diffraction ( X R D ) of the samples have shown that the 900°C dried sample exhibits H A phase with higher crystallinity. Surface force microscopy (SFM) was used to study the surface morphology of the coating. The analyses have proven that with the application of higher voltage a more compact film was produced, reducing the root mean square (rms) roughness from 90 nm to 20 n m 8 4 . 2 Reprinted from Journal of Materials Science and Engineering C, 20, Kuo M . C . , Yen S.K., The process of electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature, 153 - 160, Copyright (2002), with permission from Elsevier. 39 2.5.2 Microstructure and Phase of ECD Calcium Phosphate Coatings A s aforementioned, current density is one of the important parameters in controlling the rate of electrochemical deposition and the phase deposition. Y e n et al have performed a series of experiments at different current densities to investigate the 23 difference in phases and microstructures of E C D - C a P . Electrolyte composed of Ca(N03)2-4H20 and N H 4 H 2 P O 4 was used for deposition of CaP on pure titanium substrate, and current densities were selected as 0.3, 1.0, 2.0, and 3.0 m A / c m 2 for a deposition time of 1000s. It was found that at 0.3 m A / c m 2 , C a ( H P 0 3 H ) 2 - H 2 0 monocalcium phosphate monohydrate ( M C P M ) was the main phase and revealed a flake dendrite structure. A major D C P D and minor H A phase was found at 1 m A / c m 2 , and a fine loose plate-like structure was observed. A t 2 mA/cm , a dense plate-like structure was found, and H A became the major phase. A t 3 m A / c m , H A was again found to be the major phase, but "volcano-like" microstructures were observed. It was claimed that this is due to the "bubble-effect" from the increased current density and as a result generated higher production rate of hydrogen bubbles . The "bubble-effects" were also experienced by others electrochemical deposition studies when high current or voltage was applied to the system 8 6" 8 8 . Among these studies, "bubble effect" was generally observed to have a negative impact on coating surface coverage and coating structure. Optimization was attempted by Huo et al, in which the substrate was rotated at high speed, and thus the centrifugal force could clear away hydrogen bubbles. A study of electrochemical deposition of calcium phosphate in modified simulated body fluid (SBF) was done by Peng et a l 8 9 . S B F is an acellular fluid that has 40 inorganic ion concentrations similar to those of human extracellular fluid. Instead of applying a constant potential, a periodic pulsed potential was used for the deposition of calcium phosphate in S B F for duration of 30 minutes. The growth of coating was observed from 8 to 20 min of deposition. Sparsely distributed calcium phosphate deposits (nuclei) were found at 8 min, and the coalescence of these nuclei after 10 min of deposition has increased the surface coverage. A t 20 min of deposition, coalescence continues and an interconnected porous coating was observed. The pore size was in the range of a few nanometers to 1 u m 8 9 . X R D analysis has shown that the coating consists of O C P and amorphous H A phases. Similar to the work done by Yen et al, Huang et al have conducted an electrochemical deposition study with an electrolyte composed o f the same reagents. However, the applied voltage ranged from 1.0 to 10V and the time of deposition varied from 1 to 3 hours. Furthermore the coated specimens were subsequently treated with a hydrothermal process at various p H and temperatures. It was found that coating prepared at 60°C, 2.0 V , and 2.5 hours, contain essentially pure D C P D , and the coating morphology appeared as plate-like crystals with an estimated thickness of approximately 50 u m 9 0 . After post-hydrothermal treatment at 180°C, the coatings were found to consist of both needle-like and plate-like crystals. X R D results have demonstrated that the needle-like crystals corresponded to H A phase, and the plate-like crystal were D C P D phase. After 200°C hydrothermal treatment, the coating reveals an interlocking network of non-oriented needle-like crystals and it was proven by the authors that these crystals consist entirely of H A 9 0 . 41 Kuo et al have conducted electrochemical deposition in an electrolyte composed of Ca(NC>3)2 4H2O and NH4H2PO4. Pure titanium was used as substrate and coating was deposited at current density 1 - 2 0 m A for duration of 5 - 40 minutes at room temperature. X R D analyses of the coatings are shown in Figure 2.5-2 2 6 . With a current density of 1 m A / c m 2 , D C P D was observed to be the major phase and H A was the minor phase. The same was observed with current density 5 m A / c m 2 , but with an increased peak intensity at 26 - 25.88° and 29 = 31.77°. A t current density above 10 m A / c m 2 , D C P D peaks were not found and H A became the main phase of deposition. The authors have suggested that only at current density above 10 m A / c m 2 there w i l l be high enough concentration of OH" to convert HPO4 2 " into PO43" to subsequently precipitate H A 2 6 . 42 OTi © HA # DCPD 26 (degree) Figure 2.5-2. X P v D analysis of coatings deposited at current density o f 1, 5, 10, 15, and 20 mA/cm2 for 30 min. © Journal o f Materials Science and Engineering C , 2002, adapted by permission 3 . 2.5.3 Adhesion of H A Coatings Due to the brittle nature and adhesion concerns of H A bioceramic, there exists a challenge for the E C D hydroxyapatite coating to survive under the mechanical stress, i.e. in particular under coronary stent expansion during the angioplasty procedure. A s well known, it is difficult to form a stable bonding at interface between ceramic and metal 9 1 . 3 Reprinted from Journal o f Materials Science and Engineering C , 20, K u o M . C . , Y e n S .K . , The process of electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature, 153 - 160, Copyright (2002), with permission from Elsevier. 43 The adhesion mechanism can be divided into three groups: (1) mechanical interlocking (2) physical bonding (3) chemical bonding. To overcome the difficulty of stable bonding, an interfacially modified surface is necessary between the stainless steel metallic surface and the hydroxyapatite ceramic, for the two surfaces to bond and adhere. Interfacial adhesion is influenced by numerous variables. These variables included: stress in film (both residual stress due to deposition conditions, and due to deformation of the substrate), contamination on substrate, chemical bonding between coating and substrate, physical properties, surface roughness, and precleaning method allowing chemically modified surface 9 2. L i n et al have conducted a study on the growth o f H A on 316L stainless steel in simulated body fluid, and alkaline treatment method was used to promote the adhesive strength of H A coating 9 1 . Alkaline treatment was performed by soaking the substrate in 10M N a O H aqueous solution at 60°C for 24 hours and heat treat at 600°C. The purpose of such treatment was to form an interfacial compound to bridge the 316L stainless steel (with metallic bond) and the H A (with covalent bond). It was found that an interfacial compound of Na4Cr04 formed after the alkaline treatment, and the following reaction was suggested9 1: 8Na(OH) + Cr 2 0 3 => 2Na4Cr0 4 + 3 H 2 0 + H 2 Tensile tests were performed to measure the tensile bonding strength of the coating, in comparison between 1 0 M N a O H treated surface and 1 0 M N a O H treatment followed by 600°C, the bonding strength increased from 25 M P a to 38 M P a 9 1 . 44 The poor adhesion of electrochemically deposited H A can be also attributed to the "bubble effect" while high current (or voltage) was applied. In high current applications, the rate of OH" and H 2 generation increases. When OH" generation rate exceeds the rate of PO4 " formation, the excess OH" groups would migrate away from cathode because of electric field and diffusion . Therefore, the high pH boundary w i l l shift away from the substrate and as a result calcium phosphate precipitation wi l l take place away from the substrate surface, leading to poor coating adhesion. A s hydrogen generation in electrochemical deposition happens at the electrolyte-substrate interface, the increased generation of H 2 often led to a heterogeneous and loose coating structure. Attempts have been made in the past to eliminate such phenomena. These included E C D processing with periodic pulse voltage by Shirkhanzadeh et a l 9 3 , and cathode rotation electrodeposition process by Recently, Hou et al . In the process, the substrate was rotated at high speed (up to 1000 rpm) to remove both hydrogen bubbles and poorly adhered deposit particles. It was observed that the modified process produced a more homogenous and compact coatings which were more difficult to scrub from the substrates87. 4 5 3 SCOPE AND OBJECTIVES 3.1 Scope of the Investigation The principal motivation of the present work is search for better coatings for coronary stents, in particular search for bioceramic coatings such as hydroxyapatite, H A . Bioceramics, such as H A have been used in the medical industries for more than 20 years primarily because of their excellent biocompatibility. However, due to inferior mechanical properties such as low fracture toughness (<1.12 MPaVm ) and low flexural strength (<140 MPa) , the use o f hydroxyapatite has been limited to no-load or low-load bearing applications. Hydroxyapatite coatings have a broad range of applications in medical devices and can provide superior biocompatibility in combination with the advantage of bulk material to achieve the best desired functions of the medical device. Electrochemical deposition (ECD) of uniform hydroxyapatite coating can be achieved on complex substrates. E C D also bears other advantages, such as good control of film thickness in 0.1 - 10 urn range, low temperature of processing (20 - 60°C), uniformity, and deposition rate and owing to low cost of equipment and starting materials. The combination of the potential advantages of E C D for uniform thin f i lm H A coating on complex substrates motivated us to investigate this process for coating coronary stents. There are strong indications that hydroxyapatite thin fi lm coating on coronary artery stents can reduce or eliminate restenosis. Hence, the process for electrochemical deposition o f hydroxyapatite has been studied and optimized in the present thesis. The most significant process parameters (solution chemistry and concentration, temperature, current density, deposition time) were investigated and their effects on coating characteristics (thickness, uniformity, adhesion, phase composition) were evaluated. 46 Although there were previous studies of electrochemical deposition, very few have focused on thin fi lm (<0.5 (am) hydroxyapatite coating and none of these have focused on coatings for stents. The uniqueness of present study is to combine the technique of surface modification, electrochemical deposition, and phase optimization to produce thin film hydroxyapatite coating for cardiovascular applications. The technology has been transferred to M I V Therapeutics, a Vancouver biotechnology company, which currently evaluates the method in a series of in-vitro and in-vivo trials. 3.2 Objectives The broad objective of this study is to achieve a well-controlled and reproducible process to obtain thin f i lm hydroxyapatite coatings via electrochemical deposition, on coronary stents made of 316L stainless steel. The specific objectives of the present thesis are as follows: 1. To develop reproducible hydroxyapatite coating process via electrochemical deposition with full coverage, optimum thickness (<0.5 um), porous (~50vol%) microstructure, and sufficient adhesion to stent surface such that the coatings survive without delamination in the in-vitro stent deployment. 2. To determine the effects of electrochemical deposition (ECD) process parameters on the evolution of the resulting calcium phosphate coating. To develop an optimized electrochemical deposition process to deposit thin fi lm hydroxyapatite coating on coronary stents by studying the process parameters, such as current density, time of deposition, electrolyte concentration, p H , and temperature. 47 3. To determine the resulting hydroxyapatite coating basic properties, such as coating thickness, microstructure, elemental composition, crystallinity, mechanical integrity, and phase composition. 4. To assess the effect of substrate surface pre-treatment on the deposition process and resulting coating in terms of physical (microstructure and phase) and mechanical (adhesion) properties. To apply a substrate surface pre-treatment to achieve high level of adhesion of E C D - H A coatings. 5. To apply a post-treatment process to the coatings to attain desired phase composition (i.e. pure phase H A ) , as well as improved adhesion and mechanical integrity of the coatings. 6. To perform further characterization of the coating in industrial setting: • Qualitative mechanical assessment based on in-vitro crimping and expansion tests. • Rate of dissolution based on in-vitro dissolution tests. • Overall performance qualification based on in-vitro stent deployment and 40 mil l ion cycles fatigue test. 48 E X P E R I M E N T A L M E T H O D O L O G Y The following experimental methods were employed: 1. Measurement of the weight and thickness of electrochemically deposited coatings for different process parameters, such as current density, time of deposition, electrolyte concentration, p H , and temperature. The process stability, reproducibility and ease of control were monitored. 2. Evaluation of microstructural morphology such as uniformity and porosity, and phase composition of various E C D thin film coatings, deposited at different current densities and deposition times, and coatings deposited with substrate (316L stainless steel) surface pre-treatment and coating post- treatment. 3. Qualitative characterization of mechanical behavior o f E C D coating with various substrates surface pre-treatments and coating post-treatment. 4. E C D process optimization, based on the outcome of the above steps 1 - 3 . Application of the optimum E C D process for deposition of H A on coronary stents for further evaluation. 5. In-vitro evaluation of E C D thin film coating on coronary stents, such as crimping and expansion tests, dissolution tests, and fatigue tests. 6. In-vivo evaluation of the coated stents in porcine models (this step done entirely by the collaborating company M I V Therapeutics, Vancouver, B.C. ) 49 4.1 Sample Preparat ion Two types of substrates made of 316L medical grade stainless steel were used in this study, i.e. plate and a real coronary stent. The nominal chemical composition of 316L stainless steel is given in Table 4.1-1. Table 4.1-1. Nominal chemical composition of 316L stainless steel Element C M n Si P S Cr M o Ni N Fe Weight% (max.) 0.030 2.000 0.750 0.045 0.030 18.000 3.000 14.000 0.100 Balance Stainless steel plates were cut into rectangles with a width of 2.5 cm, a height of 3.0 cm and thickness of 0.1 cm. The actual plate area being coated was a square with 2.5 cm width and thickness of 0.1 cm. Unless otherwise stated, all plate specimens were polished with L E C O ® silicon carbide abrasive paper followed by a final 1 um mirror polish with L E C O ® diamond suspension. The polished specimens were degreased by detergent and further ultrasonically cleaned in distilled water and absolute ethanol, then dried. Stainless steel coronary stents were provided by M I V Therapeutics of Vancouver, B . C . A l l received coronary stent specimens were laser cut and electropolished using proprietary M I V T procedure. The M I V I 700 Series Coronary Stent had a 1.7 mm outside diameter, 1.5 mm inside diameter and a length of 14 mm. The M I V I stent is comprised of a series of sinusoidal-ring geometries with two discs forming a single module, a flexible curlicue was fabricated to jo in two modules together. A 14 mm M I V I stent consists of four modules. A schematic diagram of the stent is illustrated in Figure 4.1-1. 50 The strut of the stent had a width of 100 um and the disc had a diameter of 500 urn. Unless substrate surface treatment was performed, all stent specimens were used as- received after cleaning. Stents specimens were cleaned ultrasonically in absolute ethanol, and then dried in air. Figure 4.1-1. Schematic diagram of M I V I 700 Series Coronary Stent 4.1.1 Substrate Surface Modification In order to create a better interfacial bond between the electrochemically deposited coating and the substrate, a substrate surface modification was performed with the use of 10N N a O H alkaline solution 9 1 . The 316L stainless steel substrates were soaked in 10N N a O H aqueous solution at 60°C for 24 hours. After the alkali treatment, the substrates were ultrasonically cleaned with distilled water and dried at 40°C. The alkaline treated substrates were subsequently heat-treated to 500°C for 20 minutes. Surface characterizations were conducted on the alkaline treated substrates. Electrochemical depositions were performed on the alkaline treated coronary stent 51 specimens for mechanical evaluation and in-vitro characterizations, and compared with those deposited on as-received stents. 4.2 Electrochemical Deposition Electrochemical deposition was conducted in a water based electrolyte containing 0.02329M C a ( N 0 3 ) 2 4 H 2 0 and 0.04347M N H 4 H 2 P 0 4 . The p H of the electrolyte was measured by Beckman 260 p H meter with accuracy ± 0.004 p H (Beckman 260, Beckman Coulter, Inc;, Fullerton, California). The electrolyte p H was controlled at 4.5 with the addition of sodium hydroxide to the electrolyte solution. The electrolyte temperature was maintained at 45°C ± 2°C. The stainless steel substrate was used as the cathode, and a platinum foil was used as the anode. The cathodic electrochemical deposition was carried out using a current source with 0.01 raA resolution (IET Model VI-700, IET Lab Inc., Westbury, N e w York) . A l l coated specimens were rinsed in distilled water to remove residual electrolyte and dried in oven at 40°C for 1 hour. Unless otherwise stated, all depositions were performed under the above generic E C D conditions. Figure 4.2-1 illustrates a schematic diagram of the electrochemical deposition. 52 Power Supply Thermocouple - + Amperometer 0 Voltmeter Cathode (Substrate) Anode (Counter Electrode) Electrolyte Hotplate Figure 4.2-1. Schematic diagram of electrochemical deposition setup 4.2.1 Electrochemical Deposition Process Parameters Investigation In order to gain the knowledge of electrochemical deposition for the coronary stent application, understanding the influence of basic process parameters on the resulting coating is essential. Based on the generic electrochemical deposition process the following process parameters were investigated: • Current density. E C D was conducted at current density of 1 - 15 m A / c m for 5 minutes deposition time. • Deposition time. E C D was conducted for duration of 1 - 15 minutes with a current density of 1 m A / c m . • Electrolyte concentration. Concentration of E C D electrolyte was adjusted based on the calcium to phosphors (Ca/P) ratio. E C D was conducted with electrolyte 53 Ca/P ratio of 2.92, 2.63, 1.95 and 0.49, at current density of 1 m A / c m 2 for 2 minutes • Electrolyte pH. Electrolyte p H values of 3.0 - 6.0 with 0.5 increment were used for E C D at current density of 1 m A / c m 2 for 2 minutes. The specific range of p H was chosen for investigation to ensure a stable electrolyte was obtained without precipitation. • Electrolyte temperature. E C D was conducted at current density of 1 m A / c m 2 for 2 minutes deposition time with electrolyte temperature at 25, 45, and 70°C. A l l of the above electrochemical depositions were initially conducted on 316L stainless steel plates for the purpose of microstructural and phase characterization and process observation. The weight measurement and thickness estimation was performed based on the two main reaction rate determining parameters, i.e. current density and deposition time. The measurement of weight was performed on the coated specimens with various current densities and deposition times, and it is based on the weight gain per area (mg/cm 2), with a Sartorius ME235P-SD balance with 0.01 mg accuracy. The thickness (um) estimation was also performed based on weight gain on coated specimens with various current density and deposition time, and through scanning electron microscopy examination of the cross-section area. Evaluation of microstructural morphology such as uniformity and porosity, and phase characterization of the resulting coatings was conducted for all of the above specimens; the process stability and ease of control were also closely monitored for 54 process optimization. For the purpose of process consistency investigation, E C D coating was performed on five batches of ten coronary stent samples using process parameters listed in Section 5.2 (Table 5.2-1), and the coatings were characterized. 4.2.2 Electrochemical Deposition Optimization The optimization of electrochemical deposition process parameters was chosen and detailed in Chapter 5 "Results and Discussion". The optimized parameters were applied and evaluated on electropolished stainless steel coronary stent provided by M I V Therapeutics of Vancouver, B . C . The generic E C D process was performed as discussed in Chapter 4.2.1. Optimized parameters were applied on all stent substrates with a 0.90 m A applied current, and a deposition time of 1 minute. The applied current was specified to assure suitable current density was applied to the stent surface. The deposition time was adjusted to achieve desired thickness of the deposit, i.e. about 0.5 urn for the 1 minute deposition for the above conditions. 4.2.3 Phase Conversion Process It is known that crystalline H A , among various calcium phosphate phases, has the longest resorption life of approximately 2 - 3 years under in-vitro environment. To ensure the resulting electrochemically deposited coating is of stable crystalline 55 hydroxyapatite, a phase conversion process was introduced ' . The phase conversion process was accomplished by soaking the coated substrates into 0.1N N a O H aqueous solution at 75°C for 12, 24, 48, and 72 hours. The measured p H of the 0.1N N a O H solution is approximately 12.5. The converted specimens were gently rinsed with distilled water and dried at 40°C. The converted specimens were then heat treated at 300°C, 500°C, and 700°C for 20 minutes for crystallization. 4.3 Microstructural and Phase Characterizations The microstructural morphology and elemental analysis of the resulting coatings were observed by scanning electron microscopy equipped with energy dispersive x-ray spectroscopy ( S E M / E D X , Hitachi S-300N, Hitachi Ltd., Tokyo, Japan). Crystallography of the resulting coatings was analyzed by X-ray diffractometer ( X R D , Siemens D5000, Siemens, now Bruke r -AXS, Germany) with C u Ka radiation (k= 1.5418 A) and operated at a tube voltage of 40kV and a current of 30mA. The range of 29 was from 5° - 50° with a scanning rate of 0.02°/s. 4.4 In-vitro Evaluations For the application of coronary stents, the E C D - H A coating not only has to withstand deformation during manufacturing stage (i.e. stent crimping), but also at the implantation stage. Furthermore, the coating has to maintain its integrity and resist fatigue stresses in concern with the heart beat over the years after implantation in human. 56 In-vitro evaluations were targeted to assess the coating with crimping and expansion test, dissolution test, and fatigue test. 4.4.1 C r i m p i n g and Expansion Test In order to implant and deliver a coronary stent in the percutaneous transluminal coronary angioplasty procedure, a stent must be first mounted onto a balloon catheter. This procedure in the manufacturing stage is often referred as the crimping process where the stent is crimped under external pressure. The crimping procedure was performed by M I V Therapeutics of Vancouver, B . C . with M S I stent crimping machine (Machine Solutions Inc., M S I SC 513, Arizona, U S A ) . The M S I stent crimping machine applies an external pressure by using a pneumatic crimp mechanism. The E C D - H A coated stents were crimped from an original diameter of 1.7 mm to 1.0 mm onto a balloon catheter (Arriva™, InSitu Technologies Inc., Minnesota, U S A ) . Coronary stent mounted onto a balloon catheter is termed a stent delivery system (SDS). After implantation of the SDS into desired position, surgeon w i l l expanded the coronary stent with an inflation device in order to keep the artery open. The expansion test for E C D - H A coated SDS was conducted with Encore™ inflation device (Encore™ 26 inflation device, Boston Scientific, Maple Grove, M N ) . The stent was expanded from a crimped diameter of 1.0 mm to 3.0 mm with 14 atm pressure. 57 4.4.2 Dissolution Test Solubility of various calcium phosphates differs greatly, which affect the biodegradation of E C D coating. The variables that affect the dissolution not only consist of the phase and crystallinity of the coating, but variables such as p H , the specific nature of the buffer, temperature, and porosity are also important. Dissolution tests were performed with Varian dissolution apparatus (Varian V K 7 5 0 D , Varian Inc., California, U S A ) . Key features of the apparatus included precise bath temperature and rotation speed control, and the use of seal bottles to prevent dissolution media from evaporation. Dissolution tests were conduct at a bath temperature of 37°C and rotation speed at 20 rpm. Phosphate buffer saline (PBS) was used as the dissolution media because it helps to maintain constant p H (7.4) and it is isotonic. P B S contained l O m M phosphate, 140mM N a C l , and 3 m M K C l . E C D coated stents were placed into dissolution apparatus with sealed bottles of 10 m L P B S , and E C D coated stents were weighted between 30 minutes to 4 weeks. 4.4.3 Fatigue Test The objective of fatigue testing is to meet the requirements as per the " F D A Draft Guidance for the Submission of Research and Marketing Applications for Interventional Cardiology Devices" 9 4 for in vitro mechanical fatigue testing. The test should demonstrate the safety of the device from mechanical fatigue failures for at least one year of implantation life. The fatigue test is intended to provide empirical evidence for the continued structural integrity of the E C D - H A coated stents when subjected to mechanical 58 fatigue such as that which they would receive in vivo. The test is designed to simulate the stent fatigue due to the expansion and contraction of the vessel in which it is implanted. The test is accelerated in order to obtain results in a reasonably short time period. The environment for the test is phosphate buffer saline (PBS) at 37°C ± 3°C. The in-vitro simulated fatigue test was equivalent to one year of in-vivo implantation, i.e. approximately 40 mil l ion cycles of fatigue stress, which simulates heart beat rates from 5 0 - 100 beats per minute. Six stent specimens were divided into three pairs for E C D coating. One stent from each pair was implanted into the proximal end of the simulated vessel, another into the distal end. A l l three pairs were E C D coated under the optimum deposition conditions, substrates surface were modified and subsequently treated with the phase conversion process. Final E C D coated stent specimens were then tested with 40 mill ion fatigue cycles. In fatigue testing, the objective was to determine i f the E C D - H A coating is able to survive 40 mil l ion cycles in a vessel that dilated with an average percent outer-diameter (%OD) strain of 0.48%. The fatigue test was performed with a commercial EnduraTec fatigue testing machine (ElectroForce® 9100 Series, EnduraTec System Corporation, Minnesota, U S A ) . A coronary stent with the optimum E C D - H A coating was first crimped onto a balloon catheter as described in Chapter 4.4.1, and then it was deployed into a simulated vessel (artery). A high performance laser was used to continuously monitor and measure the change in outside diameter of the simulated vessel, information 59 .J was subsequently feedback to the control computer to calculate the percent radial stain of stent. Two linear actuators were used to pressurize a saline test solution linked to the test vessel, a pump assembly converts the actuator displacement to test pressures, which simulated 40 mil l ion cycles. At the end of the test cycles, the stent sample was retrieved and the P B S test solution was subsequently passed through a 0.4 um micro-filter to trap potential debris from the E C D - H A coating. Scanning electron microscopy ( S E M ) was used to examine both the coating microstructure and the saline filters for any lose debris. S E M observation was made at xlOO, x800, and x3000 magnification to examine the microstructure of coating at the end of the test to identify any existing micro-cracks or micro-peels. Energy dispersive x-ray spectroscopy ( E D X ) was used to identify the elemental composition of the coating surface or an area of interest such as filtered debris. Both S E M and E D X analyses were used to demonstrate the existence of remains of coating after the fatigue test. Filter specimen was carbon coated to create a conductive surface to avoid charging in S E M . 60 5 RESULTS AND DISCUSSION 5.1 ECD of Calcium Phosphate Coatings - Process Parameters Investigation 5.1.1 Current Density Current density, applied voltage, and electrolyte concentration are the principal variables controlling the rate of reaction in the E C D process 2 5. In order to achieve consistent results, current density instead of applied current, was chosen for investigation because different substrates (stents and plates) were interchangeably used throughout the study. The surface areas of different substrates were determined with the assumption of negligible surface roughness. The S E M observation of plate specimens deposited at current density 15, 10, 5, 3, and 1 mA/cm are shown in Figure 5.1-1 (a) - (e), respectively (other process parameters are listed in Table 5.2-1). The "bubble effects" were found on the coating surface of specimen deposited at 15 m A / c m , as shown in Figure 5.1-1 (a), where the vertical axis of specimen was horizontally aligned in the figure. Previous studies also reported similar phenomenon ' " . This effect was attributed to the increased reaction rate with increasing current density, and as a result, hydrogen bubbles generated at a vigorous 23 86 rate ' . The hydrogen bubbles rise to the surface and disturbed the formation of uniform coating. In general, the microstructure of coating deposited at 15 m A / c m 2 exhibits a fine and loose structure. Figure 5.1-1 (b) illustrates the microstructure of coating deposited at 61 2 10 mA/cm . "Bubble effect" was not observed at such current density, the coating exhibits a fine and loose structure similar to 15 m A / c m 2 , but with flakes growing on top. At 5 mA/cm , the coating exhibits a dense and non-oriented plate-like structure, as shown in Figure 5.1-1 (c). A t 3 m A / c m , the microstructure was also non-oriented plate-like, but less dense and reveals an under-layer of coating (Figure 5.1-1 (d)). A t the lowest current density of 1 m A / c m 2 , a thin layer (~1 um) of porous coating was observed as shown in Figure 5.1-1 (e). The morphology of this thin layer was similar to the under- layer in Figure 5.1-1 (d), but more porous. Other than some small amount of plate-like structures displayed at the deposition boundary (Figure 5.1-1 (e) [left]), the overall microstructure exhibits high porosity and good coverage. Due to technical difficulties, the stainless steel plates were used as-received in this study. A trend of microstructure evolution can be clearly noticed from Figure 5.1-1, as current density changes. Figure 5.1-1 (b) - (d) indicated that the current of 10 - 3 mA/cm facilitate the growth of plate-like crystals, and such microstructure is mechanically unstable since it can be easily detached from the substrate. Thus, to avoid the bubble effect and to prevent growth of plate-like crystals, 1 m A / c m 2 is more desirable for thin film uniform coating. Figure 5.1-2 illustrates the weight gain of E C D coated specimens deposited with current density 1 - 1 5 m A / c m 2 for 5 minutes deposition. The weight gain of coatings increases with increased current density. It was observed that the weight gain slightly reduced with 15 m A / c m 2 , which can be explained by the "bubble effect" where fragments of the coating were flaked off by the large volume of hydrogen bubbles generated. This observation agrees with previous studies 8 6" 8 8 62 Figure 5.1-1. S E M images of E C D coating deposited at various current densities: (a) 15 mA/cm 2 , (b) 10 m A / c m 2 , (c) 5 m A / c m 2 , (d) 3 m A / c m 2 , (e) 1 m A / c m 2 [Left: xlOO; Right: x 1,500] 63 Figure 5.1-2. Weight gain of E C D coated specimens versus current density with 5 minutes of deposition. 5.1.2 Deposition T ime Scanning electron microscopy images were obtained for E C D coatings deposited at various times to evaluation the growth morphology and structural characteristics of the evolving film on plate specimens. Current density of 1 m A / c m 2 was selected for this study based on the resulting uniform thin film. The S E M observation of plate specimens deposited for 15, 10, 5, 3, and 1 min are shown in Figure 5.1-3 (a) - (e), respectively (other process parameters are listed in Table 5.2-1). From the observation of the deposition process, an "initiation period" (-15 seconds) for formation o f coating was noticed. This phenomenon was also reported in other previous studies 8 9 ' 9 5 . A n initiation period corresponds to the time from the application of current to the beginning of calcium phosphate deposition. It is believed that the initiation period is related to the initial kinetic of reduction of water, hydrogen gas generation and increase of p H around the on electrode, and consequently precipitation (deposition) of calcium phosphate . 64 Figure 5.1-3. S E M images of E C D coating deposited with various deposition time: (a) 15 min, (b) 10 min, (c) 5 min, (d) 3 min, (e) 1 min [Left: xlOO; Right: x l500] at 1 m A / c m 2 . 65 Figure 5.1-3 (a) shows the E C D microstructure after 15 min of deposition; plate- like flake structure was observed with a dense under-layer similar to Figure 5.1-1 (d). A s the deposition time decreased to 10 min, the amount of plate-like structure visibly decreased, and revealed the under-layer (Figure 5.1-3 (b)). The under-layer exhibits some "volcano-like" structure, which is also observed in Figure 5.1-3 (c). This is likely due to the bubble effect where the H2 bubbles were trapped during the deposition on process . When the deposition time decreases to 3 min, "volcano-like" structure was no longer seen, instead sparsely distributed deposit was observed as illustrated in Figure 5.1-3 (d). Microstructure of 1 min deposit is shown in Figure 5.1-3 (e). Under x l ,500 magnification, very little material seems to be deposited, however, under xlOO magnification, a noticeable uniform film boundary was observed. Figure 5.1-4 illustrates a x20,000 magnification S E M micrograph of the 1 min deposition, and reveals a thin and uniform porous coating. Due to technical difficulties, the stainless steel plates were used as-received in this study. The weight gain of E C D specimens coated for 1 - 15 minutes with 1 mA/cm is shown in Figure 5.1-5. With increase of deposition time, an increased in weight gain was observed. A more drastic increase of weight gain was seen from 3 minutes to 5 minutes of deposition; this is believed to be a result of the growing crystals on top of the thin fi lm observed in Figure 5.1-4. It is also expected that increasing surface resistance (due to increasing coating thickness) slows down the deposition reaction. It appears that in order to avoid loosely attached plate-like structures and bubble effects, process parameters 66 should be chosen towards lower end, i.e. current density 1 m A / c m 2 and with 1 minute deposition time. x £ 0 k 0 0 0 0 2 0 KV gjJM Figure 5.1-4. High magnification S E M image of E C D coating with deposition of 1 min [x20,000] at ImA/cm 2 . 0.6 - 0.5 — E o 0.4 -- D) e ga in  0.3 - ig ht  0.2 2 0.1 • • • 0 2 4 6 8 10 12 14 16 Deposition Time (min) Figure 5.1-5. Weight gain of E C D coated specimens versus deposition time deposited at 1 mA/cm 2 . 67 5.1.3 Ca /P Ratio The composition and properties of the resulting deposit was previously found to 71 R7 R1 R Q be dependent on the composition of electrolyte ' ' ' ' . However, it was observed that for electrolyte composed of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 , the Ca/P ratio of the resulting deposits were nearly independent of the Ca/P ratio of the electrolyte. Figure 5.1-6 illustrates the different Ca/P ratio of the resulting deposits for various Ca/P ratio electrolytes evaluated i n this work through E D X . Table 5.1-1 shows the concentration of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 for the preparation of different Ca/P electrolytes. The Ca/P ratio of the deposits ranged from 1.41 - 1.57, independent of the Ca/P ratio of the electrolyte. S E M images of the various resulting deposits deposited with electrolyte of Ca/P ratio 2.92, 2.63, 1.95, and 0.49 are shown in Figure 5.1-7 (a) - (d), respectively. Among the four results, the deposit for electrolyte Ca/P 1.95 appears to be most porous (Figure 5.1-7 (c)), but all exhibit similar microstructure. It appears that, within the parameters in this study, the different Ca/P ratio of electrolyte has minimal effect on the Ca/P ratio of the resulting deposit and the resulting microstructure. In the present study, the surface of the stainless steel plate samples was polished to 1 um mirror finish, however a surface roughness created by the deposited E C D coating may cause unnecessary absorption of the generated X-ray signal, which is difficult to account for in the E D X quantification analysis. The accuracy of E D X quantitative analysis with the use of well-polished standards having a composition similar to the sample has been reported to be greater than 2% relative for major concentrations 1 0 5. The analysis of elements with concentrations less than 5wt% w i l l typically yield relative 68 accuracies of -10%, even with standards. It was also reported that for samples with rough surfaces, i.e. fracture samples or small particles, the relative accuracy may be as low as 50% [ 1 0 5 ] . Therefore, the E D X analysis in this study may not reflect the true concentration and should only be considered as an relative indicator. Table 5.1-1. The concentration of C a ( N 0 3 ) 2 4 H 2 0 and N H 4 H 2 P 0 4 for the preparation of different Ca/P ratio electrolytes. Electrolyte Concentration (M) Electrolyte Ca/P ratio NH4H2PO4 Ca(N03)2 4H20 0.0217 0.0635 2.92 0.0217 0.0572 2.63 0.0217 0.0423 1.95 0.0435 0.0212 0.49 1.8 W o a. 0 Q Ui c 3 (fl O CL O O cc CC a. re O 1.6 1.4 1.2 1.0 0.8 0.6 0.4 0.2 0.0 2.92 2.63 1.95 Ca/P Ratio of Electrolyte 0.49 Figure 5.1-6. Ca/P ratio of resulting deposit with the use of various Ca/P ratio electrolytes. Ca/P ratio was derived from E D X spectra. 69 Figure 5.1-7. S E M images of resulting deposits from various Ca/P ratio electrolytes; a) 2.92, b) 2.63, c) 1.95, and d) 0.49. [x 15,000] 5.1.4 Temperature The electrolyte temperature in the E C D process was found to be influential on the uniformity of coverage and on the coating microstructure. Figure 5.1-8 illustrates the microstructure of coatings deposited at 25°C, 45°C, and 75°C, respectively (other process parameters included: current density = 1 m A / c m 2 , deposition time = 5 minutes, electrolyte pH = 4.2, electrolyte Ca/P ratio = 1.95). It can be observed that the coating uniformly improved from Figure 5.1-8 (a) to (b), however, the microstructure appears to be denser and exhibit cracks in Figure 5.1-8 (c). 70 b) Figure 5.1-8. S E M images of resulting CaP deposits conducted in electrolyte with temperature a) 25 °C b) 45°C, c) 75°C. [x 15,000] While a constant current density was applied (i.e. constant current source, for a fixed geometry substrate), the influence of temperature on the supplied voltage was noticed. Figure 5.1-9 shows typical influence of the electrolyte temperature on the measured supply voltage. The voltage decreases as the electrolyte temperature increases, possibly due to decrease of water resistivity, as previously suggested . 71 Figure 5.1-9. Influence of electrolyte temperature on measured supply voltage, for constant current source (I = 13.77 mA) . 5.1.5 The Influence of Electrolyte p H In general, hydroxyapatite is known to dissolve in an acidic environment with p H < 2 and precipitate in a basic environment with p H > 9 [ 6 7 ] (Figure 2.3-1). However, the phase composition, crystallinity, nature of solution, temperature, and porosity also plays a role in dissolution and precipitation of calcium phosphates6. In the process of electrolyte preparation, slight precipitation was observed when the p H reached 6.0 at 45°C. A stable electrolyte without precipitation was found to lie between p H 4.0 - 5.5 at 45°C. E C D conducted with electrolyte p H 3.0 at 45°C has demonstrated non-uniform surface coverage as shown in Figure 5.1-10. S E M images of E C D coating deposited with electrolyte of p H 4.0, 4.5, and 5.5 are shown in Figure 5.1-11 (a) - (c), respectively (other process parameters are as follows: current density = 1 m A / c m , deposition time = 5 minutes, electrolyte temperature = 45°C). A l l coatings exhibit uniform surface coverage, there were no distinctive microstructural differences observed among the three deposits. 72 Therefore the range of p H 4.0 - 5.5 was concluded to be optimal for the process under consideration. Figure 5.1-10. S E M Image of E C D conducted with p H 3.0 electrolyte at 45°C. [x5,000] Figure 5.1-11. S E M images of E C D coatings deposited with various electrolyte p H : a) 4.0, b) 4.5, and c) 5.5. [x3,000] 73 5.2 ECD of Calcium Phosphate Coatings on Coronary Stents 5.2.1 Deposition Process Optimization It is obvious from the microstructural observations that both current density and deposition time plays an important role on the amount and microstructure of the E C D coatings. In order to achieve thin fi lm (<0.5 um) coating and to maintain a desire coating microstructure (i.e. porous structure" without large plates of crystal over-growth), it was determined that current density should be 1.0 m A / c m , and deposition time should be 1 minute (other standard E C D parameters kept constant, as listed in Table 5.2-1). It was observed that the application of different current densities had a more significant effect on the coating microstructure compared to deposition time, whereas the deposition time had more control of weight (or thickness) of the resulting coatings. For the purpose of depositing thin porous coating as illustrated in Figure 5.1-1 (e) on coronary stent, current density of 1 m A / c m was chosen for deposition with various deposition times. The Ca/P ratio of E C D electrolyte was chosen to be 1.95. Although there was no remarkable microstructural difference observed with various Ca/P ratio (refer to Chapter 5.1.3), the coating with Ca/P 1.95 exhibits the most porosity and highest Ca/P ratio (Figure 5.1-6., Figure 5.1-7) Deposition temperature for the optimized process was selected at 45°C, as such electrolyte temperature showed high consistency through the five batches of E C D coatings (Appendix B) , and at the same time provides a crack-free uniform resulting coating (Figure 5.1-8). The electrolyte p H was determined to be 4.5, as this electrolyte pH was found to facilitate uniform deposition. 74 Table 5.2-1. Optimum parameters for E C D of calcium phosphate coatings*. E C D Process Parameters Value Electrolyte Ca/P ratio Calcium Nitrate [ C a ( N 0 3 ) 2 4 H 2 0 ] Ammonium phosphate [ N H 4 H 2 P 0 4 ] 1.95 0.04347M 0.02329 M Current Density 1.0 m A / c m 2 Deposition Time 1 minute Electrolyte p H 4.5 Electrolyte Temperature 45°C V ~ Electrolyte solution was prepared with distilled water, typical volume of electrolyte = 400mL. Figure 5.2-1 shows S E M images of 316L stainless steel bare metal stent manufactured by M I V Therapeutics, Vancouver, B C . Minor bumps can be observed on the edge areas as seen in Figure 5.2-1 (b) & (c), believed to originate from the laser cutting process. Figure 5.2-2 to Figure 5.2-5 illustrates the E C D coating deposited with the optimum parameters (Table 5.2-1) on coronary stents for 1, 2, 3, and 5 minutes, respectively, at magnifications of xlOO, x300, x800, and x 1,500. A s can be observed in Figure 5.2-2, a thin film (-0.5 um) coating was indeed achieved with the use of optimum parameters, for 1 minute deposition (coating thickness was estimated based on cross- section evaluation of stent coating, Figure 5.2-8). The coating is observed to be covering the entire stent surface uniformly, loosely attached crystals were not found. The high conformance of coating can also be seen on the edges of the stent where the laser cutting bumps were found. The coating exhibits a porous structure demonstrating a possibility for drug encapsulation. Figure 5.2-6 shows the E D X surface analysis results for the 1 minute deposition. Calcium and phosphorous were found, confirming the existence of a calcium phosphate coating. Increasing the deposition time to 2 minutes clearly increased the density of the coating as seen in Figure 5.2-3. Nevertheless, full coverage and 75 uniformly was achieved again. Non-uniformity was observed when the deposition time was further increased to 3 and 5 minutes (Figure 5.2-4 and Figure 5.2-5). Nucleation of loosely attached structures was observed sparsely distributed over the entire stent surface in Figure 5.2-4. Large crystal plates were seen with 5 minutes deposition as observed in Figure 5.1-3 for prolonged deposition time. From the observation of both Figure 5.2-4 and Figure 5.2-5, it is noticed that the laser cutting bumps were no longer visible, indicating the increasing of thickness and density of the coating as deposition time increased. Figure 5.2-1. S E M images of bare metal stent with various magnification: a) [xlOO], b) [x800], c) [xl,500]. 76 Figure 5.2-2. S E M images of E C D coating deposited on coronary stent with optimum parameters for 1 minute, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] Figure 5.2-3. S E M images of E C D coating deposited on coronary stent with optimum parameters for 2 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [xl,500] 77 Figure 5.2-4. S E M images of E C D coating deposited on coronary stent with optimum parameters for 3 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] Figure 5.2-5. S E M images of E C D coating deposited on coronary stent with optimum parameters for 5 minutes, (a) [xlOO], (b) [x300], (c) [x800], and (d) [x 1,500] 78 Counts 1000 800 600 - 400 - 200 - Fe Ni Si Ca Fe Cr Mn Mn Fe Ni Ni T 1 | I 1 1 1 1 1 1 [ — 1 1 1 4 6 8 k e V Figure 5.2-6. E D X surface analysis of the E C D coating deposited on coronary stent with optimum parameters for 1 minute. Figure 5.2-7 shows the X-Ray diffraction pattern of the E C D coating deposited with optimum process parameters (Table 5.2-1), for 1 minute. Due to the technical difficulty of X-Ray diffraction analysis on thin film (-0.5 um), E C D coating was conducted on ten stainless steel plate substrates (as described in Chapter 4.1), then scraped off and collected onto a stainless steel plate for analysis. Two phases can be clearly noticed in the coating; the D C P D peaks (JCPDC# 01-072-0713) are at 2-theta 11.64°, 20.95°, 29.28°, 34.16°, and 41.78°, the H A peaks ( J C P D C #009-0432) are at 2- theta 25.98° and 31.94°. Stainless steel substrate has a peak at 2-theta 44.52°. However, due to the similarities of H A and O C P X-ray diffraction patterns (2-theta 25 - 40°), the two phases can not be clearly distinguished. A s indicated by the peak width, the H A 79 phase was poorly crystallized. Similar X R D analyses on H A were also reported previously, where poorly crystalline H A was characterized and identified by the broad peak w i d t h 2 3 ' 7 3 ' 8 4 ' 9 8 . Since the coating consisted of highly soluble D C P D phase and poorly crystallized H A phase, it is reasonable to believe that such coating bears high solubility. Further dissolution test results (Section 5.2.4 - Table 5.2-2) have shown that the coating lost 50% of its weight ( D i / 2 ) within 20 minutes, and was totally dissolved (Diotai) in approximately 40 minutes, in phosphate buffer saline. • c c D A 1 O , • \ 5 10 15 20 25 30 35 40 45 50 2-Theta (deg) O Dicalcium phosphate dehydrate (DCPD) A Hydroxyapatite (HA) I Octacalcium phosphate (OCP) • 316L Stainless Steel Figure 5.2-7. X - R a y diffraction of the E C D coating deposited with optimum parameters (Table 5.2-1), for 1 minute, showing a mixed phase of D C P D and H A . In addition, cross-section study was performed on the E C D coating deposited with the optimized parameters for 1 minute. Figure 5.2-8 illustrated the S E M image of the 80 coating cross-section. It was revealed that the coating thickness of an E C D coating with 1 minute deposition was approximately 0.5um. Epoxy Figure 5.2-8. Cross-section S E M image of E C D coating deposited on stent. Estimated coating thickness was approximately 0.5um. 5.2.2 In-vitro Crimping and Expansion Tests on ECD Coated Stents A qualitative coating adhesion assessment was performed by in-vitro stent crimping and expansion test on an E C D coated stents with CaP deposited at optimum parameters (Table 5.2-1). The simulation tests performed by M I V T indicated that the maximum strain experienced by stent during expansion may reach up to 15% in the bottom of " V " section of the stent2 7. Figure 5.2-9 illustrates an expanded bare metal stent. It can be seen that upon expansion, the metal surface suffers significant plastic deformation (i.e. the slip bands), which leads to delamination of coating shown in Figure 5.2-10. The deformation of a 81 coated stent after expansion can be seen in Figure 5.2-10 (a) and (b) and can be compared with a stent before expansion in Figure 5.2-1 (a). Two distinctive deformation areas can be observed: a compressive stress area in Figure 5.2-10 (c) and a tensile stress area in Figure 5.2-10 (d). The compressive stress area exhibits a buckle delamination effect as illustrated schematically in Figure 5.2-11". Observations in Figure 5.2-10 (c) have indicated that the interfacial adhesion is low and the coating buckling lead to the loss of large coating section. The tensile stress in brittle films can lead to through-thickness cracking as illustrated schematically in Figure 5.2-12 9 9. Figure 5.2-10 (d) shows tensile stress causing a series of parallel cracks with approximately uniform spacing (~5 um) indicating a poor interfacial bonding. The delamination allowed better estimation of the coating thickness, observed again to be -0.5 um, which agrees with the cross-section evaluation illustrated in Figure 5.2-8. Figure 5.2-9. (a) S E M images of an expanded bare metal stent [xlOO], (b) high magnification revealing a significantly deformed surface [x3,000]. 82 Figure 5.2-10. S E M images of expansion test result from an E C D coated stent specimen deposited with optimum parameter for 1 minute deposition, (a) Expanded area [x50] (b) Expanded area [x300] (c) Compressive stress area showing coating delamination [x800] (d) Tensile stress area showing coating delamination [x800] (a) (5c — coatina CTc substrate area of decohesion (b) 0C ^ _ ^ - > * — C c Figure 5.2-11. Compressive spallation by buckling showing localized interfacial decohesion. © Journal of Engineering Failure Analysis 2, 1995, adapted by permission 4 . 4 Reprinted from Journal o f Engineering Failure Analysis 2, 2, Strawbridge A , Evans H.E. , Mechanical failure of thin brittle coatings, 85 - 103, Copyright (1995), with permission from Elsevier. 83 Figure 5.2-12. Tensile stress in brittle film causing through-thickness cracking and interfacial delamination. © Journal of Engineering Failure Analysis 2, 1995, adapted by permission5. 5.2.3 Substrate Surface Modification for Improvement of Coating Adhesion Substrate surface modification (as described in Chapter 4 Experimental Methodology) was employed in this study to enhance the interfacial adhesion between the coating and the substrate9 1. Figure 5.2-13 shows the S E M image of a stent surface treated with 1OM N a O H aqueous solution at 60°C for 24 hours and after subsequent heat treatment at 500°C for 20 minutes. It was observed that the pretreated metal surface exhibits a nano-rough surface structure, with the characteristic size of the surface pattern on the order of 100 nm. Figure 5.2-14 illustrates the E D X analysis of the pretreated 316L stainless steel surface. While Fe, N i , and Cr are the main constituents of 316L stainless steel, it was reasonable to believe that the surface included a new compound Na4Cr04 as reported by L i n et al formed on the stent surface after surface modification 9 1 . The nano- rough surface structure formed during the surface modification process is believed to promote mechanical interlocking and thus physical bonding. Additionally, the inter- compound layer of Na4Cr04 may act as a chemical bonding bridge between the metal and ceramic. However, it has not been clarified in the present work which factor specifically 5 Reprinted from Journal of Engineering Failure Analysis 2, 2, Strawbridge A, Evans H.E. , Mechanical failure of thin brittle coatings, 85 - 103, Copyright (1995), with permission from Elsevier. 84 (i.e. nano-roughness or modified surface chemistry) helps to improve the coating adhesion to stent surface. In-vitro stent expansion was performed on a bare metal stent with surface modification as illustrated in Figure 5.2-15. Under the high magnification of 10,000x, a surface layer can be clearly seen with cracks (Figure 5.2-15 (d)). Such layer was not observed in the bare metal stent expansion test illustrated in Figure 5.2-9. It was therefore reasonable to believe that the cracked inter-compound layer observed was the indeed the ceramic Na^CrC^ layer. SE WD 4.7mm 5.OOkV x20k 2um Figure 5.2-13. S E M image of a stent surface after surface modification [x20,000]. The surface showed a nano-size roughness. 85 Counts -I Figure 5.2-15. Expansion test result from a bare metal stent specimen after surface modification, (a) [xlOO], (b) fx800], (c) [x3,000], and (d) [x 10,000]. The N a 4 C r 0 4 inter- compound can be seen with cracks. 86 Electrochemical deposition with the optimized deposition parameters (Table 5.2-1) was conducted on the surface modified stent (Figure 5.2-16). It was observed that the coating exhibits 100% coverage with excellent uniformity. The deposit on laser cut bumps as shown in Figure 5.2-16 (c) demonstrated high conformance. Observation under high magnification (x 10,000) as illustrated in Figure 5.2-16 (d) revealed a -50% porous microstructure. It appears that although electrically insulating, the Na4Cr04 surface layer is in no way disturbs the E C D process. Figure 5.2-16. S E M images of E C D coating deposited with optimum deposition parameters on surface modified stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 87 Figure 5.2-17 shows the S E M images of the coated stent after expansion test (with the use of a 3.0 mm diameter catheter) with an E C D coating deposited using the optimum deposition parameter (Table 5.2-1), on surface modified stent. There is no coating buckling or delamination. However, both Figure 5.2-17 (c) & (d) revealed a series of parallel nano-cracks (<100nm) with approximately uniform spacing. The nano-cracking with no observed coating decohesion proved an improved interfacial bonding between the E C D coating and the surface modified stent. Figure 5.2-18 shows the S E M images of a more severe expansion test results (known as "over-expansion" in industry, with the use of a 3.5 mm diameter catheter) for an E C D coating deposited with optimum deposition parameters, on surface modified stent. Figure 5.2-17. S E M images o f expansion test result from an E C D coating deposited with optimum deposition parameters on surface modified stent. Expansion performed with a 3.0 mm diameter catheter, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 88 The degree of deformation in Figure 5.2-18 (a) was notably higher compared to Figure 5.2-17 (a), although the degree of strain can not be quantified. Even though there were obvious nano-cracks (<100nm) present in the coating, yet no delamination or separation of the coating was seen in Figure 5.2-18 (d). Fracture of such coating to accommodate the strain was observed to be localized, i.e. the nano-cracks were limited to small (<100nm) areas adjacent to the pores, in the areas of the highest strain on the expanded stent. It was observed that these nano-cracks may link to form larger, 1 - 10pm long cracks, but without separation of the coating from the substrate. It is concluded that, in comparison with coating deposited on non-surface modified stent (Figure 5.2-10) the adhesion of coating deposited on stent with surface modification has shown significant improvement. The increase in adhesion is believed to be due to (i) the surface nano-roughness created during the surface modification process, thus enhanced mechanical interlocking with the E C D coating; and/or (ii) the formation of the Na4Cr04 inter-layer that have changed the surface chemistry, promoting formation of a chemical bond between the modified surface and the coating during the E C D process. 89 Figure 5.2-18. S E M images o f expansion test result from an E C D coating deposited with optimum deposition parameters on surface modified stent. Expansion performed with a 3.5 mm diameter catheter, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 5.2.4 Phase Composi t ion of E C D Ca lc ium Phosphates Coatings To prolong resorption of the E C D as-deposited mixed-phase calcium phosphate coatings, phase and crystallinity optimization was performed. A s mentioned in Chapter 5.2.1, total dissolution of as-deposited E C D coating occurred within 40 minutes in phosphate buffer saline at 37°C (pH = 7.4). The high rate of dissolution was believed to be due to the mixed-phase o f coating, including poorly crystalline H A and dicalcium phosphate dihydrate (DCPD) . Here after, the E C D deposited coating with the use of the optimum deposition parameters wi l l be described as the "as-deposited coating." In order to achieve phase conversion toward substantially pure H A , as-deposited calcium 90 phosphate coatings were submerged into 0 .1M of N a O H aqueous solution at 75°C for 12, 24, 48, and 72 hours 8 1 ' 1 0 ° . Figure 5.2-19 illustrates the X-Ray diffraction patterns of as-deposited E C D coating and after various time of the above phase conversion process. Upon the phase conversion process, D C P D peaks were no longer found in the patterns and only the H A peaks were found. The intensity of H A peaks increased as the conversion time increased. With 72 hours conversion, the four most representative peaks of H A at 2-theta 25.87°, 31.77°, 32.19°, and 32.90°, were distinctively detectable. This can be attributed to the transformation according to the following reaction 9 6: 5 C a H P 0 4 - 2 H 2 0 + 60H" o C a 5 ( P 0 4 ) 3 O H + 2 P 0 4 3 " + 1 5 H 2 0 It was reported that D C P D is an unstable phase above p H 6.9 and transforms to o 1 hydroxyapatite . Our experimental results have demonstrated the same. Although the H A peaks were more distinctive comparing to that in the as-deposited coating, H A crystallinity (as indicated by the peak width) was still considerably low. Dissolution test results (Section 5.2.4 - Table 5.2-2) have revealed that the coating lost 50% (Di/ 2 ) of its weight within 4 hours, and was totally dissolved (D-rotai) in 6 Vi hours. 91 Figure 5.2-19. X - R a y diffraction patterns of as-deposited E C D coating and the resulting coating after 12, 24, 48, and 72 hours of NaOH( a q ) phase conversion at 75°C. The microstructure of the E C D coating on stent with 12 hours conversion process is illustrated in Figure 5.2-20. In contrast to the as-deposit sample, the pores size appear to be larger (i.e. 1.0 - 1.5 jam vs. 0.2 - 0.5 urn in the as-deposited coating). This maybe attributed to the phase transformation process in which the D C P D phase was dissolved and re-precipitated as H A 1 0 0 ' 1 0 1 . Similar patterns were also found in the microstructure of E C D coating after 72 hours of the conversion process as illustrated in Figure 5.2-21. S E M observations revealed that all the E C D coatings upon the conversion process have maintained full coverage and uniformity on the stent substrate. 92 Figure 5.2-20. S E M images of a 12 hours phase conversion E C D coating deposited with optimum deposition parameters, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] L I • * J • ' f„> J •:• • .;: f S Figure 5.2-21. S E M images of a 72 hours phase conversion E C D coating deposited with optimum deposition parameters, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 93 It is generally believed that an amorphous H A coating has higher dissolution rate than a crystalline H A coating 6 5 . It is common to measure crystallinity by using diffraction techniques, such as x-ray diffraction: greater crystallinity yields a sharper 102 103 diffracted beam ' . The higher the heat treatment temperature, the larger and more perfect the crystals and thus the lower the degradation (dissolution) rate. Therefore, easily resorbable calcium phosphate materials are usually heat-treated CaP materials 3 0. High heat treatment temperatures induce H A crystals growth, and thus lead to the decrease of total porosity and pore s i z e 7 3 ' 1 0 4 . Depending on the calcium to phosphorous ratio, presence of water and impurities, amorphous H A can be transform to crystalline H A 6 ' 3 0 . Therefore, a heat treatment after the N a O H conversion was also performed. Three as-deposited calcium phosphate coatings were converted with 0 .1M N a O H aqueous solution for 12 hours at 75°C, followed by three different heat treatment temperatures at: 300°C, 500°C, and 750°C for 20 minutes. Figure 5.2-22 illustrates the X-Ray diffraction patterns of the E C D coating after 12 hours of 0.1N NaOH( a q ) phase conversion and after the subsequent 300°C, 500°C, and 750°C heat treatment of 20 minutes. A s the heat treatment temperature increases, the peaks become narrower and individual peaks of H A become more distinctive. After 300°C heat treatment, no significant change occurred in the X R D pattern; it was observed that the sample still consisted of poorly crystallized H A . A t 500°C, the signature peaks (as indicated by J C P D C #009-0432) of H A at 2-theta 25.87°, 28.96°, 31.77°, 32.19°, 32.90°, and 34.05° have became more distinctive and the intensity had notably increased. Furthermore, the width of the peaks was considerably narrower as compared to the 300°C 94 samples. Upon heat treating the coating at 700°C, a pure phase of H A was formed. It was observed that both the intensity and 2-theta have closely matched the reference X R D pattern of hydroxyapatite. The two higher intensity peaks at 31.77° and 32.19° were found to be well defined. Some of the peaks not noticeable at <700°C, i.e. at 2-theta 10.85°, 16.79°, and 46.71°, were found to have became more evidenced. This trend of phase evolution have suggested that the 0.1N N a O H conversion and the subsequent heat treatment does optimize the H A phase and does increase the crystallinity. • e 1 600 400 25 30 2-Theta (deg) A Hydroxyapatite (HA) • 316L Stainless Steel Figure 5.2-22. X - R a y diffraction patterns of 12 hours NaOH( a q ) phase converted E C D coating and after 300°C, 500°C, and 750°C for 20 minutes of heat treatment of the coating. Although the width of H A peaks became narrower as heat treatment temperature increases, it has not been clarified in the present work that it was due to the increase of crystallinity or grain size. Numerous studies suggested that the narrowing of H A peaks is 95 indication of higher crystallinity ' ' ' . Figure 5.2-23 (b) & (c) illustrates the microstructure of the 500°C and 750°C heat treated coating on stent, respectively. Although the overall microstructure of the heat treated coating was observed to be similar, the coating density was observed to have increased as compared to the pre-heat treated coating (Figure 5.2-23 (a)). It was also observed that the overall integrity of the coating appeared to have increased. The loosely attached structure found in the pre-heat treated coatings has formed a closely inter-connected porous structure. In the dissolution tests for E C D coating, it was found that the time needed for 50% coating dissolution (D1/2) increased from 20 minutes to 4 hours upon NaOH( a q ) treatment. Furthermore, the phase conversion process (with NaOH( a q ) treatment + 500°C heat treatment) prolonged E C D - H A coating dissolution up to four weeks, with only - 6 % weight lost during the test. Table 5.2-2 summarizes the dissolution results for the E C D coatings. Figure 5.2-24 shows the S E M images of an E C D - H A coated stent after NaOH( a q ) treatment and 500°C / 20 minutes heat treatment, after stent expansion. High degree of substrate deformation is noticed, since a 3.5mm diameter catheter was used for the expansion test. Examinations of both tensile and compressive areas have found no detrimental cracking or delamination of the coating. Although some fine cracks (~300nm) were noticeable, it is reasonable to believe that the heat treatment process has increased the integrity of the coating. 96 Figure 5.2-23. S E M images of E C D coating after phase conversion process [xlO,000]. (a) 12 hours N a O H ( a q ) treatment (b) 12 hours N a O H ( a q ) treatment + 500°C heat treatment (c) 12 hours N a O H ( a q ) treatment + 750°C heat treatment. Table 5.2-2. Summary of E C D coatings dissolution test data. Dissolution tests were conducted with 10 m L of phosphate buffer saline (PBS) at 37°C (pH = 7.4) with rotation speed at 20 rpm. 55 E C D Coat ing Description ** Time to D1/2 ** Time to DT otai As deposited E C D Coating 20 minutes 40 minutes E C D Coating with N a O H ( a q ) Conversion 4 hours 6 Vi hours E C D Coating with N a O H ( a q ) Conversion + 500°C heat treatment > 4 weeks > 4 weeks * All Coatings were prepared as described in Table 5.2-1. ** D 1 / 2 denotes the time of 50% coating weight lost, DT o tai denotes the time of total coating dissolution. 97 Figure 5.2-24. S E M images o f expansion test result from an as deposited E C D coating upon NaOH( a q ) treatment + 500°C heat treatment. Expansion performed with a 3.5 mm diameter catheter, (a) [x50], (b) [x300], (c) Showing the compressive stress area [x 1,500], and (d) Showing the tensile stress area [x 1,500] 5.2.5 In-vitro Fatigue Test The purpose of this test was to ensure E C D - H A coated stents are fatigue-resistant in an accelerated 40 mill ion cycles test, (i.e. simulating about 1 year of heartbeat), and to ensure no catastrophic failure occurred to the integrity o f E C D - H A coating. This is to provide empirical evidence for the continued structural integrity of the E C D - H A coated stents when subjected to mechanical fatigue such as that would receive in vivo. S E M was used to examine the coating surface at the end of the test and to examine saline filters for any lose debris. E D X was used to identify the elemental composition of the coating 98 surface or an area of interest such as filtered debris. The test was performed at M I V Therapeutics, Vancouver, B . C . A s described in Chapter 4.4.3, six E C D - H A coated stent specimens were divided into three pairs, and one stent from each pair was implanted into the proximal end of the simulated vessel, another into the distal end. The target % O D strain of the fatigue test was 0.48%. Figure 5.2-25 summarizes the actual % O D strain o f the six fatigue tested stent specimens. Percent O D strain for vessel #1 was found to be the closest to the target at -0.515%, vessel #2 exhibits the highest strain at -0 .565% and vessel #3 shows the least strain at -0.395%. Specimen Position Spcimcn ID A^eiage '• OD Strain Vessel #1 Proximal 1P 0.51 Vessel #1 Distal 1D 0.52 Vessel #2 Proximal 2P 0.57 Vessel #2 Distal 2D 0.56 Vessel #3 Proximal 3P 0.35 Vessel #3 Distal 3D 0.44 0.6 0.5 1P 1D 2P 2D 3P 3D Specimen ID Figure 5.2-25. Summary of average % O D strain for the six fatigue tested stent specimens. 99 S E M images of specimen IP after the fatigue testing are shown in Figure 5.2-26. Crystals of phosphate buffer saline (PBS) used as the test media, and N a C l were found on stent surface, as P B S contained l O m M phosphate, 140mM N a C l , and 3 m M K C l . Due to salts buildup, it was not possible to evaluate the coating by weight measurement. Instead, since PBS did not contain any calcium, the content of calcium in E C D - H A was used as an indicator of the coating existence. Figure 5.2-26 (b) illustrates the area analysis by E D X . The results have shown a high count of calcium suggesting the E C D - H A coating still remains on the stent surface (Figure 5.2-26(d)). The side edge of stent IP is shown in Figure 5.2-26 (c). While the coating microstructure on the edge area remains the same, the coating microstructure on the top surface was obviously thinner. The "thinning" of the coating was likely due to the contact between the vibrating vessel wal l and the top coating surface. However, as indicated by E D X analysis, it was reasonable to conclude that the E C D - H A coating remains adhered to the stent surface. 100 Figure 5.2-27 illustrates the S E M image of specimen 2P after the fatigue test. Even though specimens in vessel #2 suffered the highest % O D strain, there was no cracking or peeling found in the 2P coating. Similar to specimen IP, N a C l crystals were found on the stent surface. Figure 5.2-27 (b) shows the area analyzed by E D X , distinctive calcium content was found, indicating presence of E C D - H A coating. The characteristic porous microstructure of the E C D coating can be seen in Figure 5.2-27 (c). However, the coating was observed to be somewhat denser than before fatigue testing; this is believed to be due to the crystallized N a C l or K C l inside the pores during the 101 fatigue test. Nevertheless, both E D X and surface microstructure have suggested that the E C D - H A coating remains on the stent surface. Figure 5.2-27. S E M images of explanted fatigue tested specimen 2 from vessel #2 proximal end (a) [xlOO], (b) [x800], (c) [x3,000] (d) E D X analysis from surface area scan. The % O D strain of vessel#3 was the least among the three vessels, the S E M image of the specimen 3P after fatigue test is shown in Figure 5.2-28. Once again, N a C l crystals were found on the stent surface. And , similar to specimen 2P, a denser porous microstructure was observed (Figure 5.2-28 (c)). Unlike specimen IP and 2P, the coating microstructure of specimen 3P in the stent edge area and stent surface area have shown 102 minimal difference. Closer examination of Figure 5.2-28 (b) reveals a uniform microstructure from the stent edge to surface, and no "thinning" was observed, Figure 5.2-28 (c). Surface area scan with E D X analysis suggested the E C D - H A still remains with high calcium content. Figure 5.2-28. S E M images o f explanted fatigue tested specimen 2 from vessel #3 proximal end (a) [xlOO], (b) [x800], (c) [x3,000] (d) E D X analysis from surface area scan. At the end of the 40 mill ion cycles fatigue test the PBS test solution was passed through a 0.4 um micro-filter to trap potential debris released from the fracturing E C D - H A coating. Numerous debris were found on filter, and S E M and E D X characterizations 103 were performed over the different debris of interest (i.e. suspected H A debris and other contaminations). S E M and E D X characterizations indicated that there were several types of debris, varying in shape, morphology, and elemental content, present in the fatigue test system. In general, six categories of debris can be recognized. Figure 5.2-29 illustrates the six main categories of debris found, namely D l - D6. Debris size was observed to range from 5um - 30um diameter. E D X analyses of the six debris were illustrated in Figure 5.2-30 respectively. D l and D3 both contain high silicon content, as shown in Figure 5.2-30 E D X - D 1 and E D X - D 3 . It was believed that the silicon content was from the silicon-made vessel wall. The only debris found to contain calcium was D2, nevertheless, calcium content was minimal (0.37wt%). Judged from the high phosphorus content, it is believed that D2 was a PBS chlorine crystal buildup on coating surface, possibly detached during the test along with a trace of E C D - H A coating. Most important of all , none of the debris found on filter resemble the shape or microstructure of the E C D coating. Debris D2 , D4, and D6 all contain high content of chlorine from PBS based on the E D X result in Figure 5.2-30 E D X - D 2 , E D X - D 4 , and E D X - D 6 , respectively. Other commonly found debris was D5 containing high sulphur content as illustrated in Figure 5.2-29. Additional elements such as copper, zinc, and magnesium were often found incorporated in debris, the origin o f these elements was unknown. Although E D X analysis was performed to determine the presence of different elements, due to the intensity overlap caused by the filter material (i.e. carbon and 104 oxygen), the concentration result may not be referred to as the exact percentage of each element in the analyzed area. In other words, the concentration presented in the E D X analysis, may not reflect the true concentration and should only be considered as an indicator. This is because the dispersion energy used in the E D X analysis was high enough to penetrate to the substrate (filter), the elemental analysis data of the debris may also contains elements of the substrate and the surrounding area. Therefore, as seen in Figure 5.2-30, a high content of carbon and oxygen from the filter material is clearly distinguishable. Also , it was reported that for rough surface samples or small particles, the relative accuracy of E D X may be as low as 50% [ 1 0 5 1 . From the microstructural observation and elemental analysis carried out on the fatigue tested stent specimens, there was no visible crack or delamination occurred on the coating surface at x3,000 magnification. Although "thinning" of the surface coating was observed, E D X analysis have proved that the E C D - H A still remains attached to the stent surface. Furthermore, S E M observations of the debris filters have not found any debris resembling the E C D - H A coating. E D X analyses on the found debris have not suggested any debris with high calcium content originating from E C D - H A coating. 105  EDX-D2 EDX-D5 150C A Figure 5.2-30. E D X analysis of the six main categories of debris found after the fatigue test. 107 5.2.6 Reproducibi l i ty and Consistency of E C D - H A Process The E C D coating process (at optimum coating parameters as listed in Table 5.2-1) and the phase conversion process (i.e. 0.1N NaOH( a q ) + 500°C heat treatment) on the surface modified stent, were run repeatedly to evaluate process reproducibility. Five batches of E C D coatings were produced and characterized. Each batch consisted of ten 316L stainless steel bare metal stents manufactured by M I V Therapeutics, Vancouver, B C . Each stent was weighted with Sartorius M E 5 micro-balance (accuracy ±1 pg) before and after the E C D process. S E M qualification was performed with stent sample #1, #5, #10 from each batch, coating uniformity and microstructure were closely monitored. Table 2.3-1 shows the summary of the average weight of E C D coating per each batch and their yield rate. Y ie ld rate is defined as the fraction of coating processed within the specification limits. Details of records can be found in Appendix B . The qualification criterion for coating weight (specification) was set to be 62 pg ± 5 pg. A l l S E M observations of coating surface have shown uniform surface coverage and porous microstructure as illustrated in Figure 5.2-20. The average weight of the five batches was found to be 62.6 pg, with standard deviation 2.1 pg, and the average yield rate was found to be 94%. The coating on three stents (ECD-RP-002-08, ECD-RP-003-05, and E C D - RP-004-02) was found to have weighted outside of the acceptable range and was rejected. It was concluded that, in term of coating microstructure, uniformity and weight, the application of E C D - H A coatings on coronary stent exhibits high process reproducibility and consistency. 108 Table 5.2-3. Summary of five batches of E C D coating average weight and yield rate. Coating Weig ht of E C D HA Batch (ug) Stent Number ECD-RP-001 ECD-RP-002 ECD-RP-003 ECD-RP-004 ECD-RP-005 01 62 60 61 64 65 02 63 62 59 23 64 03 66 63 61 63 64 04 65 61 64 67 63 05 64 60 73 65 65 06 60 62 61 62 62 07 61 61 60 61 59 08 62 42 62 63 61 09 66 65 63 61 65 10 65 64 62 66 59 Average (u,g): 63.4 62 0 61.4 63.6 62 7 \ Overall Average (pg): Standard Deviation (u.g): 2 0 1.6 1.4 2.0 2.2 Overall Standard Deviation (u,g): Yield Rate: 100% 90% 90% 90% 100% Overall Yield Rate: > v ' : 94% ''K * Rejected samples w ere excluded from average and standard deviation calculation [ j 5.2.6.1 Errors in ECD-HA Characteristics and Process Parameters Measurement Within the group of 50 stents used for reproducibility study, stents weights were in the range of 15.243 mg - 16.897 mg. According to M I V Therapeutics, the weight variation was attributed to the stent manufacture process, i.e. laser cutting and electropolishing. It was also suggested by M I V Therapeutics that under these manufacture process variations, the surface area of the stent may vary as much as 5%. In consideration of the current density (1 mA/cm 2 ) applied under the optimum coating condition, the surface area variation can impose a 5% error in current density measurement. 109 Electrolyte temperature fluctuation was observed during the process reproducibility study. A two degree Celsius variation was observed while the optimum coating temperature (45°C) was being used. It was believed that the fluctuation was due to the hot-plate on/off cycling. Measurement of electrolyte p H was done with Beckman 260 p H meter with an inhered error of ± 0.004 pH. Chemical preparation for the E C D electrolyte was performed with Sartorius ME235P-SD balance with 0.01 mg accuracy. Time of deposition was controlled by switching the current supply on/off, human error was expected to be ± 2 seconds. Coating weight measurement for stent was performed with Sartorius M E 5 micro-balance with ±1 p.g accuracy. Based on the reproducibility and consistency study of E C D - H A process, the cumulative errors described above were found to have no significant implication on the coating microstructure and coverage. The average weight measurement was found to be 62.6 ug, with standard deviation 2.1 u.g. This measurement was found to be well within the qualification criterion for coating weight (62 ± 5 p,g). 110 6 CONCLUSIONS In this study, electrochemical deposition (ECD) was used to deposit uniform calcium phosphate coatings on 316L stainless steel coronary stents. The influence of the E C D process parameters (deposition time, current density, electrolyte temperature, p H , and Ca/P ratio) on the resulting deposition morphology was investigated. The research results in the following conclusions: 1. Hydroxyapatite coatings can be successfully deposited through E C D on 316L stainless steel coronary stents. 2. The E C D current density and deposition time play an important role on the coating characteristics, in particular the amount and microstructure of coating being deposited. In order to achieve uniform thin film (<0.5 urn) coating on the complex surface of coronary stents, current density should be 1.0 m A / c m and deposition time should be 1 minute. These conditions were determined for the electrolyte concentration of 0.02329M calcium nitrate and 0.04347M ammonium phosphate (Ca/P 1.95), deposition temperature at 45°C, and p H = 4.5. 3. It was observed that the level of current density has a more significant effect on the resulting Ca/P microstructure, whereas the deposition time has more control of weight or thickness of the resulting coating. The coating microstructure exhibits plate-like crystals when high current density (> 3mA/cm ) was applied. The 111 optimum stent coating current density revealed a porous coating with uniform coverage on the complex stent surface. 4. E C D with electrolyte temperature at 45°C exhibits good reproducibility of coating microstructure and thickness (~0.5 urn), as demonstrated through a five batches of ten stents reproduction evaluation. The average weight of the five batches was found to be 62.5 |ug, with standard deviation 2.1 jug, and the average yield rate was found to be 94%. 5. The electrolyte Ca/P ratio (0.49 - 2.92) and p H (4.5 - 5.5) selected for this study have not shown significant impact on the microstructure o f the resulting E C D - H A coatings. 6. Stent surface modification was employed to improve coating adhesion and integrity. The modification involved soaking of stent in ION NaOH( a q ) at 60°C for 24 hours, and subsequently heat-treatment at 500°C for 20 minutes. In-vitro stent crimping and expansion tests found that the application of such substrate surface modification procedure has remarkably improved the adhesion between the E C D coating and the stent substrate, while maintaining the desired coating microstructure and phase composition. 7. X-Ray diffraction studies have confirmed that the as-deposited E C D mixed-phase calcium phosphate coating can be subsequently transformed into pure H A without 112 detrimental effect on the coating microstructure. The post-treatment process involved a 0.1N NaOH( a q ) phase conversion at 75°C for 12 hours and a 500°C heat treatment for 20 minutes. 8. The process parameters for optimized E C D - H A coatings on stents include: • Electrolyte Ca/P ratio Calcium Nitrate [ C a ( N 0 3 ) 2 4 H 2 0 ] Ammonium phosphate [ N H 4 H 2 P 0 4 ] 1.95 0.02329 M 0.04347M • Current Density • Deposition Time • Electrolyte p H • Electrolyte Temperature 1.0 m A / c m 2 1 minute 4.5 45°C 9. The standard 40 mil l ion cycles fatigue test validated the safety and reliability of the optimized E C D - H A coatings with the incorporation of the substrate surface treatment and the phase conversion processes. S E M and E D X analyses of the stent specimens retrieved from fatigue test have shown no sign of cracking or delamination. Filter analysis have further verified that there were no E C D - H A coating debris >0.4um detached from the stent substrate. 113 7 RECOMMENDATIONS FOR F U T U R E W O R K Based on the previous studies done on electrochemical deposition and this current research, the recommendation for future work can be listed as follows: • Study of electrochemical deposition on various substrates. The difference in substrate material and geometry strongly influence the process of deposition, and certainly poses multiple challenges. Application of E C D on other implantable biomedical devices should be highlighted, such as Co-Cr alloys for stents or T i alloys for implants. • Quantification o f the mechanical properties o f E C D coating to allow a more complete understanding of the coating mechanical behavior under the influence of varies E C D parameters. • There are indications that E C D in diluted electrolyte at particularly higher current densities (>10mA/cm ) may yield nanostructure coatings. These conditions should be explored further. • Drug encapsulation in combination with electrochemical deposition technology. Competition to bring to markets polymer-free drug elution stent is vigorous. Investigation of drug encapsulation or impregnation into E C D - H A should focus on maximizing the drug content in the coatings, and optimization of the drug release profile. 114 Electrochemical co-deposition of organo-ceramic coatings. 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Butterworth-Heinemann, Stoneham, M A , 1992. 123 APPENDIX A - PRELIMINARY RESULTS O N E C D CO- DEPOSITION OF O R G A N O - C E R A M I C COATINGS The study of electrochemical co-deposition was not fully complete in this part of the research project, however, preliminary experimental work was performed to investigate the feasibility of such process. The objective of this study was to (i) co- deposit calcium phosphate coating with a polymer for mechanical properties enhancement; (ii) confirmation of the possibility of drug co-deposition technique. Previous study done by Fan et al has demonstrated a uniform collagen fibril/octacalcium phosphate composite coating by electrolytic deposition. Preliminary results indicated that the composite coating may have higher elastic modulus and more scratch resistance go ' than monolithic porous calcium phosphate coating . Polyvinyl alcohol ( P V A ) is a water-soluble polymer, it was chosen as the model reagent for co-deposition. Electrochemical co-deposition was conducted with the optimized process parameters listed in Table 5.2-1, on bare metal stents without the application of substrate surface treatment. Polyvinyl alcohol was dissolved in the standard E C D electrolyte before deposition. Two different electrolytes were prepared with 0.1 wt% P V A and 0.8wt% P V A . Figure A 1 and Figure A 3 illustrate the resulting E C D co-deposited coatings on coronary stents with 0.1 wt% and 0.8wt% P V A , respectively. Both of these coatings demonstrated good uniformity and full coverage. It was obvious that the co-deposition with 0.8wt% P V A exhibits a denser structure compared to the 0.1 wt% P V A . Figure A 2 124 and Figure A 4 illustrate the stent expansion test results of the 0.1 wt% and 0.8wt% P V A , respectively. Even though both substrates have not been modified for coating adhesion enhancement, the stent expansion results revealed dramatic improvement over E C D coating of CaP alone on bare metal stent as illustrated in Figure 5.2-10. These results confirmed the possibility of E C D co-deposition with P V A for coating integrity enhancement. Although the fundamental mechanism of co-deposition was not clear at this stage, it was reasonable to believe that the dissolved P V A was encapsulated while calcium phosphate was being deposited. 1 1 ^ -1 A . Hum 1. .OOkV x900 iUum Figure A 1. S E M images of an E C D coating co-deposited with 0.1 wt% of P V A on coronary stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [xl0,000] 125 Figure A 2. S E M images of an expanded stent coated with E C D co-deposited with 0.1 wt% P V A . (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] Figure A 3. S E M images of an E C D coating co-deposited with 0.8wt% P V A on coronary stent, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [xl0,000] 126 Figure A 4. S E M images of an expanded stent coated with E C D co-deposited with 0.8wt% P V A . (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000] 127 APPENDIX B D E T A I L E D RECORDS OF REPRODUCIBILITY STUDY Table B 1. Preparation record for E C D coating batch E C D - R P - 0 0 1 . Substrate Marks & Identity: Rcf No. Size: Description: Preparation Date: Batch Number Prepared By: Electrolyte Ca:P Ratio: Applied Current (mA): Electrolyte Temp. ("C): Electrolyte pH: Deposition Time (min): Surface Treatment Phase Conversion: Sintering Temp. ("C): ECDCqating Yes 500 °C T 2 0"min Sample Name X D - R P - O O I - O I - JD-RP-COi-O/ ]CD-RP-.001-03: C O - R P 0 0 1 3-1 : D - R P - 0 0 1 - 0 5 : icD-RP-oor-O'e' -JD-RP-001-07 E C D - R P - 0 0 1 - 0 8 - .ECD-RP^001'-09: ECDr-RP-001-10 Average: Bare Stent Weight (mg) : 5. SOT- I S . 7 1 6 - : 1 5 . 7 . 4 5 - 1 6 . 0 8 7 1 6 . 1 8 7 1 6 . 2 1 9 1 5 : . 7 8 9 : 1 5 - . ; 8 5 2 - 1 6 . 7 1 6 1 5 . 7 9 8 15.971 Final Stent Weight (mg) 15:. 6,63- 15.779 15.811:-. 16:i52" 16.251. •16.279 15.850 15.914 16.782 15.863 16.034 Coating Weight (ug) 63 66 65' 64 60 61 62'- 66. 65 63.4 Pass/Fail Pass Pass Pass Pass Pass; Pass- Pass Pass Pass Standard Dev. 0.317 Yield Rate of ECD-RP-001 -100' 128 Table B 2. Preparation record for E C D coating batch ECD-RP-002 . Substrate Marks & Identity: r Ref No. Size: Description: MIVI 700 S e r i e s C o r o n a r y S t e n t ECD Coating Preparation Date: J a n u a r y 23, 2006 "l^^^^^^^Bli^BSiB1!^!! Batch Nurrben E C D - R P - OO'™"" ' [.. ' Prepared By: Electrolyte Ca:P Ratio: Appliucl Cum-nt (mA): bl--ctn*kte Temp. (°C): Eli'utr»Kti.' pH: Deposition Time (mm) Surface Treatment Pli.i-.t- Cunvuisiuii: Yes* Sintering Temp. ("C): 500 C {20 min) y, •• • /",• • < Sample Name ECD-RP-002-0:l Bare Stent Weight (mg) 15.768 Final Stent Weight (mg) ~15 . 8 2 F "~ Coating We ( U C j l ight 60 Pass/Fail P a s s ;ECD-R-P-002-02 15.785 15.847 62- •ECD-RP-002-03 15.743 15.806 '63 Pass :ECD-RP-002-04 16.220 16.281 61 Pass ECD-RP-002-05 16.735: 16.7 95 60 ECD-RP-002-06 •16.234. 16.296 : "62 Pass ECD^RP-002-07- 15 . 982' 16.0,43 Pass ECD-RP-002- F a i l P a s s Pass Table B 3. Preparation record for E C D coating batch ECD-RP-003 . Substr.itt- Marks & Identity: Ref No. Size: Description: MIV I ~7 0^"^Tr i e7"F<^onl ECDL Coating Pri'|j;iratiu:i Ddtu: January 31, 2006 Batch Number ECD-RP-003 liiliM^iiiiiiMli^^^B^i^^W^i^^psi^^S Pn-piin-d B>: Electrolyte Ca:P Ratio: Applied Current (mA): Electrolyte Temp. (°C): 45°c ± 2 °c ^^^^^^^^^JlBll^Bifl^BSi^HSl El'ictrolvt'- pH: Deposition Time (mm): l T o ^ ^ ^ ^ ^ ^ ^ ^ ^ ^ ^ t l i l l l i l l t i r i l ^ ^ Surface Treatment Phase Conversion: Sintering Temp. ("C): " Sample Name B.iro Stunt Weight (mg) Final Stent Weight (mg) Coating Woight (ug) Pass/Fail ECD-RP-003-01 16.024 16.085 61 Pass ECD-RP-003-02 15.655 15.714 59 Pass ECD-RP-003-03 15 . 979 16.040 61 Pass ••- ECD-RP-003-04 16.246 16.310 64 Pass te F.C:; ? . ? 0 0 3 Ob . 16 897 16.970 73 F a i l ECD-RP-O.03:-O:6:: •• • 15.786 15.847 61 -Pass . • ECD-RP-003^.0-7 • -1-5-.-2 43- - •• ; :5.303 60 Pass '. • ECD-RP-003-08 16.099 16 161 62 "PaSS': •:: Pi :' ECD-RP-003'-0;9 16 565 16 628 63 .Pass.. Pi ECD-RP-003-10 35.662 • 15.724" - : : :62 • Pass Average: 16.016 16.07B Standard Dev. 0". 452 ' ' V 0.455 Yield Rate of ECD-RP-003=90°l Table B 4. Preparation record for E C D coating batch ECD-RP-004 . IF Substr.ito M i r l n <K lil>-nllt\ I. <: Rtf Nn 14 mm Description: MIVI 700 Series Coronary Stent -:&i ECD Coatinq Preparation Date: February 8, 2006 Batch Number Prepared By: b"l-::trolyte Ca:P Ratio: Applied Current (niA): Electrolyte Temp. (°C): 45°C ± 2 °C ^^^^^^^^^^^^^^^^^^^^^^ Electrolyte pH: 4 .5 ^^^^^^ îp^^^H î̂ î ^Bî ^^S^^^KS Deposition Time (min): Surface Treatment Ph.ise Conversion: Sintering Temp. (°C): "500 "°c"720 min) • • ' Sample Name Bjrr> Stent Weight (mg) Final Stent Weight (mg) Coating Weight (ug) Pass/Fail ECD-RP-004-01 15.277 15.341 64 :" ; D - R P - O O 4 - 0 2 15 .'349- ' Fa J. 1 .CD-RP-004-03 • 16.021 16.084 . 63 ••' . ' . Pass ! CD-RP-0.04:-0,4.' 16.730 16.797 67 Pass •••CD-RP-004.̂ 05 15.97 6 ' 16.041 65 Pans - ED-RP-0.04-0:6 15.991 16.053 62 Pass • CD-RP-004-07 16.767' ' • = 16.828 61 Pass •.CD-RP-004-08 16.765 16.828 63 Pass ECD'-RP-0:04'-0:9: - 15.429 15.490 61 Pass ECD-RP-004-10: 15.657 :. :i:5:. : • 66 Average 15 996 ..16.056 63. 6 Standard Dev. 0.557 0.561 2.0 Yield'Rate of ECD-RP-004 ' =90 Table B 5. Preparation record for E C D coating batch ECD-RP-005 . Substrate Marks & Identity: : i Ref No. i .1 Size: Description: MIVI 700 Series Coronary Ster ECD Coatinq Preparation Date: February 13, 2 0 t J ^ ^ ( i | l | ! | f l l l^^Bl^B^BilH Batch Number Prepared By: Electrolyte Ca:P Ratio: Applied Current (mA): bl-ctfolyte Temp. (°C): 4 5°C ± 2 °C ...JIIRIM Electrolyte pH: Deposition Time (min): Surface Treatment Phase Conversion: Sintering Temp. (°C): Sample Name Bare Stent Weight (mg) Final Stent Weight (mg) Coating Weight (ug) Pass/Fail 1 ECD-RP-005-01 '. 15.767 15.832 65 Pass ECD-RP-005-02 16.341 16.405 64 Pass ECD-RP-005-03 16.792 16.856 64 Pass | ECD-RP-005-0;4:: 16.139 16.202 63 Pass . ECD-RP-00!5"-05 15.672 15.737 65 Pass ECD-RP-005-06 15.887 15.949 62 Pass ECD-RP-005-07 16.099 16.158 59 Pass | ECD-RP-005-08 16.754 16.815 61 Pass ECD-RP-005-09 15.780 15.845 65 Pass • iEGD-RP-005-10 • 16.024 16.083 59' Pass Average: 16.126 16.188 62.7 Standard Dev. 0.375 0.375 2.2 Yield Rate of ECD-RP-005 =100%

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