Open Collections

UBC Theses and Dissertations

UBC Theses Logo

UBC Theses and Dissertations

Calcium phosphate coatings on coronary stents by electrochemical deposition Tsui, Manus Pui-Hung 2006

You don't seem to have a PDF reader installed, try download the pdf

Item Metadata

Download

Media
[if-you-see-this-DO-NOT-CLICK]
ubc_2006-0741.pdf [ 22.71MB ]
[if-you-see-this-DO-NOT-CLICK]
[if-you-see-this-DO-NOT-CLICK]
Metadata
JSON: 1.0078873.json
JSON-LD: 1.0078873+ld.json
RDF/XML (Pretty): 1.0078873.xml
RDF/JSON: 1.0078873+rdf.json
Turtle: 1.0078873+rdf-turtle.txt
N-Triples: 1.0078873+rdf-ntriples.txt
Original Record: 1.0078873 +original-record.json
Full Text
1.0078873.txt
Citation
1.0078873.ris

Full Text

C A L C I U M P H O S P H A T E COATINGS O N C O R O N A R Y STENTS B Y 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 o f Metals and Materials Engineering, The University o f British Columbia, Vancouver, B . C . , Canada, 2004  A THESIS S U B M I T T E D IN P A R T I A L F U L F I L L M E N T OF THE R E Q U I R E M E N T FOR THE D E G R E E OF  M A S T E R OF 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 STUDIES  (Materials Engineering)  T H E U N I V E R S I T Y OF BRITISH C O L U M B I A October 2006  © Manus Pui-Hung Tsui, 2006  ABSTRACT Calcium phosphate ceramic coatings, especially hydroxyapatite attracted much attention  i n the orthopedics  and dentistry  ( H A ) , have  due to their  excellent  biocompatibility and bioactivity. A m o n g the different methods o f calcium phosphate coatings processing, electrochemical deposition ( E C D ) 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 o f the E C D process parameters (deposition time, current density, electrolyte temperature, p H , and investigated.  Ca/P ratio)  on the  resulting deposition morphology  was  Scanning electron microscopy ( S E M ) 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 - 2 H 0 , D C P D ) and hydroxyapatite ( C a ( P O ) ( O H ) , H A ) 4  2  10  were present in the uniform -0.5 um thick as-deposited coating.  4  2  However, a,post- -  treatment process, including a 0.1N N a O H ( ) phase conversion at 75°C and a 500°C heat aq  treatment produced a pure phase H A coating. The final deposit revealed a highly porous surface morphology which could be useful for drug encapsulation. W i t h 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 m i l l i o n 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.  ii  T A B L E OF CONTENTS ABSTRACT  ii  T A B L E OF C O N T E N T S  iii  LIST OF T A B L E S  vi  LIST OF FIGURES  vii  ACKNOWLEDGMENTS  xiii  1  2  INTRODUCTION  1  1.1  Biomaterial Coatings for Stents  2  1.2  Calcium Phosphates and Hydroxyapatite  3  1.3  Electrochemical Deposition o f H A Coatings  5  1.4  Motivation and Focus o f the Present Study  6  LITERATURE REVIEW 2.1  Coronary Heart Disease ( C H D )  2.2  Coronary Artery Stent  8 8 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  2.3.1  Calcium Phosphate Bioceramics  19 22  2.3.1.1  Bioresorption and Biodegradation  23  2.3.1.2  Mechanical Properties  25  2.3.2  Hydroxyapatite Bioceramics  27  iii  2.4  4  Thermal Stability  28  2.3.2.2  Mechanical Properties  29  Biomaterial Coatings  30  2.4.1  Hydroxyapatite Coatings  31  2.4.2  Processing o f Hydroxyapatite Coatings  32  2.5  3  2.3.2.1  Electrochemical Deposition  34  2.5.1  Electrochemical Deposition o f Hydroxyapatite Coatings  36  2.5.2  Microstructure and Phase o f E C D Calcium Phosphate Coatings  40  2.5.3  Adhesion o f H A Coatings  43  SCOPE A N D OBJECTIVES  46  3.1  Scope o f the Investigation  46  3.2  Objectives  47  EXPERIMENTAL METHODOLOGY 4.1  Sample Preparation  4.1.1 4.2  Substrate Surface Modification  Electrochemical Deposition  49 50 51 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 5  Fatigue Test  58  RESULTS A N D DISCUSSION 5.1  61  E C D o f 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 o f Electrolyte p H  72  5.2  E C D o f 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 o f E C D Calcium Phosphates Coatings  90  5.2.5  In-vitro Fatigue Test  98  5.2.6  Reproducibility and Consistency o f E C D - H A Process  5.2.6.1  Errors  in  ECD-HA  Characteristics  and  108 Process  Parameters  Measurement  109  6  CONCLUSIONS  '.  Ill  7  RECOMMENDATIONS FOR FUTURE WORK  114  REFERENCES APPENDIX  A  116 -  PRELIMINARY  RESULTS  O N E C D CO-DEPOSITION  OF  ORGANO-CERAMIC COATINGS  124  APPENDIX B  128  D E T A I L E D RECORDS OF REPRODUCIBILITY S T U D Y  v  LIST O F T A B L E S Table 2.3-1. Various calcium phosphate bioceramics  23  Table 2.3-2. Mechanical and Physical Properties o f Calcium Phosphates  26  Table 2.4-1. Purposes o f biomaterial coatings  30  Table 2.4-2. Summary o f techniques used for deposition o f hydroxyapatite coatings..... 34 Table 4.1-1. Nominal chemical composition o f 316L stainless steel  50  Table 5.1-1. The concentration o f C a ( N 0 ) 4 H 0 and N H H P 0 for the preparation o f 3  2  2  4  2  4  different Ca/P ratio electrolytes  69  Table 5.2-1. Optimum parameters for E C D o f calcium phosphate coatings* Table 5.2-2. Summary o f E C D coatings dissolution test data.  75  Dissolution tests were  conducted with 10 m L o f phosphate buffer saline ( P B S ) at 37°C (pH = 7.4) with rotation speed at 20 rpm  97  Table 5.2-3. Summary o f five batches o f E C D coating average weight and yield rate. 109 Table B 1. Preparation record for E C D coating batch E C D - R P - 0 0 1  128  Table B 2. Preparation record for E C D coating batch E C D - R P - 0 0 2  129  Table B 3. Preparation record for E C D coating batch E C D - R P - 0 0 3  ..130  Table B 4. Preparation record for E C D .coating batch E C D - R P - 0 0 4  131  Table B 5. Preparation record for E C D coating batch E C D - R P - 0 0 5  132  vi  LIST O F FIGURES Figure 2.1-1. Fat and cholesterol accumulated oh the inside o f coronary arteries  9  Figure 2.1-2. Schematic o f percutaneous transluminal coronary angioplasty technique. A guide-wire is placed across the blocked section o f 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 o f percutaneous transluminal coronary angioplasty technique with the use o f 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 o f the product o f calcium and phosphate concentrations plotted against p H values o f solution saturated with respect to various calcium phosphate phases in the ternary system Ca(OH)2-H3P04-H20. Calculated for 37°C  Figure 2.5-1.  25  Cathodic polarization curve o f T i substrate i n a Ca(N03)2 4 H 2 O and NH4H2PO4  electrolyte  39  Figure 2.5-2. X R D analysis o f coatings deposited at current density o f 1, 5, 10, 15, and 20 43  m A / c m 2 for 30 min Figure 4.1-1. Schematic diagram o f M I V I 700 Series Coronary Stent  51  Figure 4.2-1. Schematic diagram o f electrochemical deposition setup  53  Figure 5.1-1. S E M images o f E C D coating deposited at various current densities: (a) 15 m A / c m , (b) 10 m A / c m , (c) 5 m A / c m , (d) 3 m A / c m , (e) 1 m A / c m 2  2  [Left: xlOO; Right: x 1,500]  2  2  2  63  vii  Figure 5.1-2. Weight gain o f E C D coated specimens versus current density with 5 minutes o f deposition  64  Figure 5.1-3. S E M images o f E C D coating deposited with various deposition time: (a) 15 min, (b) 10 min, (c) 5 min, (d) 3 min, (e) 1 m i n x l 5 0 0 ] at 1 m A / c m  [Left: xlOO; Right: 65  2  Figure 5.1-4. H i g h magnification S E M image o f E C D coating with deposition o f 1 min [x20,000] at I m A / c m  2  67  Figure 5.1-5. Weight gain o f E C D coated specimens versus deposition time deposited at 1 mA/cm  2  67  Figure 5.1-6. Ca/P ratio o f resulting deposit with the use o f various Ca/P ratio electrolytes. Ca/P ratio was derived from E D X spectra  69  Figure 5.1-7. S E M images o f 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 o f resulting CaP deposits conducted i n electrolyte with temperature a) 25 °C b) 45°C, c) 75°C. [x 15,000] Figure 5.1-9.  71  Influence o f electrolyte temperature on measured supply voltage, for constant current source (I = 13.77 m A )  72  Figure 5.1-10. S E M Image o f E C D conducted with p H 3.0 electrolyte at 45°C. [x5,000] 73 Figure 5.1-11. S E M images o f 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 o f bare metal stent with various magnification: a) [xlOO], b) [x800],c) [x 1,500]  76 viii  Figure 5.2-2. S E M images o f 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 o f 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 o f 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 o f 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 o f 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 o f the E C D coating deposited with optimum parameters (Table 5.2-1), for 1 minute, showing a mixed phase o f D C P D and H A . . . 80 Figure 5.2-8. Cross-section S E M image o f E C D coating deposited on stent. Estimated coating thickness was approximately 0.5 urn Figure 5.2-9.  81  (a) S E M images o f 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 o f 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] Figure 5.2-11.  83  Compressive spallation by buckling showing localized interfacial decohesion  Figure 5.2-12.  83  Tensile stress in brittle film causing through-thickness cracking and interfacial delamination  84  Figure 5.2-13. S E M image o f 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 o f a surface modified stent Figure 5.2-15.  86  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 Figure 5.2-16.  86  S E M images o f 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 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 m m diameter catheter, (a) [xlOO], (b) [x800], (c) [x3,000], and (d) [x 10,000]  88  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 m m 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 o f as-deposited E C D coating and the resulting coating after 12, 24, 48, and 72 hours o f N a O H ( ) phase conversion at aq  75°C  92  Figure 5.2-20. S E M images o f 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 o f 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 o f 12 hours N a O H ( ) phase converted E C D aq  coating and after 300°C, 500°C, and 750°C for 20 minutes o f heat treatment o f the coating  95  Figure 5.2-23. S E M images o f 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( ) treatment + 750°C heat treatment.. 97 aq  Figure 5.2-24. S E M images o f expansion test result from an as deposited E C D coating upon N a O H ( ) treatment + 500°C heat treatment. aq  Expansion performed  with a 3.5 m m 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] Figure 5.2-25.  98  Summary o f average % O D strain for the six fatigue tested stent specimens  99  xi  Figure 5.2-26. S E M images o f explanted fatigue tested specimen I P 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 o f 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 Figure 5.2-28.  102  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  103  Figure 5.2-29. The six main categories o f debris found on filter after the fatigue test.. 106 Figure 5.2-30. E D X analysis o f the six main categories o f debris found after the fatigue test  107  Figure A 1. S E M images o f an E C D coating co-deposited with 0.1 wt% o f 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 o f 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 o f 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] Figure A 4.  126  S E M images o f 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  xii  ACKNOWLEDGMENTS I wish to express my sincere gratitude to the many people who have kindly provided me with great contribution to the completion o f this thesis.  I owe particular  thanks to my supervisor, D r T o m 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 o f Materials Engineering for all their support. A l s o , 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 o f education.  xiii  1  INTRODUCTION Coronary artery disease ( C A D ) is the leading cause o f death in North America. In  2004, approximately 54% o f all cardiovascular deaths are due to coronary artery disease . 1  C A D occurs when fat (cholesterol) deposit blocks the arteries, reducing the oxygen supply to the heart muscle.  Angioplasty is a technique o f 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 o f C A D . A stent is a stainless steel or C o - C r 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 o f blockage. When the balloon is inflated, the stent expands, locks i n place and forms a scaffold.  Stent stays i n 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 o f blockage (restenosis) in the artery in all cases. The bare metal stent may trigger a natural inflammatory response resulting in proliferation o f smooth muscle cells , thus, narrowing or re-closing o f the artery, which often requires a 2  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. produce different symptoms depending on where it forms.  A thrombus w i l l  A coronary thrombus w i 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 ( T I A ) 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 o f coated coronary stents should be less restenotic,  thrombogenic, be more acceptable by body environment, and should be capable o f locally deliver drug to the surrounding tissue.  1.1  Biomaterial Coatings for Stents A biomaterial is any synthetic material used to replace part o f a living system or to  function in intimate contact with living tissues . The ability o f such material, device or 6  system to perform without a clinically significant host response i n a specific application is defined as biocompatibility.  The performance o f a biomedical device is greatly  depended on the material properties, design o f the device, surgical techniques, health o f patient, and the surface properties.  Because of the complex interface o f biomaterial-  biological medium, surface properties is often the key to determine success o f 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 o f the biomaterial to a specific application without detrimentally affecting its bulk properties.  Ultimately, biomaterial  coatings combine the advantages o f 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  C a l c i u m Phosphates and Hydroxyapatite Calcium phosphates have been used since 1969 for manufacturing various forms o f  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 temperature . 6  Calcium phosphate ceramics, particularly  hydroxyapatite ( H A ) , have received much attention and clinically applied in orthopedics and dental industries due to their excellent biocompatibility " . H A is recognized as 7  9  osteoconductive and is able to accelerate bone ingrowth and attachment to the surface o f implant during the early stages after implantation ' . 8  9  In addition, it can induce tissue  ingrowth and formation o f chemical bonding to achieve better implant fixation.  HA 3  interacts with the biological environment through a complex dissolution, precipitation, and ion exchange process ' 11  1 2  .  Following the introduction o f H A to the biological  environment, a partial dissolution o f the surface is initiated causing the release o f C a , HPO4 " and PO4 ", increasing the supersaturation o f the micro-environment with respect 2  3  to the stable ( H A ) phase. carbonate  This saturation leads back to the precipitation hydroxyl-  apatite ( H C A ) layer.  The polycrystalline H C A phase is equivalent in  composition and structure to the mineral phase o f bone ' 10  H A coating can prevent the  fibrous tissue encapsulation o f implants and reduce the potentially harmful release o f metallic ion from the implant into the biological environment . Therefore, H A coatings 12  on metallic substrate were applied to achieve better biocompatibility.  Hydroxyapatite Caio(P0 )(OH)2 is chemically similar to the mineral component o f 4  bones and hard tissues i n mammals. Hydroxyapatite crystallizes into hexagonal crystal structure and has a unit cell dimension a = 9.432 A and c = 6.881 A . H A is 10:6 and the density is 3.219 g/cm  3  [ 6  \  The Ca/P ratio o f  The mechanical properties o f synthetic  calcium phosphate vary widely as a function o f their crystallinity and porosity. Limitations o f H A application relate to its brittleness ( K i c < 0.9 M P a - m  ) and low  strength (generally, compressive strength <300 M P a , tensile < 50 M P a )  Therefore,  [ 1 3 ]  .  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 o f  orthopedics, there is accumulating evidence that H A performs equally well in blood circulation environment . 14  4  1.3  Electrochemical Deposition of HA Coatings Different deposition technologies o f H A have been investigated i n the past  including;  plasma  electrophoretic  spray " , 15  deposition , 21  17  sol-gel ,  and  pulse  18  electrochemical  laser  deposition ,  sputtering ,  19  deposition " . 22  20  Electrochemical  24  deposition ( E C D ) o f calcium phosphate coatings at room (or near-room) temperature has been investigated since the 1990s. surfaces  E C D is particularly attractive for coating internal  o f porous metallic substrates or highly complex surface.  The resulting  properties o f H A can be controlled in E C D by regulating the electrochemical potential, current density, electrolyte concentration, and temperature o f the process, as these conditions directly determine the deposition rate, morphology, and chemical composition of the H A coatings . 25  In the cathodic electrochemical deposition method, cathodic reactions are used to generate O H " groups and increase p H at the electrode. Metal ions or complexes, which are stable in the bulk solution at low p H , are hydrolyzed by electro generated base at the electrode surface to form colloidal particles ' . 23  26  A l l applications o f 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 i n this respect as there is a significant strain o f stent during expansion, as large as 15% .  Due to brittleness o f 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 o f 5  separated coating have the potential to clog up blood supply causing fatal  affect.  Therefore, one o f the crucial factors for E C D - H A film 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 o f the stent.  Bio-resorption or dissolution is a major concern o f i n the biological environment. In terms o f 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 o f 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  M o t i v a t i o n 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. temperature, current  The E C D parameters, such as deposition  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 o f 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 o f coating adhesion. The coatings phase composition and crystallinity are  investigated  and  conversion process.  optimized through  introduction o f a post-deposition  phase  Mechanical behavior o f the coatings is characterized qualitatively  through observation o f the coatings behavior on expanded stents. H A was applied on coronary stents followed by a simulated stent implantation procedure to ensure high reliability o f coating, i.e. no separation, delamination or visible damage.  Finally,  performance o f 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  2.1  LITERATURE REVIEW  C o r o n a r y H e a r t Disease ( C H D ) Coronary heart disease ( C H D ) , also called coronary artery disease ( C A D ) , is the  result of the atheromatous plaque accumulation, which composed o f cholesterol, fatty compounds, and a blood-clotting material (fibrin), within the walls o f the arteries that supply blood to the myocardium- the middle layer o f the heart ' 31  illustrates accumulated atheromatous  3 2  .  plaques on the inside o f arteries.  Figure 2.1-1 While the  atheromatous plaques initially expand into the arteries walls, they w i l l finally expand into the lumen o f the vessel, affecting the blood flow through the arteries. atheromatous  When  plaques obstruct less than 70 percent o f the diameter o f the vessel,  symptoms o f obstructive C A D are rarely observed.  However, as the plaques obstruct  more than 70 percent o f the diameter o f the vessel, symptoms o f obstructive C A D w i l l starts to develop .  A t this stage o f the obstruction, the symptoms o f ischemic heart  disease are often first noted during increased workload o f the heart. A s the obstruction develops, there may be near complete blockage o f the lumen o f the coronary artery, restricting the flow o f 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 o f chronic coronary ischemia, including symptoms o f angina at rest and flash pulmonary edema . 1  8  Figure 2.1-1. Fat and cholesterol accumulated on the inside o f coronary arteries.  Coronary artery disease is the leading cause o f death i n North America. In 2004 approximately 54% o f all cardiovascular deaths were due to coronary artery disease . 1  Due to the varying severity levels o f 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 o f symptoms or to reverse atherosclerosis; (2) to relieve symptoms; and (3) to lower an individual's risk o f 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 o f 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 o f the narrowing, number o f arteries being affected, the location o f the narrowing, and individuals' factors, such as age and general health.  Percutaneous transluminal coronary angioplasty is a technique o f 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 o f a coronary artery stent together with balloon catheter i n P T C A has been significantly increased during the past 10 years.  Figure 2.1-3 illustrates the use o f 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% o f the 20,000 P T C A procedures in United K i n g d o m . 34  10  Figure 2.1-2. Schematic o f 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 o f percutaneous transluminal coronary angioplasty technique with the use o f 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 o f coronary arties. Stent is commonly mounted on a balloon catheter as a single unit, with the use o f fluoroscopic screening and a radiopaque marker, the unit is carefully positioned across the site o f lesion. Inflation o f the balloon catheter results i n expansion and deployment o f the stent circumferentially to the surface o f 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 o f a n i m a l . In 1983, 35  Dotter et al and Gragg et al implanted the first "stent-like device", a wire'spring o f nitinol, into the canine coronary arteries with a catheter . In 1985, Palmaz et al reported 36  the introduction o f 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 o f these  devices are implanted i n the arteries. It was reported that 3 5 % o f the animals exhibited partial or total thrombotic occlusion .  It was in 1986 that the first implantation o f  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 o f 229 successful deliveries of intracoronary stent placements in 230 lesions in 213 patients . It was reported that the 37  subacute thrombosis rate reached 14% and 40% o f restenosis was determined in the following 6 months. The high level o f thrombosis rate had convinced cardiologists that  12  stent, as a foreign object, when implanted w i l l cause high thrombogenicity. In the early 1990s, the Belgian-Netherlands Stent study ( B E N E S T E N T - 1 ) Study ( S T R E S S )  39  3 8  and Stent Restenosis  were carried out to compare the conventional coronary balloon  angioplasty with stent implantation.  The results have shown beneficial effects o f  coronary stenting following balloon angioplasty, in comparison to angioplasty procedure alone.  It was reported that among the 407 patients in the studies, the incidence o f  restenosis was significantly lower after stent implantation ( B E N E S T E N T 22%, S T R E S S 32%) when compared with after balloon angioplasty alone ( B E N E S T E N T 32%, S T R E S S 42%) ' . 3 8  The beneficial effect o f coronary stenting has been attributed to the larger  3 9  37  •  acute lumen dimensions and to the prevention o f constrictive remodeling . It was until 1994, that the first F D A approval o f stent device was issued to Palmaz-Schatz stent.  Many pioneer researches have underlined major problems with the use o f 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 o f the vessel lumen diameter by at least 50% at the site o f angioplasty procedure, is the other problem. Nevertheless, the combination o f the technical advances and previous research data helped to define the ideal characters o f today's stent as follows: ' 36  37  •  Flexible  •  L o w unconstrained profile  •  Radio-opaque  •  Non-Thrombogenic  13  •  Biocompatible  •  Reliable  •  H i g h radial strength  •  Circumferential coverage  •  L o w surface area  Alloys routinely used for the manufacture o f 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 o f 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 i n the artery. Nitinol or N i T i alloy consists o f - 5 5 % N i c k e l and - 4 5 % titanium. N i t i n o l is gaining popularity in the design o f stent due to its shape memory capabilities.  Under a  mechanical load, nitinol can deform reversibly up to 10% o f strain by the creation o f a in  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 Nitinol have limited its applications ' . Tantalum is another option for materials in stent manufacture, though questions have been raised because o f its exaggerated radiopacity ' . 37  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 ' . 37  40  Although coronary artery stenting results in a larger acute lumen and prevents elastic retraction o f 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 o f an injured vessel.  The complex restenosis  mechanism following the stent implantation is due to the f o l l o w i n g ' : 38  41  •  The inner elastic membrane induced by injuries.  •  The secretion o f mitotic substances and growth factors.  •  The prolonged and continuous stress created by the stent.  •  The chronic irritation targeted by the presence o f a foreign object.  In 2002, the worldwide market o f coronary and peripheral intervention was about U . S . $5 billion 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 ' 42  4 3  , hydrogen-rich amorphous silicon carbide ' 44  4 5  , amorphous titanium  oxide , and g o l d . Some o f these inert or tissue friendly coatings do not cause platelet 46  47  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 o f metallic ions from the stainless steel to the surrounding tissue . 43  The inertness o f gold  was investigated by Kastrati et al with the expectation o f increasing the stent biocompatibility and radiopacity. However, a randomized trial has suggested the gold coated stents exhibit an increased possibility o f restenosis when compared with stainless steel stents . It was further reported that the undesired result may be related to a coating 47  defect rather than the gold itself.  This demonstrates the importance o f 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 o f the  16  commercialized heparin stents, B X Velocity by Johnson & Johnson, immobilizes heparin on the stent surface and remains free to inhibit thrombus formation . 49  2.2.2  Biodegradable Stents One o f 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 o f a metallic stent may trigger chronic inflammation and i n turn stimulate restenosis.  T o avoid the  latter problem, a biodegradable stent has been proposed. A biodegradable stent exhibits similar mechanical functionality o f a metallic stent but w i l l then slowly degrade as the artery become stable. poly-L-lactase . 50  Stackle et al were the first to develop biodegradable stents with  In 2000, Tamai et al have reported the results o f the first 6-month  clinical trial o f 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% . 51  2.2.3  Radioactive Stents It was demonstrated that a l o w dose o f radiation can effectively inhibit or  decelerate the proliferation o f vegetative c e l l s  3 7 , 4 9  ' . It is believed that this phenomenon 5 2  can lead to a reduction o f intimal hyperplasia and as a result decreases the chance o f restenosis.  Herlerin et al have reported the use o f radioactive stent i n an animal study . 52  The radioactive stents used are made o f steel bombarded with C o , 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 o f hyperplasia . This demonstrates the complexity between radiation and vessel response.  While the therapeutic  and toxic limits still have not been  determined  accurately, the rate o f restenosis at the edges o f radiated area remains a major problem . 49  In general, there still exist unsolved and unknown problem with the use o f radiation in form o f short therapy and radioactive stents.  2.2.4  Drug Eluting Stents Drug eluting stents ( D E S ) that can efficiently deliver biologically active agents  locally to the vascular wall have been under intensive research i n recent years.  Early  results with drug eluting stents were often unsatisfactory due to uncertain require dosage, uncertain coating . 37  delivery duration, poor  drug efficacy, and  proinflammatory  polymeric  More recent studies on D E S with the use o f anti-proliferation drugs have  generated excellent results. anti-proliferation  drugs.  Sirolimus and Paclitaxel are the two most commonly used Sirolimus  is  a  macrolide  antibiotic  that  possesses  immunosuppressant activity, it is also known to be an inhibitor o f smooth muscle cells ( S M C ) proliferation and migration . 53  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 o f the cell cytoskeleton structure . 54  Other anti-proliferation drugs under investigation includes:  everolimus, tacrolimus, estradiol, dexamethasone, and angiopeptin.  18  Two o f 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 o f these D E S use polymer as a vehicle to carry the drug eluting into surrounding tissue through diffusion and convection . 55  While concerns have been raised with  polymer coating, non-polymeric stents A C H I E V E ™ coated with paclitaxel by C o o k ® is under investigation.  However, clinical studies have been inconclusive. The question  between the potential long term degradation and inflammatory effect o f polymer coating and control release rate o f non-polymeric coating are still i n debate. Although there can be no doubt that the development and implementation o f drug-eluting stent in past clinical studies has been a major milestone that w i l l evolve the approach to the management o f coronary artery disease, yet, it is also certain that more data are required to fully understand the scope o f its benefits and to identify its limitations. In spite o f the different drugs being investigated and their different therapeutic effects, the technique o f coating, and drug delivery platform are still the key component leading to the success o f 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  hybrid o f ionic and covalent bonds.  Ceramic bonding is generally a  The strong bonding forces o f ceramic materials  result in properties such as high elastic modulus and hardness, high melting points, low thermal expansion, and good chemical resistance . 6  However, the bonding nature and  minimum number o f slip system o f 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 o f 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 o f ceramics have been realized, and these "bioceramics" are being widely used in orthopedics and dental industries. Bioceramics can be manufactured i n 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 . Bio-inert ceramics are inert in 6  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 . Alumina has been used i n orthopedics and 57  dental surgery for almost 20 years, it properties include high hardness, inertness, l o w CO  coefficient o f friction and wear resistance candidate for joint replacement.  .  These have made alumina as the ideal  Medical grade zirconia has been developed for use i n 20  total joint prostheses because o f its high fracture toughness and tensile strength.  Such  improved properties o f zirconia have enable manufacture o f femoral head into smaller diameter than present generation o f alumina. Pyrolytic carbon is commonly used in coating o f artificial heart valves for the last 30 years . Properties that make this material 59  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 o f bio-resorbable ceramics varies  among different materials. Most o f the bio-resorbable ceramics are i n forms o f calcium phosphate.  Due to the resorbable nature o f most calcium phosphate, they have been  often used as potential bone defect fillers.  In such application, the calcium phosphate  filler would fill the void and gradually dissolve away, being replaced by host tissues. Two critical issues with the development o f bio-resorbable ceramics are: maintenance o f stability and strength during the degradation period, and matching resorption rate to the natural repair rates o f the body tissues. Tricalcium phosphate is one example; it has been used as temporary bone substitute , where it has to maintain sufficient strength at first to 60  allow bone ingrowth, then eventually degrade and be replaced by endogenous bone. Other uses o f 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 o f bone infection.  It was reported that  TCP/cysteine composite releases the antibiotics over a period o f 3 weeks at the site o f  21  infection.  The authors suggested that antibiotics released from T C P amino acid  composites effectively utilized in the treatment o f bone infection . 61  Bio-active ceramics are surface reactive and bond to adjacent tissue some time after implantation . 12  In many cases, the interfacial bond strength is equivalent to or greater  than the cohesive strength o f the implant material. Bioactive glass and hydroxyapatite are  examples  phosphates,  o f bio-active ceramics.  hydroxyapatite  does  not  Unlike the break  down  other in  bio-resorbable  physiological  calcium  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 o f bio-active ceramics have been found to be unsuitable for loadbearing 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  2.3.1  2 6  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 o f implants and also coating for i m p l a n t s " ' ' ' . 7  Depending on the Ca/P ratio, water content, impurities, and temperature,  9  17  26  63  calcium  phosphate can be crystallized into different forms . There are several calcium phosphate 6  ceramics that are considered biocompatible. O f these, most are bio-resorbable and w i l l dissolve when exposed to physiological environments. Some o f these materials include: 22  amorphous calcium phosphate ACP, dicalcium octacalcium  phosphate  OCP  phosphate D C P ( C a H P C ^ H i O ) ,  (Ca8H2(P04)65H20),  tetracalcium  phosphate  TTCP  (Ca4P209), alpha-tricalcium phosphate a-TCP (Ca3(P04)2), beta-tricalcium phosphate (3TCP (Ca (P0 ) ), and hydroxyapatite H A (Cai (PO )6(OH) ). These forms of CaP are 3  4  2  0  4  2  summarized in Table 2.3-1 . 64  Table 2.3-1. Various calcium phosphate bioceramics Chemical Name  Abbreviation  Chemical Formula  Ca/P Ratio  Amorphous calcium phosphate  ACP  -  -  Dicalcium phosphate  DCP  CaHP0  Dicalcium phosphate dihydrate  DCPD  CaHP0 2H 0  1.00  Octacalcium phosphate  OCP  Ca H (P0 ) 5H 0  1.33  Alpha-tricalcium phosphate  a-TCP  a-Ca (P0 )  1.50  Beta-tricalcium phosphate  P-TCP  B-Ca (P0 )  Hydroxyapatite  HA  Ca (PO ) (OH)  Tetracalcium phosphate  TTCP  Ca P 0  1.00  4  4  8  2  2  4  6  2  3  4  3  10  4  4  4  6  2  2  1.50  2  2  1.67 2.00  9  2.3.1.1 Bioresorption and Biodegradation The dissolution products o f CaP bioceramics can be readily assimilated by the human  body.  Physiochemical  Bioresorption dissolution-  the biomaterial, Physical  and  biodegradation  is  generally  controlled  by  degradation depending on the local p H and the solubility o f disintegration-  particles, and Biological factors-  degradation due to disintegration into small  degradation cause by biological responses leading to  local p H 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 c r y s t a l l i n i t y ' . While 65  66  various variables w i l l have an effect on the biodegradation o f calcium phosphate, the general order o f 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 o f various forms o f CaP as a function o f p H is illustrated in Figure 2.3-1 . A s seen in Figure 2.3-1 hydroxyapatite is relatively insoluble compare to other 67  calcium phosphate phases, it is the only stable phase above p H 4.2. B e l o w p H 4.2, dicalcium phosphate dihydrate ( D C P D ) is the stable phase. It is often observed that unstable phases o f calcium phosphate w i l l dissolve and re-precipitate into stable phase at a given p H . The p H o f normal physiological environment is 7.2, however, this may decrease to as low as 5.5 in the region o f tissue injuries or inflammation, and slowly return to 7.2 over time.  24  0.0  30  40  50  ^ 60  70  pH  80  90  Figure 2.3-1. Logarithm o f the product o f calcium and phosphate concentrations plotted against p H values o f solution saturated with respect to various calcium phosphate phases in the ternary system Ca(OH)2-H P04-H 0. Calculated for 37°C. 3  2  © American Chemical Society, 1980, adapted by permission  2.3.1.2  1  M e c h a n i c a l Properties The biomedical applications o f calcium phosphates depend greatly on their  mechanical properties; however, the poor mechanical properties o f ceramic often limit its application. Table 2.3-2 summarizes the mechanical and physical properties o f 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  phosphates . 6  The wide range o f variations i n properties is due to the variation i n  structure, phase composition and manufacturing process.  Tensile and compressive  strength and fatigue resistance o f H A (like most ceramics) are highly dependent on the total volume o f porosity. For example, Jarcho et al reported compressive strength o f 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 l o w compressive strength value o f 138 M P a was reported by Rao et al, i n which the H A specimen was heat treated at 900°C for 0.5 hour , and likely produced high porosity fraction i n the material. 69  Furthermore, the presence o f porosity in form of either micropores (<lum) or macropores (>10um) can affect both the compressive and tensile strength o f the material . 10  Table 2.3-2. Mechanical and Physical Properties o f Calcium Phosphates Value  Property  Elastic Modulus (GPa)  4.0-117  Compressive Strength (MPa)  70-300  Bending Strength (MPa)  147  Hardness (Vickers, GPa)  1.1-6.0 0.27  Poisson's Ratio  1.90-3.27  B u l k Density (g/cm ) J  Theoretical Density (g/cm ) 3  Facture Strength (MPa)  3.16 12.8-60.4  26  2.3.2  H y d r o x y a p a t i t e Bioceramics Hydroxyapatite ( H A ) is chemically similar to the mineral component o f 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 o f H A is 10:6 and the theoretical density is 3.219 g/cm  . The mechanical properties o f  hydroxyapatite vary widely as many other calcium phosphates. Depending o f the final heat treatment conditions, Ca/P ratio, and the presence o f water and impurities, calcium phosphate can transform into H A or P-TCP. In many final products both H A and p - T C P phase coexist. A bioceramic consisted o f a mixture o f H A and P - T C P was first described 7ft  as biphasic calcium phosphate ( B C P ) by Nery et al in 1986 . It was further researched and reported that the bioactivity o f 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 s s u e ' ' . 7  8  64  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 o f powders, porous blocks, or beads.  Bone  substitute or bone filler was most often used i n the case o f bone cancer and bone augmentation, where a large section o f 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 o f nanocrystals o f 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 . 14  Endothelial cell is a layer that lines the cavities o f the heart and  the blood vessels. The reaction o f vascular endothelium cell to implanted biomaterial is of great importance because they interact in each step o f tissue integration. concluded that H A nanocrystals  exhibit high biocompatibility for  endothelium, and do not acquire a proinflammatory  It was  microvascular  or thrombogenic  phenotype.  Furthermore, endothelium was observed to be functioning toward angiogenesis - the formation o f new blood vessels.  2.3.2.1 Thermal Stability Hydroxyapatite is a hydrated calcium phosphate material.  The stability o f  calcium phosphate depends on the temperature o f preparation and the partial pressure o f water. H A was found to be the stable phase up to 1360°C at 500mmHg partial pressure of water. A n d , in the absence o f water content, p - T C P is found to be stable phase . A t 71  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 )6(OH) <=> 2 p- C a ( P 0 ) + C a ^ O o + H 0 4  2  3  4  2  2  C a i o ( P 0 ) ( O H ) <=> 3 p- C a ( P 0 ) + C a O + H 0 4  6  2  3  4  2  2  28  A t dry environment and at temperature above 1350°C, P - T C P w i l l form. Both H A and PT C P are very tissue compatible and o s t e o i n d u c t i v e  6,71  .  Therefore, heat treating H A  under vacuum condition can induce decomposition at lower temperature, promoting the formation o f 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  M e c h a n i c a l Properties Hasimah et al have conducted a study on the effect o f heat treatment temperature  on the mechanical properties o f hydroxyapatite , it was reported that upon heat treating 72  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 . 73  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  B i o m a t e r i a l Coatings Manufacturing o f medical devices often incorporates a surface coating process.  The design o f medical devices is generally based on the bulk properties o f 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 o f 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 o f a medical device for a specific application without detrimentally affect the bulk properties o f the biomaterial.  Table 2.4-1  summarizes the purposes o f  biomaterial coatings. The ultimate objective o f biomaterial coatings on medical devices is to combine the mechanical advantages o f metallic implant and the biocompatible properties o f coating to achieve the best desired biocompatible environment. Table 2.4-1. Purposes o f 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 o f the implant surface without delivery o f therapeutic  or other agents.  30  Whereas, active surface coatings are aimed to deliver bioactive compounds or drugs to influence the biological response i n order to achieve the desired function.  In coronary  stent applications, passive coating has often been applied as a thromboresistant coating; phosphorylcholine (PC) is one o f 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 o f complex coronary lesions . Drug eluting coatings are the 74  one o f 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 . 40  2.4.1  Hydroxyapatite Coatings Due to the highly bioactive and biocompatible nature o f 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 o f osseous applications, hydroxyapatite  coating can prevent fibrous tissue encapsulation on metallic implant, at the same time improves the integration o f bone into the implant, and thus provides a strong bond Q  between them .  In 1991, Furlong et al have performed a histological section o f H A  coated stem, which shown good osteointegration and formation o f new vital bone. There was no evidence o f an inflammatory reaction or o f fibrous tissue formation . 75  Further  31  study on H A coated porous metallic implant demonstrated a clear acceleration o f new bone ingrowth . 75  It is known that various metal ions can affect the functionality o f osteoblast cell , and H A coating was also observed to be able reduce the release o f metallic ions from the implant into the physiological surrounding. Sousa et al have reported that a film o f 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 o f metallic ion 77  .  This barrier effect o f H A was claimed to be a result o f metal phosphate formation or  incorporation o f metal ions i n the H A structure.  2.4.2  Processing of H y d r o x y a p a t i t e Coatings In the past, many different H A deposition methods have been reported, for  example: plasma spray " , sol-gel , pulse laser deposition , chemical vapor deposition, 15  17  18  19  sputtering , electrophoretic deposition , and electrochemical deposition ' . 20  21  22  24  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 o f coating material is irradiated by a high power source. Therefore these methods consume high amount o f 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 o f all, the high temperature o f deposition may lead to detrimental effects on the o f target substrate.  Plasma spray deposition o f H A has become the most popular and commercial method for coating H A on orthopedic devices, however, technical issues still remain. One o f the issues is the formation o f 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 o f plasma sprayed H A on titanium alloy and found that other than crystalline H A , amorphous calcium phosphate with Ca/P ratios o f 0.6 - 1 . 0 are present, also a calcium titanate phase was detected at the coating78  substrate interface  . Another issue is the poor adhesion o f H A by plasma spray. Radin  et al have found delamination o f H A coating from titanium substrate in an in-vitro stability study . 79  Table 2.4-2 summarizes the general advantages and disadvantages o f  the techniques used for the deposition o f hydroxyapatite.  33  Table 2.4-2. Summary o f techniques used for deposition o f hydroxyapatite coatings Technique Plasma Spray  Thickness Range 1 0 - lOOOum  Advantages • • • •  Sol Gel  Pulsed Laser Deposition  Sputtering  2.5  • • • • • • •  Line of sight process High temperature induced High cost May decrease fatigue life of substrate High capital cost Expensive raw materials May require control atmosphere  <l[im  • • •  High purity Low cost process Low heat treatment process temp. ( 2 0 0 - 6 0 0 ° C )  0.5-2mm  • • •  Cost effective Rapid deposition rate Uniform coating thickness on flat substrate  • •  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  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  0.02-lmm  • • •  Electro-deposition  Rapid deposition rate Dense under well control spray very high melting point materials Wide range of applications  Disadvantages  0.1-50um  • • • "  •  •  Line of sight process Cannot coat complex or porous substrate High capital cost  Electrochemical Deposition The electrochemical deposition ( E C D ) 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. reduction or an oxidation reaction.  The electrochemical synthesis can be a  B y manipulating the applied potential, the E C D  reaction can be continuously varied. There are a number o f features that differentiate E C D from other deposition technologies:  (1) The reaction o f electrochemical process  takes place within the electric double layer o f 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 o f a net charge on the solid surface (usually negative) causing a localized layer o f 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 o f conformal coating on complex shaped substrates. (3) Process temperature is often low (<70°C) since it is limited by the boiling point o f 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 o f 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 o f deposition instruments are relatively low compared to other technologies.  Film  thickness can be tailored over a wide range from ~0.1 um up to millimeter range while still maintaining its usefulness for a desired a p p l i c a t i o n ' . The thickness and chemical 23  26  composition o f the E C D coatings can be well controlled through adequate selection o f the process  parameters,  such  as  substrate composition and  electrolyte  composition,  electrolyte concentration, p H , and temperature, and mode o f applied potential (constant or periodic).  The two most important parameters in determining the course o f E C D  35  reactions are current density and applied potential, either one o f these parameters can be controlled as a function o f time to control the rate o f coating deposition.  2.5.1  E l e c t r o c h e m i c a l 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 i n the bulk solution  is relatively low (typically 4.0<pH<6.0).  The hydrolysis reaction and various cathodic  reactions w i l l consume water and produce OH", resulting in an increase o f 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 o f an aqueous  solution saturated with Ca(H P04)2 as an eletrolyte .  Shirkhanzadeh prepared an  81  2  electrolyte for bioactive calcium phosphate coating by dissolution o f hydroxyapatite into N a C l solution and further adjusted the p H by H Q  8 2  .  A solution saturated with  C a H P 0 4 - 2 H i O and added K N O 3 and N a H C 0 3 was prepared as electrolyte by Royer el 83  at .  Electrochemical deposition o f hydroxyapatite ( H A ) has been conducted i n the  mixed aqueous solution o f C a ( N 0 ) 4 H 0 and 3  2  2  NH H P0 4  2  2 3 4  ' ' . 8 4  8 5  According to a study by Shirkhanzadeh et al, the p H value within the diffusion layer o f the cathode during E C D w i l l be increased by the following two reactions: 2 H 0 + 2e- - » H + 2 0 H "  (1)  2 H + 2e"  (2)  2  +  2  H  2  36  The concentration o f phosphate HPO4 " is determined by the following reactions: H+ + P 0 " o H P 0 " 3  (3)  2  4  4  H P 0 " + OH"  H 0 + P0 "  2  (4)  3  4  2  4  With the extra O H " being produced, solubility limit o f H A is reached (as illustrated i n Figure 2.3-1) and consequently H A is deposited on the cathodic surface by the following reaction: 10Ca  2+  + 6P0 " + 20H"  Ca (PO ) (OH)  3  4  10  4  6  2  (5)  Although this has been accepted as the general reaction sequence leading to H A deposition, details o f cathodic reactions have been studied v i a cathodic polarization study to suggest further process controls.  In a cathodic reactions study done by Y e n et al,  hydroxyapatite was coated v i a electrochemical method on pure titanium . 23  Y e n 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 0 + 4e" -> 4 0 H " 2  H P 0 " + H 0 + 2e" - » H P 0 " + 2 0 H " 2  (2)  4  2  2  3  Applied Voltage at [-0.3 to -1. 1 V] 2 H + 2e" - » H +  H P 0 " + e" 2  (3)  4  2  HP0 " + A H 2  l  4  Applied Voltage at [-1.1 to-1.5 V] 2 H P 0 " + 2e" - » 2 P 0 " + H 2  3  4  (4)  2  4  2  Applied Voltage at [-1.5 to -3.0 V] 2 H 0 + 2e"-> H + 2 0 H " 2  2  37  It was found that at an applied voltage of-0.7 V , C a ( H P 0 H ) - H 0 and C a ( H P 0 ) H 0 3  2  2  2  4  2  2  were formed. A t -1.25V, C a H P 0 4 - 2 H 0 were found to be the major phase and H A were 2  found to be the minor phase.  A n d , at -1.55V, H A became the major phase o f 2  23  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 o f the reaction : 26  (1)  Reduction of oxygen at [-0.4 V] 0  (2)  2  + 2 H 0 + 4e" -¥ 4 0 H " 2  Reduction of H P0 ' 2  4  and 2HP0 ' 2  4  2 H P 0 - + 2e" -» 2 P 0 " + H 2  at [-0.4 to -1.6 V]  3  4  4  2  2 H P 0 " + 2e" -» 2 H P 0 " + H 2  2  (3)  4  4  2  Reduction of water at [-1.6 to -3.0 V] 2 H 0 + 2e"-* H + 2 0 H " 2  2  It can be seen that both Y e n et al and K u o et al have suggested very similar cathodic reactions.  Figure 2.5-1 shows the cathodic polarization curve o f titanium  substrate in the electrolyte compose o f C a ( N 0 ) 4 H 0 and N H H P 0 . It was found that 3  2  2  4  2  4  the greater amount o f hydroxyls produced from reduction o f water at voltage -1.6 to 3.0 V promotes the formation o f a purer H A . 2 6  38  "I'TITJ  ' '  1 , 1  10"  7  1  r""l"i"llTI|  '  • • • » tl ll  10  s  1  l-TTTTITJ  t  I I I Itill  11  1  r " " I I T IT  .  • 1 1 I 1 111  -  1x10"  5  1  r'TTI'TTTJ  1  I I I •Fill  1x10"  1  |-  •  • • • t i ••!  T  T  TTIT|  10"  4  10"  3  2  1  fT'TTTTFI  •  .  • i • • 1.1  10"  1  I/Area (A/cm ) 2  Figure 2.5-1. Cathodic polarization curve o f Ti substrate i n a Ca(NC>3)2 4H2O and N H 4 . H P 0 electrolyte. 2  4  © Journal of Materials Science and Engineering C, 2002, adapted by permission . 2  The  electrochemical deposition o f H A has also been conducted i n an electrolyte  composed o f calcium acetate, acetic acid, and sodium phosphate . 84  Manso et al have  subsequently dried the coated sample at 100°C and 900°C, X - r a y diffraction ( X R D )  o f 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 o f higher voltage a more compact film was produced, reducing the root mean square (rms) roughness from 90 nm to 20 n m . 84  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. 2  39  2.5.2  Microstructure and Phase of ECD Calcium Phosphate Coatings A s aforementioned,  current density is one o f the important parameters  in  controlling the rate o f electrochemical deposition and the phase deposition. Y e n et al have performed a series o f experiments at different current densities to investigate the 23  difference in phases and microstructures o f E C D - C a P  .  Electrolyte composed o f  Ca(N03)2-4H20 and N H 4 H 2 P O 4 was used for deposition o f CaP on pure titanium substrate, and current densities were selected as 0.3, 1.0, 2.0, and 3.0 m A / c m deposition time o f 1000s.  It was found that at 0.3 m A / c m , 2  2  for a  Ca(HP0 H) -H 0 3  2  2  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 , and a 2  fine loose plate-like structure was observed. A t 2 m A / c m , 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 o f hydrogen bubbles  . The "bubble-effects" were also  experienced by others electrochemical deposition studies when high current or voltage was applied to the system " . 86  88  A m o n g 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 o f electrochemical deposition o f calcium phosphate  in modified  simulated body fluid ( S B F ) was done by Peng et a l . S B F is an acellular fluid that has 89  40  inorganic ion concentrations similar to those o f human extracellular fluid.  Instead o f  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. observed from 8 to 20 m i n o f deposition.  The growth o f coating was  Sparsely distributed calcium phosphate  deposits (nuclei) were found at 8 min, and the coalescence o f these nuclei after 10 min o f deposition has increased the surface coverage.  A t 20 m i n o f deposition, coalescence  continues and an interconnected porous coating was observed. The pore size was in the range o f a few nanometers to 1 u m . X R D analysis has shown that the coating consists 89  of O C P and amorphous H A phases.  Similar to the work done by Y e n 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 o f 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 o f approximately 50 u m . After post-hydrothermal treatment at 180°C, the coatings were found to consist 90  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 o f H A . 9 0  41  K u o et al have conducted electrochemical deposition i n 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 o f 5 -  40 minutes at room  temperature. X R D analyses o f the coatings are shown in Figure 2.5-2 . W i t h a current 26  density o f 1 m A / c m , D C P D was observed to be the major phase and H A was the minor 2  phase. The same was observed with current density 5 m A / c m , but with an increased 2  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 o f deposition. The authors have suggested that only at current density above 10 m A / c m there w i l l be high enough 2  concentration o f O H " to convert HPO4 " into PO4 " to subsequently precipitate H A . 2  3  2 6  42  OTi  © HA # DCPD  26 (degree)  Figure 2.5-2. X P v D analysis o f 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  A d h e s i o n of H A Coatings Due to the brittle nature and adhesion concerns o f 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 . 91  Reprinted from Journal o f Materials Science and Engineering C , 2 0 , K u o M . C . , Y e n S . K . , The process o f electrochemical deposited hydroxyapatite coatings on biomedical titanium at r o o m temperature, 153 - 160, Copyright (2002), with permission from Elsevier. 3  43  The adhesion mechanism can be divided into three groups: (1) mechanical interlocking (2) physical bonding (3) chemical bonding. T o overcome the difficulty o f 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 o f the substrate), contamination on substrate, chemical bonding between coating and substrate, physical properties, surface roughness, and precleaning method allowing chemically modified surface . 92  L i n et al have conducted a study on the growth o f H A on 3 1 6 L stainless steel i n simulated body fluid, and alkaline treatment method was used to promote the adhesive strength o f H A coating . Alkaline treatment was performed by soaking the substrate in 91  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 o f  Na4Cr04 formed after the alkaline treatment, and the following reaction  was suggested : 91  8Na(OH) +  C r 0 => 2 N a 4 C r 0 + 3 H 0 + H 2  3  4  2  2  Tensile tests were performed to measure the tensile bonding strength o f 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 o f 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 o f OH" and H generation increases. When O H " generation rate exceeds the rate 2  of PO4 " formation, the excess OH" groups would migrate away from cathode because o f electric field and diffusion  . Therefore, the high p H boundary w i l l shift away from the  substrate and as a result calcium phosphate precipitation w i 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 o f H often led to a heterogeneous and loose coating structure. Attempts have 2  been made in the past to eliminate such phenomena. with  periodic  pulse  voltage  by  Shirkhanzadeh  electrodeposition process by Recently, H o u et al .  These included E C D processing et  al , 93  and  cathode  rotation  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  substrates . 87  45  3  3.1  SCOPE A N D O B J E C T I V E S  Scope of the Investigation The principal motivation o f 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 o f their excellent biocompatibility.  However, due to inferior  mechanical properties such as low fracture toughness (<1.12 MPaVm ) and low flexural strength (<140 M P a ) , the use o f hydroxyapatite has been limited to no-load or low-load bearing applications.  Hydroxyapatite coatings have a broad range o f applications in  medical devices and can provide superior biocompatibility in combination with the advantage of bulk material to achieve the best desired functions o f the medical device. Electrochemical deposition ( E C D ) o f uniform hydroxyapatite coating can be achieved on complex substrates.  E C D also bears other advantages, such as good control o f film  thickness in 0.1 - 10 urn range, low temperature o f processing (20 - 60°C), uniformity, and deposition rate and owing to low cost o f equipment and starting materials.  The  combination o f the potential advantages o f E C D for uniform thin film H A coating on complex substrates motivated us to investigate this process for coating coronary stents. There are strong indications that hydroxyapatite thin film coating on coronary artery stents can reduce or eliminate restenosis.  Hence, the process for electrochemical  deposition o f hydroxyapatite has been studied and optimized i n 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 o f electrochemical deposition, very few have focused on thin film (<0.5 (am) hydroxyapatite coating and none o f these have focused on coatings for stents.  The uniqueness o f present study is to combine the technique o f  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 o f in-vitro and in-vivo trials.  3.2  Objectives The broad objective o f this study is to achieve a well-controlled and reproducible  process to obtain thin film hydroxyapatite coatings v i a electrochemical deposition, on coronary stents made o f 316L stainless steel. The specific objectives o f the present thesis are as follows: 1. To develop reproducible hydroxyapatite coating process v i a 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 o f electrochemical deposition ( E C D ) process parameters on the evolution o f the resulting calcium phosphate coating. To develop an optimized electrochemical deposition process to deposit thin film hydroxyapatite coating on coronary stents by studying the process parameters,  such  as  current  density,  time  o f 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 o f substrate surface pre-treatment on the deposition process and resulting coating in terms o f physical (microstructure and phase) and mechanical (adhesion) properties. T o apply a substrate surface pre-treatment to achieve high level o f adhesion o f 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 o f the coatings. 6. To perform further characterization o f the coating in industrial setting: •  Qualitative mechanical assessment based on in-vitro crimping and expansion tests.  •  Rate o f dissolution based on in-vitro dissolution tests.  •  Overall performance qualification based on in-vitro stent deployment and 40 million cycles fatigue test.  48  EXPERIMENTAL METHODOLOGY The following experimental methods were employed: 1. Measurement o f the weight and thickness o f electrochemically deposited coatings for different process parameters, such as current density, time o f deposition, electrolyte concentration, p H , and temperature.  The process  stability, reproducibility and ease o f control were monitored. 2.  Evaluation o f microstructural morphology such as uniformity and porosity, and phase composition o f 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 posttreatment.  3.  Qualitative characterization o f 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 o f the above steps  1-3.  Application o f the optimum E C D process for deposition o f H A on coronary stents for further evaluation. 5. In-vitro evaluation o f E C D thin film coating on coronary stents, such as crimping and expansion tests, dissolution tests, and fatigue tests. 6. In-vivo evaluation o f 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 P r e p a r a t i o n Two types o f substrates made o f 316L medical grade stainless steel were used i n  this study, i.e. plate and a real coronary stent.  The nominal chemical composition o f  316L stainless steel is given in Table 4.1-1.  Table 4.1-1. Nominal chemical composition o f 316L stainless steel Element  C  Mn  Si  P  S  Cr  Mo  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 o f 2.5 cm, a height o f 3.0 cm and thickness o f 0.1 cm. The actual plate area being coated was a square with 2.5 c m width and thickness o f 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 u m mirror polish with L E C O ® diamond suspension. The polished specimens were degreased by detergent and further ultrasonically cleaned i n distilled water and absolute ethanol, then dried.  Stainless steel coronary stents were provided by M I V Therapeutics o f 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 m m outside diameter, 1.5 m m inside diameter and a length o f 14 mm. The M I V I stent is comprised of a series o f sinusoidal-ring geometries with two discs forming a single module, a flexible curlicue was fabricated to j o i n two modules together.  A 14 m m M I V I stent  consists o f four modules. A schematic diagram o f the stent is illustrated i n Figure 4.1-1. 50  The strut o f the stent had a width o f 100 um and the disc had a diameter o f 500 urn. Unless substrate surface treatment was performed, all stent specimens were used asreceived after cleaning. Stents specimens were cleaned ultrasonically i n absolute ethanol, and then dried in air.  Figure 4.1-1. Schematic diagram o f 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 o f 10N N a O H alkaline solution . 91  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 ) 4 H 0 and 0.04347M N H H P 0 . 3  2  2  4  2  4  The p H o f 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 M o d e l VI-700, I E T 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 o f the electrochemical deposition.  52  Power Supply  -+ Thermocouple  0  Amperometer  Voltmeter  Anode (Counter Electrode)  Cathode (Substrate)  Electrolyte Hotplate  Figure 4.2-1. Schematic diagram o f electrochemical deposition setup  4.2.1  Electrochemical Deposition Process Parameters Investigation In order to gain the knowledge o f 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 o f 1 - 15 m A / c m for 5 minutes deposition time.  •  Deposition  time.  E C D was conducted for duration o f 1 - 15 minutes with a  current density o f 1 m A / c m . •  Electrolyte concentration.  Concentration o f 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 o f 2.92, 2.63, 1.95 and 0.49, at current density o f 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 o f 1 m A / c m for 2 minutes. The specific range o f p H 2  was chosen for investigation to ensure a stable electrolyte was obtained without precipitation. •  Electrolyte temperature.  E C D was conducted at current density o f 1 m A / c m for 2  2 minutes deposition time with electrolyte temperature at 25, 45, and 70°C.  A l l o f the above electrochemical depositions were initially conducted on 316L stainless steel plates for the purpose o f 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 o f 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 ), with a Sartorius M E 2 3 5 P - S D balance with 0.01 mg accuracy. 2  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 o f the cross-section area.  Evaluation o f microstructural morphology such as uniformity and porosity, and phase characterization o f the resulting coatings was conducted for all o f the above specimens; the process stability and ease o f control were also closely monitored for  54  process optimization. For the purpose o f process consistency investigation, E C D coating was performed on five batches o f 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 o f 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 o f Vancouver, B . C .  The generic E C D process was performed  as discussed i n Chapter 4.2.1.  Optimized parameters were applied on all stent substrates with a 0.90 m A applied current, and a deposition time o f 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 o f 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 o f approximately 2 - 3 ensure the  years under in-vitro environment.  resulting electrochemically deposited  coating is o f stable  To  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. solution is approximately 12.5. distilled water and dried at 40°C.  The measured p H o f the 0.1N N a O H  The converted specimens were gently rinsed with 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 o f 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 B r u k e r - A X S , Germany) with C u Ka radiation (k= 1.5418 A ) and operated at a tube voltage o f 4 0 k V and a current o f 30mA. The range o f 29 was from 5° - 50° with a scanning rate o f 0.02°/s.  4.4  In-vitro Evaluations For the application o f 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 i n 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 a n d E x p a n s i o n 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 o f Vancouver, B . C . with M S I stent crimping machine (Machine Solutions Inc., M S I S C 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 o f 1.7 mm to 1.0 m m 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 o f the S D S 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 S D S was conducted with E n c o r e ™ inflation device (Encore™ 26 inflation device, Boston Scientific, Maple Grove, M N ) . The stent was expanded from a crimped diameter o f 1.0 m m to 3.0 m m with 14 atm pressure.  57  4.4.2  Dissolution Test Solubility o f various calcium phosphates differs greatly, which affect  the  biodegradation o f E C D coating. The variables that affect the dissolution not only consist of the phase and crystallinity o f 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 ) . K e y features o f the apparatus included precise bath temperature and rotation speed control, and the use o f 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 ( P B S ) 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 o f 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 o f fatigue testing is to meet the requirements as per the " F D A Draft  Guidance for the Submission o f Research and Marketing Applications for Interventional Cardiology Devices"  9 4  for in vitro mechanical fatigue testing.  The test should  demonstrate the safety o f 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 o f 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 o f the vessel in which it is implanted. The test is accelerated i n 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 o f in-vivo  implantation, i.e.  approximately 40 million cycles o f 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 o f 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 million fatigue cycles.  In fatigue testing, the objective was to determine i f the E C D - H A coating is able to survive 40 million cycles in a vessel that dilated with an average percent outer-diameter (%OD) strain o f 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 o f the simulated vessel, information  59  .J  was subsequently feedback to the control computer to calculate the percent radial stain o f 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 million cycles.  A t the end o f 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 o f 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 o f the coating surface or an area o f interest such as filtered debris. Both S E M and E D X analyses were used to demonstrate the existence o f remains o f coating after the fatigue test. Filter specimen was carbon coated to create a conductive surface to avoid charging in S E M .  60  5  R E S U L T S A N D 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 o f reaction in the E C D process . 25  In order to achieve  consistent results, current density instead o f applied current, was chosen for investigation because different substrates (stents and plates) were interchangeably used throughout the study. The surface areas o f different substrates were determined with the assumption o f negligible surface roughness.  The S E M observation o f plate specimens deposited at current density 15, 10, 5, 3, and 1 m A / c m 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 o f  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  rate '  86  .  The hydrogen bubbles rise to the surface and disturbed the formation o f  uniform coating.  In general, the microstructure o f coating deposited at 15 m A / c m exhibits a fine 2  and loose structure. Figure 5.1-1 (b) illustrates the microstructure o f coating deposited at  61  10 m A / c m 2.  "Bubble effect" was not observed at such current density, the coating  exhibits a fine and loose structure similar to 15 m A / c m , but with flakes growing on top. 2  At 5 m A / c m , 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 o f coating (Figure 5.1-1 (d)).  A t the lowest  current density o f 1 m A / c m , a thin layer (~1 um) o f porous coating was observed as 2  shown in Figure 5.1-1 (e). The morphology o f this thin layer was similar to the underlayer in Figure 5.1-1 (d), but more porous. Other than some small amount o f 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 o f microstructure evolution can be clearly noticed from Figure 5.1-1, as current density changes. mA/cm  facilitate the  Figure 5.1-1 (b) - (d) indicated that the current o f 10 - 3 growth o f 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 o f plate-like crystals, 1 m A / c m is more desirable 2  for thin film uniform coating.  Figure 5.1-2 illustrates the weight gain o f E C D coated  specimens deposited with current density 1 - 1 5 m A / c m for 5 minutes deposition. The 2  weight gain o f coatings increases with increased current density. It was observed that the weight gain slightly reduced with 15 m A / c m , which can be explained by the "bubble 2  effect" where fragments o f the coating were flaked off by the large volume o f hydrogen bubbles generated. This observation agrees with previous studies " 86  88  62  Figure 5.1-1. S E M images o f E C D coating deposited at various current densities: (a) 15 m A / c m , (b) 10 m A / c m , (c) 5 m A / c m , (d) 3 m A / c m , (e) 1 m A / c m [Left: xlOO; Right: x 1,500] 2  2  2  2  2  63  Figure 5.1-2. Weight gain o f E C D coated specimens versus current density with 5 minutes o f deposition.  5.1.2  Deposition T i m e Scanning electron microscopy images were obtained for E C D coatings deposited  at various times to evaluation the growth morphology and structural characteristics o f the evolving film on plate specimens.  Current density of 1 m A / c m was selected for this 2  study based on the resulting uniform thin film. The S E M observation o f plate specimens deposited for 15, 10, 5, 3, and 1 m i n are shown in Figure 5.1-3 (a) - (e), respectively (other process parameters are listed in Table 5.2-1).  From the observation o f the  deposition process, an "initiation period" (-15 seconds) for formation o f coating was noticed. This phenomenon was also reported in other previous s t u d i e s ' . A n initiation 89  95  period corresponds to the time from the application of current to the beginning o f calcium phosphate deposition.  It is believed that the initiation period is related to the initial  kinetic of reduction o f water, hydrogen gas generation and increase o f p H around the on  electrode, and consequently precipitation (deposition) o f calcium phosphate . 64  Figure 5.1-3. S E M images o f 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 l 5 0 0 ] at 1 m A / c m . 2  65  Figure 5.1-3 (a) shows the E C D microstructure after 15 m i n o f deposition; platelike 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 o f 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 o f 1 m i n deposit is shown i n Figure 5.1-3 (e). Under x l , 5 0 0  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 o f the 1 m i n 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 o f E C D specimens coated for 1 - 15 minutes with 1 m A / c m is shown in Figure 5.1-5. W i t h increase o f deposition time, an increased in weight gain was observed. A more drastic increase o f weight gain was seen from 3 minutes to 5 minutes of deposition; this is believed to be a result o f the growing crystals on top o f the thin film 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 and with 1 minute 2  deposition time.  x£0k  0000  20  KV  gjJM  Figure 5.1-4. H i g h magnification S E M image o f E C D coating with deposition o f 1 min [x20,000] at I m A / c m . 2  0.6 •  0.5  •  —  E 0.4 -o D)  •  ight gain  e  0.3 0.2  2  0.1  0  2  4  6  8  10  12  14  16  Deposition Time (min)  Figure 5.1-5. Weight gain o f 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  be dependent on the composition o f electrolyte '  R7  R Q  R1  '  '  '  . However, it was observed  that for electrolyte composed o f C a ( N 0 ) 4 H 0 and N H H P 0 , the Ca/P ratio o f the 3  2  2  4  2  4  resulting deposits were nearly independent o f the Ca/P ratio o f the electrolyte. Figure 5.1-6  illustrates the different Ca/P ratio o f 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 o f C a ( N 0 ) 4 H 0 and N H H P 0 3  2  2  4  2  4  for the preparation o f different Ca/P electrolytes.  The  Ca/P ratio of the deposits ranged from 1.41 - 1.57, independent o f the Ca/P ratio o f the electrolyte. S E M images o f the various resulting deposits deposited with electrolyte o f 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 o f electrolyte has minimal effect on the Ca/P ratio o f the resulting deposit and the resulting microstructure.  In the present study, the surface o f 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 o f E D X quantitative  analysis with the use o f well-polished standards having a composition similar to the sample has been reported to be greater than 2 % relative for major concentrations . The 105  analysis o f elements with concentrations less than 5wt% w i l l typically yield relative  68  accuracies o f - 1 0 % , 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 o f C a ( N 0 ) 4 H 0 and N H H P 0 for the 3  2  2  4  2  4  preparation o f different Ca/P ratio electrolytes.  Electrolyte Concentration (M) Ca(N0 ) 4H 0  NH4H2PO4  3 2  0.0217 0.0217 0.0217 0.0435  Electrolyte Ca/P ratio  2  0.0635 0.0572 0.0423 0.0212  2.92 2.63 1.95 0.49  1.8 W  o 1.6 a. 0  Q  1.4  Ui c 1.2  3  (fl 1.0  O CL 0.8  O O  0.6  cc  0.4  a.  0.2  CC  re O 0.0 2.92  2.63  1.95  0.49  Ca/P Ratio of Electrolyte  Figure 5.1-6. Ca/P ratio o f resulting deposit with the use o f various Ca/P ratio electrolytes. Ca/P ratio was derived from E D X spectra.  69  Figure 5.1-7. S E M images o f 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 o f 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 p H = 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 o f 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 o f temperature on the supplied voltage was noticed.  Figure 5.1-9 shows typical influence o f the electrolyte temperature on the  measured supply voltage. The voltage decreases as the electrolyte temperature increases, possibly due to decrease o f water resistivity, as previously suggested .  71  Figure 5.1-9. Influence o f electrolyte temperature on measured supply voltage, for constant current source (I = 13.77 m A ) .  5.1.5  T h e 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  (Figure 2.3-1). However, the  [ 6 7 ]  phase composition, crystallinity, nature o f solution, temperature, and porosity also plays a role in dissolution and precipitation o f calcium phosphates . 6  In the process o f 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 coverage as shown i n Figure 5.1-10.  non-uniform surface  S E M images o f E C D coating deposited with  electrolyte o f 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 o f p H 4.0 - 5.5 was concluded to be optimal for the process under consideration.  Figure 5.1-10. S E M Image o f E C D conducted with p H 3.0 electrolyte at 45°C. [x5,000]  Figure 5.1-11. S E M images o f 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 5.2.1  ECD of Calcium Phosphate Coatings on Coronary Stents 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 o f the E C D coatings. In order to achieve thin film (<0.5 um) coating and to maintain a desire coating microstructure (i.e. porous structure" without large plates o f 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 i n Table 5.2-1). It was observed that the application o f different current densities had a more significant effect on the coating microstructure compared to deposition time, whereas the deposition time had more control o f weight (or thickness) o f the resulting coatings. For the purpose o f depositing thin porous coating as illustrated in Figure 5.1-1 (e) on coronary stent, current density o f 1 m A / c m was chosen for deposition with various deposition times. The Ca/P ratio o f 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 o f 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 p H was found to facilitate uniform deposition.  74  Table 5.2-1. Optimum parameters for E C D o f calcium phosphate coatings*. Value  E C D Process Parameters Electrolyte Ca/P ratio Calcium Nitrate [ C a ( N 0 ) 4 H 0 ] 3  2  2  A m m o n i u m phosphate [ N H H P 0 ] 4  2  4  Current Density Deposition Time Electrolyte p H Electrolyte Temperature V  1.95 0.04347M 0.02329 M 1.0 m A / c m 1 minute 4.5 45°C  2  ~  Electrolyte solution was prepared with distilled water, typical volume of electrolyte = 400mL.  Figure 5.2-1  shows S E M images o f 316L stainless steel bare metal stent  manufactured by M I V Therapeutics, Vancouver, B C . M i n o r bumps can be observed on the edge areas as seen i n 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 o f 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 o f optimum parameters, for 1 minute deposition (coating thickness was estimated based on crosssection evaluation o f 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 o f 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 o f a calcium phosphate coating. Increasing the deposition time to 2 minutes clearly increased the density o f 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 o f 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 o f thickness and density of the coating as deposition time increased.  Figure 5.2-1. S E M images o f bare metal stent with various magnification: a) [xlOO], b) [x800], c) [xl,500].  76  Figure 5.2-2. S E M images o f 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 o f 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 o f 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 o f 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  Fe  800  600  -  Ni  400  Ca  Fe  200  -  Cr Si Mn  Fe  Mn T  1  |  4  I  1  1  1  1  1  Ni 1  6  Ni [—1  8  1  1  keV  Figure 5.2-6. E D X surface analysis o f the E C D coating deposited on coronary stent with optimum parameters for 1 minute.  Figure 5.2-7 shows the X - R a y diffraction pattern o f the E C D coating deposited with optimum process parameters (Table 5.2-1), for 1 minute.  Due to the technical  difficulty of X - R a y 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. T w o 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 2theta 25.98° and 31.94°. Stainless steel substrate has a peak at 2-theta 44.52°. However, due to the similarities o f 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 o f highly soluble  DCPD  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% o f its weight (Diotai)  (Di/ ) 2  within 20 minutes, and was totally dissolved  in approximately 40 minutes, in phosphate buffer saline.  •  c  cD A  5  10  15  20  ,•  O  1  25  30  \ 35  40  45  50  2-Theta (deg) O  Dicalcium phosphate dehydrate  (DCPD)  A  Hydroxyapatite (HA) I Octacalcium phosphate  •  316L Stainless Steel  (OCP)  Figure 5.2-7. X - R a y diffraction o f the E C D coating deposited with optimum parameters (Table 5.2-1), for 1 minute, showing a mixed phase o f 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 o f the  80  coating cross-section. It was revealed that the coating thickness o f 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 C a P 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 o f " V " section o f the stent . 27  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 o f coating shown in Figure 5.2-10.  The deformation o f 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 o f large coating section.  The tensile stress in brittle films can lead to through-thickness  cracking as illustrated schematically in Figure 5.2-12 . Figure 5.2-10 (d) shows tensile 99  stress causing a series o f parallel cracks with approximately uniform spacing (~5 um) indicating a poor interfacial bonding. The delamination allowed better estimation o f the coating thickness, observed again to be -0.5 um, which agrees with the cross-section evaluation illustrated i n 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 o f 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  CTc  —  coatina substrate area of decohesion  (b) 0C  ^  _  ->*—  ^  Cc  Figure 5.2-11. Compressive spallation by buckling showing localized interfacial decohesion. © Journal of Engineering Failure Analysis 2, 1995, adapted by permission . 4  Reprinted from Journal o f Engineering Failure Analysis 2 , 2, Strawbridge A , E v a n s H . E . , Mechanical failure of thin brittle coatings, 85 - 103, Copyright (1995), with permission from Elsevier.  4  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 permission . 5  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 substrate . 91  Figure 5.2-13 shows the S E M image o f 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 o f the surface pattern on the order o f 100 nm. Figure 5.2-14 illustrates the E D X analysis o f the pretreated 316L stainless steel surface. While Fe, N i , and Cr are the main constituents o f 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 . The nano91  rough surface structure formed during the surface modification process is believed to promote mechanical interlocking and thus physical bonding.  Additionally, the inter-  compound layer o f 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 o f 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 x 2 0 k  2um  Figure 5.2-13. S E M image o f 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 intercompound can be seen with cracks. 4  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 - 5 0 % 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 o f 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 o f a 3.0 m m diameter catheter) with an E C D coating deposited using the optimum deposition parameter (Table 5.2-1), on surface modified stent. buckling or delamination.  There is no coating  However, both Figure 5.2-17 (c) & (d) revealed a series o f  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 o f a more severe expansion test results (known as "over-expansion" in industry, with the use of a 3.5 m m 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 o f deformation in Figure 5.2-18 (a) was notably higher compared to Figure 5.2-17 (a), although the degree o f strain can not be quantified. Even though there were obvious nano-cracks (<100nm) present in the coating, yet no delamination or separation o f the coating was seen in Figure 5.2-18 (d).  Fracture o f 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, i n the areas o f 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 o f 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 o f 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 o f the Na4Cr04 inter-layer that have changed the surface chemistry, promoting formation o f 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 C o m p o s i t i o n of E C D C a l c i u m Phosphates Coatings To prolong resorption o f the E C D as-deposited mixed-phase calcium phosphate  coatings, phase and crystallinity optimization was performed. A s mentioned i n Chapter 5.2.1, total dissolution o f 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 ( D C P D ) . Here after, the E C D deposited coating with the use o f the optimum deposition parameters w i 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 . 1 M o f N a O H aqueous solution at 75°C for 12, 24, 48, and 72 hours ' ° . 81  10  Figure 5.2-19 illustrates the X - R a y diffraction patterns o f as-deposited E C D coating and after various time o f the above phase conversion process. U p o n the phase conversion process, D C P D peaks were no longer found i n 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 o f 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 : 96  5CaHP0 -2H 0 + 60H" o 4  2  Ca (P0 ) OH + 2P0 " + 15H 0 3  5  4  3  4  2  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/ ) o f its 2  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 o f as-deposited E C D coating and the resulting coating after 12, 24, 48, and 72 hours of NaOH( ) phase conversion at 75°C. aq  The microstructure o f 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 i n the microstructure  of E C D coating after 72 hours o f the conversion process as illustrated i n 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 o f 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]  I  L  • • ' f„>  J  •:•  • .;: f  *  J  S  Figure 5.2-21. S E M images o f 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 . 65  It is common to measure crystallinity by using  diffraction techniques, such as x-ray diffraction: greater crystallinity yields a sharper 102  diffracted beam  103  '  . 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 . 30  High heat treatment temperatures induce H A crystals growth, and thus lead to the decrease o f total porosity and pore s i z e ' 7 3  1 0 4  . Depending on the calcium to phosphorous  ratio, presence o f water and impurities, amorphous H A can be transform to crystalline HA ' 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 . 1 M 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 - R a y diffraction patterns o f the E C D coating after 12 hours of 0.1N N a O H ( ) phase conversion and after the subsequent 300°C, 500°C, and aq  750°C heat treatment o f 20 minutes. A s the heat treatment temperature increases, the peaks become narrower and individual peaks o f 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 o f poorly crystallized H A . A t 500°C, the signature peaks (as indicated by J C P D C #009-0432) o f 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 o f the peaks was considerably narrower as compared to the 300°C  94  samples. Upon heat treating the coating at 700°C, a pure phase o f 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 o f 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 o f 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 o f 12 hours NaOH( ) phase converted E C D coating and after 300°C, 500°C, and 750°C for 20 minutes o f heat treatment o f the coating. aq  Although the width o f 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 o f crystallinity or grain size. Numerous studies suggested that the narrowing o f H A peaks is 95  indication o f higher crystallinity '  '  '  .  Figure 5.2-23 (b) & (c) illustrates the  microstructure o f the 500°C and 750°C heat treated coating on stent, respectively. Although the overall microstructure o f 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 o f the coating appeared to have increased. The loosely attached structure found i n 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( ) treatment. Furthermore, the aq  phase conversion process (with NaOH( ) treatment + 500°C heat treatment) prolonged aq  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 N a O H ( ) aq  treatment and 500°C / 20 minutes heat treatment, after stent expansion. H i g h degree o f substrate deformation is noticed, since a 3.5mm diameter catheter was used for the expansion test.  Examinations o f both tensile and compressive areas have found no  detrimental cracking or delamination o f the coating.  Although some fine cracks  (~300nm) were noticeable, it is reasonable to believe that the heat treatment process has increased the integrity o f the coating.  96  Figure 5.2-23. S E M images o f E C D coating after phase conversion process [xlO,000]. (a) 12 hours N a O H treatment (b) 12 hours N a O H treatment + 500°C heat treatment (c) 12 hours N a O H treatment + 750°C heat treatment. ( a q )  ( a q )  ( a q )  Table 5.2-2. Summary o f E C D coatings dissolution test data. Dissolution tests were conducted with 10 m L o f phosphate buffer saline (PBS) at 37°C ( p H = 7.4) with rotation speed at 20 rpm. **  55  E C D C o a t i n g Description A s deposited E C D Coating E C D Coating with N a O H ) Conversion E C D Coating with N a O H ) Conversion + 500°C heat treatment  T i m e to D1/2 20 minutes 4 hours > 4 weeks  ( a q  ( a q  **  T i m e to D tai 40 minutes To  6 Vi hours > 4 weeks  * All Coatings were prepared as described in Table 5.2-1. ** D denotes the time of 50% coating weight lost, D ai denotes the time of total coating dissolution. 1 / 2  Tot  97  Figure 5.2-24. S E M images o f expansion test result from an as deposited E C D coating upon NaOH( ) 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] aq  5.2.5  In-vitro Fatigue Test The purpose o f this test was to ensure E C D - H A coated stents are fatigue-resistant  in an accelerated 40 million cycles test, (i.e. simulating about 1 year o f 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 o f 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 o f the coating  98  surface or an area o f 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 o f the simulated vessel, another into the distal end. The target % O D strain o f 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 . 5 6 5 % and vessel #3 shows the least strain at -0.395%.  Specimen Position Vessel Vessel Vessel Vessel Vessel Vessel  #1 #1 #2 #2 #3 #3  Proximal Distal Proximal Distal Proximal Distal  A^eiage '• OD Strain 0.51 0.52 0.57 0.56 0.35 0.44  Spcimcn ID 1P 1D 2P 2D 3P 3D  0.6 0.5  1P  1D  2P  2D  3P  3D  Specimen ID  Figure 5.2-25. Summary o f average % O D strain for the six fatigue tested stent specimens.  99  S E M images o f specimen IP after the fatigue testing are shown in Figure 5.2-26. Crystals o f 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 P B S did not contain any calcium, the content o f calcium i n E C D - H A was used as an indicator o f the coating existence. Figure 5.2-26 (b) illustrates the area analysis by E D X . The results have shown a high count o f calcium suggesting the E C D - H A coating still remains on the stent surface (Figure 5.2-26(d)). The side edge o f stent I P 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" o f the coating was likely due to the contact between the vibrating vessel w a l 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 o f 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 o f E C D - H A coating.  The  characteristic porous microstructure o f 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 o f 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 o f vessel#3 was the least among the three vessels, the S E M image o f 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. A n d , similar to specimen 2P, a denser porous microstructure was observed (Figure 5.2-28 (c)). Unlike specimen I P and 2P, the coating microstructure o f 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 million cycles fatigue test the P B S 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 i n the fatigue test system. In general, six categories o f debris can be recognized. Figure 5.2-29 illustrates the six main categories o f debris found, namely D l - D 6 . Debris size was observed to range from 5um - 30um diameter.  E D X analyses o f 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 D 2 , nevertheless, calcium content was minimal (0.37wt%). Judged from the high phosphorus content, it is believed that D 2 was a P B S chlorine crystal buildup on coating surface, possibly detached during the test along with a trace o f E C D - H A coating. Most important o f all, none o f the debris found on filter resemble the shape or microstructure o f the E C D coating. Debris D 2 , D 4 , and D6 all contain high content o f chlorine from P B S 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 o f 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 o f 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 o f the debris may also contains elements o f the substrate and the surrounding area. Therefore, as seen in Figure 5.2-30, a high content o f carbon and oxygen from the filter material is clearly distinguishable. A l s o , it was reported that for rough surface samples or small particles, the relative accuracy o f 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" o f 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 o f 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-D5  EDX-D2 150C  A  Figure 5.2-30. E D X analysis o f the six main categories o f debris found after the fatigue test.  107  5.2.6  R e p r o d u c i b i l i t y a n d 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( ) + 500°C heat treatment) on the aq  surface modified stent, were run repeatedly to evaluate process reproducibility. Five batches o f E C D coatings were produced and characterized. Each batch consisted o f 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 o f the average weight o f E C D coating per each batch and their yield rate. Y i e l d rate is defined as the fraction o f coating processed within the specification limits.  Details of records can be found i n Appendix B .  The  qualification criterion for coating weight (specification) was set to be 62 pg ± 5 pg. A l l S E M observations o f coating surface have shown uniform surface coverage and porous microstructure as illustrated in Figure 5.2-20. The average weight o f 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, E C D - R P - 0 0 3 - 0 5 , and E C D RP-004-02) was found to have weighted outside o f the acceptable range and was rejected. It was concluded that, in term o f coating microstructure, uniformity and weight, the application o f E C D - H A coatings on coronary stent exhibits high process reproducibility and consistency.  108  Table 5.2-3. Summary o f five batches o f E C D coating average weight and yield rate. Coating Weig ht of E C D HA B a t c h (ug) Stent Number 01 02 03 04 05 06 07 08 09 10 Average (u,g):  ECD-RP-001  ECD-RP-002  ECD-RP-003  ECD-RP-004  ECD-RP-005  62 63 66 65 64 60 61 62 66 65 63.4  60 62 63 61 60 62 61 42 65 64 62 0  61 59 61 64 73 61 60 62 63 62 61.4  64 23 63 67 65 62 61 63 61 66 63.6  65 64 64 63 65 62 59 61 65 59 62 7  20  1.6  1.4  2.0  2.2  100%  90%  90%  90%  100%  \ Overall Average (pg): Standard Deviation (u.g): Overall Standard Deviation (u,g): Yield Rate: Overall Yield Rate:  >  v  '  94% ''  :  * Rejected samples w ere excluded from average and standard deviation calculation  K  [  j  5.2.6.1 Errors in ECD-HA Characteristics and Process Parameters Measurement Within the group o f 50 stents used for reproducibility study, stents weights were in the range o f 15.243 mg - 16.897 mg. According to M I V Therapeutics, the weight variation was attributed to the stent manufacture electropolishing.  It was also suggested  process, i.e. laser cutting and  by M I V Therapeutics that under these  manufacture process variations, the surface area o f the stent may vary as much as 5%. In consideration o f the current density (1 m A / c m ) applied under the optimum coating 2  condition, the surface  area variation can impose a 5% error i n 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 o f electrolyte p H was done with Beckman 260 p H meter with an inhered error o f ± 0.004 p H . Chemical preparation for the E C D electrolyte was performed with Sartorius M E 2 3 5 P - S D 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 o f 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 ( E C D ) was used to deposit uniform calcium  phosphate coatings on 316L stainless steel coronary stents. The influence o f 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  o f coating being  deposited. In order to achieve uniform thin film (<0.5 urn) coating on the complex surface o f 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 o f 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 o f current density has a more significant effect on the resulting Ca/P microstructure, whereas the deposition time has more control o f weight or thickness o f the resulting coating.  The coating microstructure exhibits  plate-like crystals when high current density (> 3 m A / c m ) 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 o f coating microstructure and thickness (~0.5 urn), as demonstrated through a five batches of ten stents reproduction evaluation. The average weight o f 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 N a O H ( ) at 60°C for 24 hours, aq  and subsequently heat-treatment at 500°C for 20 minutes. In-vitro stent crimping and expansion tests found that the application o f 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 - R a y 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 N a O H ( ) phase conversion at 75°C for 12 hours and a 500°C heat aq  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 ) 4 H 0 ] 3  2  2  A m m o n i u m phosphate [ N H H P 0 ] 4  • • • •  9.  Current Density Deposition Time Electrolyte p H Electrolyte Temperature  2  1.95 0.02329 M 0.04347M  4  1.0 m A / c m 1 minute 4.5 45°C  2  The standard 40 million cycles fatigue test validated the safety and reliability o f the optimized E C D - H A coatings with the incorporation o f the substrate surface treatment and the phase conversion processes. S E M and E D X analyses o f the stent specimens retrieved from fatigue test have shown no sign o f 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 FUTURE 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 o f electrochemical deposition on various substrates.  The difference in  substrate material and geometry strongly influence the process o f deposition, and certainly poses multiple challenges.  Application o f E C D on other implantable  biomedical devices should be highlighted, such as C o - C r 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 o f 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 o f drug encapsulation or impregnation into E C D - H A should focus on maximizing the drug content in the coatings, and optimization o f the drug release profile.  114  Electrochemical co-deposition of organo-ceramic coatings.  The co-deposition  studies should emphasize feasibility o f co-depositing pharmaceutical agents for drug eluting purpose and co-depositing reinforcing agents, such as polymer, for coating mechanical properties improvement (Appendix A ) , such that thicker (several microns thick) coatings with higher capacity for drug encapsulation can be processed.  115  REFERENCES  1. American Heart Association, "Cardiovascular Disease Statistics,"  2006,  1 (2005).  2. F. Faccioni, P. Franceschetti, M . Cerpelloni and M . E . Fracasso. "In vivo study on metal release from fixed orthodontic appliances and D N A damage i n oral mucosa cells," American Journal o f Orthodontics and Dentofacial Orthopedics, 124, 687-93 (2003). 3. J.J. Popma and M . T u l l i . "Drug-eluting stents," Cardiol. C l i n . , 24, 217, (2006). 4. C M . Gibson, D . Karmpaliotis and L . Kosmidou, et al. "Comparison o f effects o f bare metal versus drug-eluting stent implantation on biomarker levels following percutaneous coronary intervention for non-ST-elevation acute coronary syndrome," A m . J. Cardiol., 97, 1473-7 (2006). 5. R. Wessely, A . Kastrati and A . Schomig. "Late restenosis i n patients receiving a polymer-coated sirolimus-eluting stent," A n n . Intern. M e d . , 143, 392-4 (2005). 6. Joon B . Park and Joseph D . Bronzino. "Biomaterials: Principles and Applications," 264 (August 29, 2002). 7. K . A . Hing, S . M . Best, K . E . Tanner, W . Bonfield and P A . Revell. "Biomechanical assessment of bone ingrowth i n porous hydroxyapatite," Journal o f Materials ScienceMaterials in Medicine, 8, 731-6 (1997). 8. G . Willmann, "Coating o f implants with hydroxyapatite material connections between bone and metal," Advanced Engineering Materials, 1, 95-105 (1999). 9. W . Suchanek and M . Yoshimura. "Processing and properties o f hydroxyapatite-based biomaterials for use as hard tissue replacement implants," J. Mater. Res., 13, 94-117 (1998). 10. L . L . Hench, "Bioceramics - from Concept to Clinic," American Ceramic Society Bulletin, 72, 93-8 (1993). 11. X . L u and Y . Leng. "Theoretical analysis o f calcium phosphate precipitation i n simulated body fluid," Biomaterials, 26, 1097-108 (2005). 12. W . Suchanek and M . Yoshimura. "Processing and properties o f hydroxyapatite-based biomaterials for use as hard tissue replacement implants," J. Mater. Res., 13, 94-117 (1998). 13. S.L. Shi, W . Pan, M . H . Fang and Z . Y . Fang. "Reinforcement o f hydroxyapatite bioceramic by addition o f T i 3 S i C 2 , " J A m Ceram Soc, 89, 743-5 (2006).  116  14. S. Pezzatini, R. Solito and L . Morbidelli, et al. "The effect o f hydroxyapatite nanocrystals on microvascular endothelial cell viability and functions," Journal o f Biomedical Materials Research Part a, 76A, 656-63 (2006). 15. K . Degroot, R. Geesink, C . P . A . T . K l e i n and P. Serekian. "Plasma Sprayed Coatings of Hydroxylapatite," J. Biomed. Mater. Res., 21, 1375-81 (1987). 16. J. Weng, Q. L i u , J . G . C . Wolke, X . D . Zhang and K . deGroot. "Formation and characteristics o f the apatite layer on plasma-sprayed hydroxyapatite coatings in simulated body fluid," Biomaterials, 18, 1027-35 (1997). 17. H . Liang, B . Shi, A . Fairchild and T. Cale. "Applications o f plasma coatings in artificial joints: an overview," Vacuum, 73, 317-26 (2004). 18. C M . Lopatin, V . Pizziconi, T . L . Alford and T. Laursen. "Hydroxyapatite powders and thin films prepared by a sol-gel technique," Thin Solid Films, 326, 227-32 (1998). 19. C K . Wang, J . H . C L i n , C P . Ju, H . C . Ong and R . P . H . Chang. "Structural characterization o f pulsed laser-deposited hydroxyapatite film on titanium substrate," Biomaterials, 18, 1331-8 (1997). 20. T. Seno, Y . Izumisawa and I. Nishimura, et al. "The interfacial strength in sputteringhydroxyapatite-coating implants with Arc-deposited surface," Journal o f Veterinary Medical Science, 65, 419-22 (2003). 21.1. Zhitomirsky, "Electrophoretic hydroxyapatite coatings and fibers," Mater Lett, 42, 262-71 (2000). 22. H . B . H u , C . J . L i n , P . P . Y . L u i and Y . Leng. "Electrochemical deposition o f hydroxyapatite with vinyl acetate on titanium implants," Journal o f Biomedical Materials Research Part a, 6 5 A , 24-9 (2003). 23. S.K. Y e n and C M . L i n . "Cathodic reactions o f electrolytic hydroxyapatite coating on pure titanium," Mater. Chem. Phys., 77, 70-6 (2003). 24.1. Zhitomirsky, "New developments in electrolytic deposition o f ceramic films," American Ceramic Society Bulletin, 79, 57-63 (2000). 25.1. Zhitomirsky, "Cathodic electrodeposition o f ceramic and organoceramic materials. Fundamental aspects," A d v . C o l l o i d Interface Sci., 97, 279-317 (2002). 26. M . C . K u o and S . K . Y e n . "The process o f electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature," Materials Science & Engineering C-Biomimetic and Supramolecular Systems, 20, 153-60 (2002). 27. M I V I Technologies Inc., "Structural Analysis o f M I V I Technologies Coronary Stent Model C8-6: Finite Element Analysis Report," 1, 17-80 (2000).  117  28. K . A . Gross, W . Walsh and E . Swarts. "Analysis o f retrieved hydroxyapatite-coated hip prostheses," J. Therm. Spray Technol., 13, 190-9 (2004). 29. H . Gierse and K . Donath. "Reactions and complications after the implantation o f Endobon including morphological examination o f explants," Arch. Orthop. Trauma Surg., 119, 349-55 (1999). 30. L . M . Sun, C C . Berndt, K . A . Khor, H . N . Cheang and K . A . Gross. "Surface characteristics and dissolution behavior of plasma-sprayed hydroxyapatite coating," J. Biomed. Mater. Res., 62, 228-36 (2002). 31. American Heart Association, "Heart Attack," 2006, 1 (2006). 32. Heart & Stoke Foundation, "Heart Disease," 2006, 1 (2006). 33. A . H . Gershlick, "Coronary disease - Role o f stenting i n coronary revascularisation," Heart, 86, 104-12 (2001). 34. M . M . Gandhi and K . D . Dawkins. "Intracoronary stents," B r . M e d . J., 318, 650-3 (1999). 35. P . W . Serruys and D . Keane. "Randomized Trials o f Coronary Stenting Introduction," J. Interv. Cardiol., 7, 331- (1994). 36. C L . Zollikofer, F . Antonucci, G . Stuckmann, P. Mattias and E . K . Salomonowitz. "Historical Overview on the Development and Characteristics o f Stents and Future Outlooks," Cardiovasc. Intervent. Radiol., 15, 272-8 (1992). 37. Kutyk M J . B . and Serruys P W . "Historical Overview, "pp. 1-16 i n CORONARY STENT Current Perspectives. Edited by Kutyk M J . B . and Serruys P W . Martin Dunitz, United Kingdom, 1999. 38. P . W . Serruys, P. Dejaegere and F. Kiemeneij, et al. " A Comparison o f BalloonExpandable-Stent Implantation with Balloon Angioplasty i n Patients with CoronaryArtery Disease," N . Engl. J. M e d . , 331,489-95 (1994). 39. D . L . Fishman, M . B . Leon and D.S. Bairn, et al. " A Randomized Comparison o f Coronary-Stent Placement and Balloon Angioplasty in the Treatment o f Coronary-Artery Disease," N . Engl. J. M e d . , 331,496-501 (1994). 40. Ioannis K . Stefanidis, Vasilios A . Tolis, Dimitrios G . Sionis and Lambros K . Michals. "Development i n Intracoronary Stents," H J C , 43, 63-7 (2002). 41. V . Rajagopal and S.G. Rockson. "Coronary restenosis: A review o f mechanisms and management," A m . J. M e d . , 115, 547-53 (2003).  118  42. F. A i r o l d i , A . Colombo and D . Tavano, et al. "Comparison o f diamond-like carboncoated stents versus uncoated stainless steel stents in coronary artery disease," A m . J. Cardiol., 93, 474-7 (2004). 43. K . Gutensohn, C . Beythien and J. Bau, et al. "In vitro analyses o f diamond-like carbon coated stents: Reduction o f metal ion release, platelet activation, and thrombogenicity," Thromb. Res., 99, 577-85 (2000). 44. B . Heublein, C . Ozbek and K . Pethig. "Silicon carbide-coated stents: Clinical experience in coronary lesions with increased thrombotic risk," Journal o f Endovascular Surgery, 5,32-6 (1998). 45. M . Unverdorben, M . Schywalsky and D . Labahn, et al. "Stents coated with hypothrombogenic amorphous silicon carbide - preliminary results i n the N e w Zealand White Rabbit," Perfusion, 13, 124,+ (2000). 46. N . Huang, P. Y a n g and X . Cheng, et al. "Blood compatibility o f amorphous titanium oxide films synthesized by ion beam enhanced deposition," Biomaterials, 19, 771-6 (1998). 47. A . Kastrati, A . Schomig and J. Dirschinger, et al. "Increased risk o f restenosis after placement o f gold-coated stents - Results o f a randomized trial comparing gold-coated with uncoated steel stents in patients with coronary artery disease," Circulation, 101, 2478-83 (2000). 48. Thomson Healthcare, "MedlinePlus Drug Information: Heparin," 1 (2006). 49. E.J. Topol and P . W . Serruys. "Frontiers in interventional cardiology," Circulation, 98, 1802-20(1998). 50. A . Colombo and E . Karvouni. "Biodegradable stents - "Fulfilling the mission and stepping away"," Circulation, 102, 371-3 (2000). 51. H . Tamai, K . Igaki and E . K y o , et al. "Initial and 6-month results o f biodegradable poly-l-lactic acid coronary stents in humans," Circulation, 102, 399-404 (2000). 52. R. Albiero and A . Colombo. "European high-activity P-32 radioactive stent experience," J. Invasive Cardiol., 12,416-21 (2000). 53. R . N . Saunders, M . S . Metcalfe and M . L . Nicholson. "Rapamycin i n transplantation: A review o f the evidence," Kidney Int., 59, 3-16 (2001). 54. E . K . Rowinsky and R . C . Donehower. "Drug-Therapy - Paclitaxel (Taxol)," N . Engl. J. M e d . , 332, 1004-14(1995). 55. C . J . Creel, M A . L o v i c h and E.R. Edelman. "Arterial paclitaxel distribution and deposition," Circ. Res., 86, 879-84 (2000).  119  56. C21.01, "C242-01 Standard Terminology o f Ceramic Whitewares and Related Products," A S T M International, 57. L . L . Hench, "Bioceramics - from Concept to Clinic," J A m Ceram Soc, 74, 1487-510 (1991). 58. E . B . Riska, "Ceramic-on-ceramic total hip replacement," Current Orthopaedics, 13, 320-4 (1999). 59. R . H . Dauskardt, R . O . Ritchie, J.K. Takemoto and A . M . Brendzel. "Cyclic Fatigue and Fracture in Pyrolytic Carbon-Coated Graphite Mechanical Heart-Valve Prostheses Role of Small Cracks in Life Prediction," J. Biomed. Mater. Res., 28, 791-804 (1994). 60. A . Lasserre and P . K . Bajpai. "Ceramic drug-delivery devices," Crit. Rev. Ther. Drug Carrier Syst., 15, 1-56 (1998). 61. Morris L and Bajpai P K . "Development o f a resorbable tricalcium phosphate (TCP) amine antibiotic composite," In Biomedical Materials and Devices, 293-300 (1989). 62. J.E. Lemons, "Ceramics: Past, present, and future," Bone, 19, S121-8 (1996). 63. A . Tonino, C. Oosterbos, A . Rahmy, M . Therin and C . Doyle. "Hydroxyapatitecoated acetabular components - Histological and histomorphometric analysis o f six cups retrieved at autopsy between three and seven years after successful implantation," Journal of Bone and Joint Surgery-American Volume, 83 A , 817-25 (2001). 64. M . Vallet-Regi and J . M . Gonzalez-Calbet. "Calcium phosphates as substitution o f bone tissues," Progress i n Solid State Chemistry, 32, 1-31 (2004). 65. L . M . Sun, C C . Berndt, K . A . Khor, H . N . Cheang and K . A . Gross. "Surface characteristics and dissolution behavior o f plasma-sprayed hydroxyapatite coating," J. Biomed. Mater. Res., 62, 228-36 (2002). 66. P. Ducheyne, S. Radin and L . K i n g . "The Effect of Calcium-Phosphate Ceramic Composition and Structure on Invitro Behavior .1. Dissolution," J. Biomed. Mater. Res., 27, 25-34(1993). 67. P. Koutsoukos, Z . Amjad, M . B . Tomson and G . H . Nancollas. "Crystallization o f Calcium Phosphates - Constant Composition Study," J. A m . Chem. S o c , 102,1553-7 (1980). 68. M . Jarcho, C H . Bolen, M . B . Thomas, J. Bobick, J.F. K a y and R . H . Doremus. "Hydroxylapatite Synthesis and Characterization i n Dense Polycrystalline Form," J. Mater. Sci., 11, 2027-35 (1976). 69. W . R . Rao and R.F. Boehm. "Study o f Sintered Apatites," J. Dent. Res., 53,1351-4 (1974).  120  70. G . Daculsi, R . Z . Legeros, E . Nery, K . Lynch and B . Kerebel. "Transformation o f Biphasic Calcium-Phosphate Ceramics Invivo - Ultrastructural and Physicochemical Characterization," J. Biomed. Mater. Res., 23, 883-94 (1989). 71. N . Rangavittal, A . R . Landa-Canovas, J . M . Gonzalez-Calbet and M . Vallet-Regi. "Structural study and stability o f hydroxyapatite and beta-tricalcium phosphate: T w o important bioceramics," J . Biomed. Mater. Res., 51, 660-8 (2000). 72. M . Noor Hasimah and O. Radzali. "The effect o f sintering temperature on the mechanical properties o f hydroxyapatite," School of Material and Minerial Resources Engineering, Universiti Sains Malaysia. 73. G . Muralithran and S. Ramesh. "The effects o f sintering temperature on the properties of hydroxyapatite," Ceram. Int., 26, 221-30 (2000). 74. E . Grenadier, A . Roguin and I. Hertz, et al. "Stenting very small coronary narrowings (< 2 mm) using the biocompatible phosphorylcholine-coated coronary stent," Catheterization and Cardiovascular Interventions, 55, 303-8 (2002). 75. R . J . Furlong and J.F. Osborn. "Fixation o f Hip Prostheses by Hydroxyapatite Ceramic Coatings," Journal o f Bone and Joint Surgery-British Volume, 73, 741-5 (1991). 76. S.R. Sousa and M A . Barbosa. "Effect of hydroxyapatite thickness on metal ion release from Ti6A14V substrates," Biomaterials, 17, 397-404 (1996). 77. S.R. Sousa and M . A . Barbosa. "The effect o f hydroxyapatite thickness on metal ion release from stainless steel substrates," Journal o f Materials Science-Materials i n Medicine, 6,818-23 (1995). 78. H . X . J i , C . B . Ponton and P . M . Marqutis. "Microstructural Characterization o f Hydroxyapatite Coating on Titanium," Journal of Materials Science-Materials i n Medicine, 3, 283-7 (1992). 79. S.R. Radin and P. Ducheyne. "Plasma Spraying Induced Changes o f CalciumPhosphate Ceramic Characteristics and the Effect on Invitro Stability," Journal o f Materials Science-Materials in Medicine, 3, 33-42 (1992). 80. G . H . A . Therese and P . V . Kamath. "Electrochemical synthesis o f metal oxides and hydroxides," Chemistry o f Materials, 12, 1195-204 (2000). 81. J. Redepenning, T. Schlessinger, S. Burnham, L . Lippiello and J. Miyano. "Characterization o f electrolytically prepared brushite and hydroxyapatite coatings on orthopedic alloys," J. Biomed. Mater. Res., 30, 287-94 (1996). 82. M . Shirkhanzadeh, "Bioactive Calcium-Phosphate Coatings Prepared by Electrodeposition," J. Mater. Sci. Lett., 10, 1415-7 (1991).  121  83. P. Royer and C . Rey. "Calcium-Phosphate Coatings for Orthopedic Prosthesis," Surf Coat Technol, 45, 171-7 (1991). 84. M . Manso, C . Jimenez, C . Morant, P. Herrero and J . M . Martinez-Duart. "Electrodeposition o f hydroxyapatite coatings in basic conditions," Biomaterials, 21, 1755-61 (2000). 85. M . Shirkhanzadeh, "Calcium-Phosphate Coatings Prepared by Electrocrystallization from Aqueous-Electrolytes," Journal o f Materials Science-Materials i n Medicine, 6, 90-3 (1995). 86. X . L . Cheng, M . Filiaggi and S.G. Roscoe. "Electrochemically assisted coprecipitation o f protein with calcium phosphate coatings on titanium alloy," Biomaterials, 25, 5395-403 (2004). 87. X . H . Hou, X . L i u , J . M . X u , J. Shen and X . H . L i u . " A self-optimizing electrodeposition process for fabrication o f calcium phosphate coatings," Mater Lett, 50, 103-7 (2001). 88. Y . W . Fan, K . Duan and R . Z . Wang. " A composite coating by electrolysis-induced collagen self-assembly and calcium phosphate mineralization," Biomaterials, 26, 1623-32 (2005) . 89. P. Peng, S. Kumar, N . H . Voelcker, E . Szili, R . S . Smart and H . J . Griesser. "Thin calcium phosphate coatings on titanium by electrochemical deposition i n modified simulated body fluid," Journal o f Biomedical Materials Research Part a, 76A, 347-55 (2006) . 90. L . Y . Huang, K . W . X u and J. L u . " A study o f the process and kinetics o f electrochemical deposition and the hydrothermal synthesis o f hydroxyapatite coatings," Journal o f Materials Science-Materials in Medicine, 11, 667-73 (2000). 91. F . H . L i n , Y . S . Hsu, S . H . L i n and J.S. Sun. "The effect o f Ca/P concentration and temperature o f simulated body fluid on the growth o f hydroxyapatite coating on alkalitreated 316L stainless steel," Biomaterials, 23, 4029-38 (2002). 92. K . K o s k i , J. Holsa, J. Ernoult and A . Rouzaud. "The connection between sputter cleaning and adhesion o f thin solid films," Surf Coat Technol, 80, 195-9 (1996). 93. M . Shirkhanzadeh, "Electrochemical Preparation o f Bioactive Calcium-Phosphate Coatings on Porous Substrates by the Periodic Pulse Technique," J. Mater. Sci. Lett., 12, 16-9(1993). 94. U . S . Food and Drug Administration, "Selected Guidance Documents Applicable to Combination Products- Draft Guidance for the Submission o f Research and Marketing Applications for Interventional Cardiology Devices: P T C A Catheters, Atherectomy Catheters, Lasers, Intravascular Stents,"  122  95. S. Ban and S. Maruno. "Effect o f P h Buffer on Electrochemical Deposition o f Calcium-Phosphate," Japanese Journal o f Applied Physics Part 2-Letters, 33, L I 545-8 (1994). 96. J. Redepenning and J.P. Mcisaac. "Electrocrystallization o f Brushite Coatings on Prosthetic Alloys," Chemistry o f Materials, 2, 625-7 (1990). 97. T.S. Light, S. Licht, A . C . Bevilacqua and K . R . Morash. "The fundamental conductivity and resistivity o f water," Electrochemical and Solid State Letters, 8, E l 6-9 (2005). 98. M . C . K u o and S . K . Y e n . "Immersion characteristics o f electrolytic calcium phosphate coated T i in simulated physiological fluid," J. Mater. Sci., 39, 2357-63 (2004). 99. A . Strawbridge and H . E . Evans. "Mechanical Failure o f T h i n Brittle Coatings," Eng. Failure Anal., 2, 85-103 (1995). 100. M . Kumar, H . Dasarathy and C . Riley. "Electrodeposition o f brushite coatings and their transformation to hydroxyapatite i n aqueous solutions," J. Biomed. Mater. Res., 45, 302-10 (1999). 101. Y . Han, K . W . X u and J . A . L u . "Morphology and composition o f hydroxyapatite coatings prepared by hydrothermal treatment on electrodeposited brushite coatings," Journal o f Materials Science-Materials in Medicine, 10, 243-8 (1999). 102. M . Leoni and P. Scardi. "Nanocrystalline domain size distributions from powder diffraction data," Journal o f Applied Crystallography, 37, 629-34 (2004). 103. M . P . Mahabole, R . C . Aiyer, C . V . Ramakrishna, B . Sreedhar and R . S . Khairnar. "Synthesis, characterization and gas sensing property o f hydroxyapatite ceramic," B u l l . Mater. Sci., 28, 535-45 (2005). 104. O . E . Petrov, E . Dyulgerova, L . Petrov and R. Popova. "Characterization o f calcium phosphate phases obtained during the preparation o f sintered biphase Ca-P ceramics," Mater Lett, 48, 162-7 (2001). 105. C . Richard Brundle, Charles A . Evans, Jr. and ShaunWilson. "Surfaces, Interfaces,  Thin Films," pp. 120-9 in Encyclopedia of Materials Characterization. Edited by Lee E . Fitzpatrick. Butterworth-Heinemann, Stoneham, M A , 1992.  123  APPENDIX  A  -  PRELIMINARY  RESULTS  ON  ECD CO-  DEPOSITION O F O R G A N O - C E R A M I C C O A T I N G S The study o f electrochemical co-deposition was not fully complete in this part of the research  project,  however, preliminary experimental work was performed  investigate the feasibility o f such process. deposit  calcium phosphate  coating  with  to  The objective o f this study was to (i) coa  polymer  for  mechanical  properties  enhancement; (ii) confirmation o f the possibility o f 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 o f substrate surface treatment.  Polyvinyl alcohol was dissolved in the  standard E C D electrolyte before deposition.  T w o 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. demonstrated good uniformity and full coverage.  Both o f these coatings  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 o f 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 o f CaP alone on bare metal stent as illustrated i n Figure 5.2-10.  These results  confirmed the possibility o f E C D co-deposition with P V A for coating integrity enhancement. Although the fundamental mechanism o f 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 x 9 0 0  iUum  Figure A 1. S E M images o f an E C D coating co-deposited with 0.1 wt% o f 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 o f 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  DETAILED 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: ECDCqating 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):  Sample Name XD-RP-OOI-OI-  Yes 500 °C T 2 0"min  Bare Stent Weight (mg) : 5. SOT-  Final Stent Weight (mg) 15:. 6,63-  Coating Weight  IS.716-  15.779  63  Pass  15.811:-.  66  Pass  (ug)  Pass/Fail Pass  JD-RP-COi-O/ ]CD-RP-.001-03:  :15.7.45-  C O - R P 0 0 1 3-1  16.087  16:i52"  65'  Pass  :D-RP-001-05:  16.187  16.251.  64  Pass;  icD-RP-oor-O'e' -JD-RP-001-07  16.219  •16.279  60  Pass-  15:.789  15.850  61  Pass  ECD-RP-001-08-  15-.;852-  15.914  62'-  Pass  .ECD-RP^001'-09: ECDr-RP-001-10  16.716  16.782  66.  1 5 . 7 98  15.863  65  Average:  15.971 0.317  16.034  63.4  S t a n d a r d Dev.  :  Pass  Yield Rate of ECD-RP-001 -100'  128  Table B 2. Preparation record for E C D coating batch E C D - R P - 0 0 2 .  Substrate r  Marks & Identity: Ref No. Size:  MIVI 700 S e r i e s C o r o n a r y  Description:  Stent  ECD Coating Preparation Date: Batch Nurrben  J a n u a r y 2 3 , 2006 E C D - R P - OO'™""  "l^^^^^^^Bli^BSiB !^!! 1  ' [.. '  Prepared By: Electrolyte Ca:P Ratio: Appliucl Cum-nt (mA): bl--ctn*kte Temp. (°C): Eli'utr»Kti.' pH: Deposition Time (mm) Surface Treatment Yes*  Pli.i-.t- Cunvuisiuii: Sintering Temp. ("C):  500  C {20 m i n )  Sample Name  Bare Stent Weight (mg)  ECD-RP-002-0:l  15.768  ;ECD-R-P-002-02  15.785  Final Stent Weight (mg) ~15 . 8 2 F "~ 15.847  y, •• • /",• • <  Coating Weight (UCjl  60  Pass/Fail Pass  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  ECD^RP-002-07ECD-RP-002-  15 . 982'  16.0,43  Pass Pass Fail Pass Pass  Table B 3. Preparation record for E C D coating batch E C D - R P - 0 0 3 .  Substr.ittMarks & Identity: Ref No. Size:  MIV I ~7 0^"^Tr i e7"F<^onl  Description:  Pri'|j;iratiu:i Ddtu:  ECDL Coating January 31, 2006  Batch Number  liiliM^iiiiiiMli^^^B^i^^W^i^^psi^^S  ECD-RP-003  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: lTo^^^^^^^^^^^tlilllilltiril^^  Deposition Time (mm): Surface Treatment Phase Conversion: Sintering Temp. ("C):  "  B.iro Stunt Final Stent Weight (mg) Weight (mg) 16.024 16.085 ECD-RP-003-01 ECD-RP-003-02 15.655 15.714 ECD-RP-003-03 15 . 979 16.040 ECD-RP-003-04 16.246 16.310 16 897 F.C:; ? . ? 0 0 3 Ob . 16.970 ECD-RP-O.03:-O:6:: •• • 15.786 15.847 ECD-RP-003^.0-7 • :5.303 -1-5-.-2 43- - ••; ECD-RP-003-08 16.099 16 161 16 565 ECD-RP-003'-0;9 16 628 15.724" ECD-RP-003-10 35.662 • Sample Name  ••-  te •  Pi ' :  Pi  Average:  16.016  16.07B  Standard Dev.  0". 452  ' ' V 0.455  Yield Rate of  ECD-RP-003=90°l  Coating Woight (ug) 61 59 61 64 73 61 60 62 63 : :62 :  Pass/Fail Pass Pass Pass Pass Fail -Pass . Pass '. • "PaSS': •:: .Pass..  •  Pass  Table B 4. Preparation record for E C D coating batch E C D - R P - 0 0 4 .  IF  Substr.ito M  irln  <K  I.  lil>-nllt\  <:  Rtf Nn  Description:  14 mm MIVI 700 S e r i e s Coronary S t e n t  -:&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:  ^^^^^^^^^^^^^^^^^^^^^^ ^^^^^^^ip^^^H^i^i^^Bi^^^S^^^KS  4 .5  Deposition Time (min): Surface Treatment Ph.ise Conversion: Sintering Temp. (°C):  "500 "°c"720 min)  Bjrr> Stent Weight (mg)  Sample Name  ECD-RP-004-01 :" ; D - R P - O O 4 - 0 2  .CD-RP-004-03 CD-RP-0.04:-0,4.' •••CD-RP-004.^05 - ED-RP-0.04-0:6 • CD-RP-004-07 •.CD-RP-004-08 ECD'-RP-0:04'-0:9 ECD-RP-004-10: Average Standard Dev.  !  :  Final Stent Weight (mg)  15.341 15.277 15 .'349- ' 16.084 . • 16.021 16.797 16.730 16.041 15.97 6 ' 16.053 15.991 16.767' ' • = 16.828 16.828 16.765 15.490 15.429 15.657 :. :i 5:. : • ..16.056 15 996 0.557 0.561  Yield'Rate of ECD-RP-004 ' =90  :  • • '  Coating Weight (ug)  Pass/Fail  64 Fa J. 1 63 ••' . ' . Pass 67 Pass Pans 65 Pass 62 Pass 61 Pass 63 Pass 61 66 63. 6 2.0  Table B 5. Preparation record for E C D coating batch E C D - R P - 0 0 5 .  Substrate Marks & Identity:  :  i  Ref No.  i  .1  Size: MIVI 700 S e r i e s C o r o n a r y Ster  Description:  ECD Coatinq Preparation Date:  t  February 13, 2 0  J ^ ^ ( i | l | ! | f lll^^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):  Final Stent Weight (mg) 15.832  Coating Weight  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  • 16.024  16.083  59'  Pass  Average:  16.126  16.188  62.7  S t a n d a r d Dev.  0.375  0.375  2.2  Sample Name 1  |  ECD-RP-005-01 ECD-RP-005-02  !  |  • iEGD-RP-005-10  Bare Stent Weight (mg) '. 15.767  Yield Rate of ECD-RP-005 =100%  (ug) 65  Pass/Fail Pass  

Cite

Citation Scheme:

    

Usage Statistics

Country Views Downloads
United States 15 0
Germany 10 0
France 7 0
Russia 5 0
China 5 22
United Kingdom 2 8
India 2 0
New Zealand 1 0
Taiwan 1 0
Japan 1 0
City Views Downloads
Unknown 11 0
Bad Soden am Taunus 9 0
Boardman 6 0
Saint Petersburg 5 0
Ashburn 5 0
Shanghai 2 0
Jalalpur 2 0
Denver 1 0
Bath 1 8
Mountain View 1 0
Hamilton 1 0
Redmond 1 0
Putian 1 0

{[{ mDataHeader[type] }]} {[{ month[type] }]} {[{ tData[type] }]}
Download Stats

Share

Embed

Customize your widget with the following options, then copy and paste the code below into the HTML of your page to embed this item in your website.
                        
                            <div id="ubcOpenCollectionsWidgetDisplay">
                            <script id="ubcOpenCollectionsWidget"
                            src="{[{embed.src}]}"
                            data-item="{[{embed.item}]}"
                            data-collection="{[{embed.collection}]}"
                            data-metadata="{[{embed.showMetadata}]}"
                            data-width="{[{embed.width}]}"
                            async >
                            </script>
                            </div>
                        
                    
IIIF logo Our image viewer uses the IIIF 2.0 standard. To load this item in other compatible viewers, use this url:
http://iiif.library.ubc.ca/presentation/dsp.831.1-0078873/manifest

Comment

Related Items