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The effect of organic solvents on sol-gel hydroxyapatite and its application as biocoating Hakimimehr, Dorna 2001

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THE EFFECT OF ORGANIC SOLVENTS ON SOL-GEL HYDROXY APATITE ' AND ITS APPLICATION AS BIOCOATING by DORNA HAKIMIMEHR B.A-.Sc, TEHRAN UNIVERSITY OF SCIENCE AND TECHNOLOGY, Iran, 1999 A THESIS SUBMITTED IN PARTIAL FULFILMENT OF THE REQUIREMENTS FOR THE DEGREE OF MASTER OF APPLIED SCIENCE in THE FACULTY OF GRADUATE STUDIES (Department of Metals and Materials Engineering) We accept this thesis as conforming to the required standard THE UNIVERSITY OF BRITISH COLUMBIA October 2001 © Dorna Hakimimehr, 2001 UBC Special Collections - Thesis Authorisation Form http://www.library.ubc.ca/spcoll/thesauth.html In p r e s e n t i n g t h i s t h e s i s i n p a r t i a l f u l f i l m e n t of the requirements f o r an advanced degree at the U n i v e r s i t y pf B r i t i s h Columbia, I agree that the L i b r a r y s h a l l make i t f r e e l y a v a i l a b l e f o r reference and study. I f u r t h e r agree that permission f o r extensive copying of t h i s t h e s i s f o r s c h o l a r l y purposes may be granted by the head of my department or by h i s or her r e p r e s e n t a t i v e s . I t i s understood that copying or p u b l i c a t i o n of t h i s t h e s i s f o r f i n a n c i a l g a i n s h a l l not be allowed without my w r i t t e n permission. Department of tvAj VoAs OwrA tAo\Wf.CT.U &r\J\.ypor.V The U n i v e r s i t y of B r i t i s h Columbia Vancouver, Canada Date O f 4 | | l7Qp\ 1 of 1 10/11/2001 2:37 PM 11 ABSTRACT Hydroxyapati te (HAp) is widely used by the biomedical industry due to its excellent biocdmpatibil ity. The sol-gel p rocess offers a relatively low temperature procedure to produce hydroxyapatite powders and thin fi lms (<1um). The focus of the present study is to investigate the effect of different organic solvents (methanol, ethanol, and propanol) on the sol-gel H A p in the system of triethyl phosphite and ca lc ium nitrate. X -ray diffraction analys is , thermal gravimetric analys is , differential thermal analys is , and scanning electron microscopy (SEM) are used to investigate the phase evolution and the morphology of the HAp produced in this sol-gel sys tem. Our results show that different organic solvents induce different H A p formation pathway. In methanol and propanol -based sys tems HAp forms as a result of transformation of intermediate crystall ine ca lc ium phosphate phases (such as C a 3 ( P 0 4 ) 2 and C a 2 P 2 0 7 ) . Formation of H A p in e thanol -based system is attributed to crystall ization from an amorphous, intermediate apatite phase. Titanium substrates were coated using the HAp so l . IR spect roscopy of the coatings revealed that carbonated hydroxyapatite is present in the coat ing, which is similar to that of the natural bone. In-vitro bioactivity test conf i rmed the bioactive nature of the coat ings while no apparent difference was observed in the bioactivity of the coat ings obtained from different solvent sys tems. The type of solvent, concentrat ion of the solut ion, and heat treatment time affect the quality of the coat ings on stainless steel wires. The coat ings obtained from low concentrat ion (1M) methanol -based solution, heat treated at 500°C for a period of 10 minutes were the most satisfactory. Coronary stents were also coated with HAp using this sol-gel sys tem. The coat ing remained on the stent after the expans ion of the stent using an angioplasty bal loon. i i i TABLE OF CONTENTS Abstract ii Table of contents iii List of Figures vi Aknowledgments viii Chapter 1 1 Introduction 1 1.1 Biomaterials 1 1.1.1 H y d r o x y a p a t i t e 2 1.2 Sol-gel process 3 1.3 Focus of the present study 4 Chapter 2 6 Literature review 6 2.1 Biomaterials 6 2.2 Calcium phosphate ceramics 7 2.2.1 M e c h a n i c a l p r ope r t i e s 8 2.2.2 B i o r e s o r p t i o n a n d b i o d e g r a d a t i o n 9 2.2.3 M e c h a n i s m of n e w b o n e f o r m a t i o n 10 2.3 Calcium phosphate bone cements 11 2.4 Hydroxyapatite-based biomaterials 11 2.4.1 H A p p o w d e r s 11 2.4.2 D e n s e H A p c e r a m i c s 12 2.4.3 P o r o u s H A p c e r a m i c s 12 2.4.4 H A p - b a s e d c e r a m i c c o m p o s i t e s 12 2.4.5 H A P / b i o a c t i v e g l a s s c o m p o s i t e s 13 2.4.6 H A p c o a t i n g s 13 2.4.7 H A P / p o l y m e r c o m p o s i t e s 14 2.5 Sol-Gel HAp 14 iv 2.5.1 The Sol-Gel process 14 2.5.2 Ceramic coatings by sol-gel 17 2.5.3 Sol-gel HAp materials 20 2.5.4 Sol-gel HAp coatings 24 2.6 Stents 28 Chapter 3 33 Scope and objectives 33 3.1 Scope of investigation 33 3.2 Objectives 34 Chapter 4 36 Experimental methodology 36 4.1 General procedures 36 4.1.1 Materials 36 4.1.2 Analysis 37 4.2 Powder preparation and characterization 37 4.3 Coating preparation and characterization 39 4.3.1 Titanium substrates 39 4.3.2 Stainless steel substrates 39 4.3.2.a Wires 39 4.3.2.b Stents 40 Chapter 5 42 Experimental Results and Discussion 42 5.1 Phase evolution in sol-gel hydroxyapatite 42 5.1.1 Results 42 5.1.1 .a X-ray diffraction 42 5.1.1.b Thermal Analysis 46 5.1.1 .c Electron microscopy 48 5.1.2 Discussion 49 5.2 HAp coating characterization 53 5.2.1 Coatings on titanium substrates 53 5.2.1 .a Electron microscopy examination 53 5.2.1.b Infrared spectroscopy 57 V 5.2.1 .c In-vitro bioactivity test 58 5.2.2 Stainless steel substrates 60 5.2.2.1 Stainless steel wires 60 5.2.2.1 .a surface treatment 60 5.2.2.1 .b Effect of coating solution 62 5.2.2.1 .c Effect of heat treatment time 65 5.2.2.2 Stents 68 Chapter 6 73 Summary and conclusions 73 6.1 Summary .'. 73 6.2 Conclusions 76 Chapter 7 80 Recommendation for future work 80 Nomenclature 81 References 82 Appendix 1 88 Appendix II 89 VI LIST OF FIGURES Figurel .1 Schematic of a sol-gel processes and ceramic products 3 Figure 2.1 Phase diagram of CaO and P 20 5 binary system. From [2] 8 Figure 2.3 Stages of the spin-coating process. From [12] 18 Figure 2.2 (A) stages of dip coating process: (a-e) batch, (f) continuous. (B) Detail of the liquid flow patterns in area 3 of the continuous process. U is the withdrawal speed, S is the stagnation point, 5 is the boundary layer, and, h is the thickness of the fluid film. From [12] 19 Figure 2.4 Stent Placed inside the artery 29 Figure 4.1 Flow chart of so-gel HAp synthesis procedure 38 Figure 5.1 XRD patterns for methanol-based gels calcined at (a) 375°C (b) 425°C, (c) 445°C, (d) 460°C, (e) 500°C and (f) 545°C 43 Figure 5.2 XRD patterns for ethanol-based gels calcined at (a) 375°C (b) 425°C, (c) 445°C, (d) 460°C, (e) 500°C and (f) 545°C 44 Figure 5.3 XRD patterns for propanol-based gels calcined at (a) 375°C (b) 425°C, (c) 445°C, (d) 460°C, (e) 500°C and (f) 545°C 45 Figure 5.4 Thermal gravimetric analysis of the dried gels 46 Figure 5.5 Differential thermal analysis of the (a) methanol, (b) ethanol, and (c) Propanol-derived dried gels 47 Figure 5.6 Scanning electron micrographs of (a) methanol-based, (b) ethanol-based, (c) propanol-based gels calcined at 500°C 48 Figure 5.7 Scanning electron micrograph of (a) surface of titanium plate, treated with phosphoric acid, and (b) titanium surface after sandblasting 53 Figure5.8 Surface of chemically treated in phosphoric acid (85%) at 50-60°C for 30 minutes titanium substrates after spin coating with (a) methanol, (b) ethanol, and (c) propanol-based solutions 54 Figure 5.9 Energy dispersive spectrometry of (a) titanium etched surface and titanium surface after coating with (b) methanol, (c) ethanol and (d) propanol-based solutions 55 Figure 5.10 Sandblasted titanium surfaces after coating with solutions obtained from (a) methanol, (b) ethanol and (c) propanol solvent systems 56 Figure 5.11 FTIR spectra of the surface of sandblasted titanium substrates after coating with (a) methanol, (b) ethanol and (c) propanol-based solutions 58 Figure 5.12 SEM micrographs of surface morphology of titanium substrates coated with (a) methanol, (b) ethanol, and (c) propanol-based solutions, after incubation in SBF 59 Figure 5.13 SEM micrographs of (a) as received stainless steel wire and wires after exposure to (b) nitric acid, (c) phosphoric acid, and (d) hydrochloric acid...61 V l l Figure 5.14 SEM micrographs of stainless steel wires treated with HCI (2.4 N) for (a) 10, (b) 20, (c) 30, (d) 40, (e) 50, and (f) 60 minutes at 75°C and neutralized for 4 minutes. 62 Figure 5.15 SEM micrographs of bent stainless steel wire samples, dip coated in (a) methanol, (b) ethanol, and (c) propanol-based solutions after unbending 64 Figure 5.16 SEM picture of bent wire samples coated with dilute (1M) (a) methanol, and (b) ethanol-based solutions after unbending 65 Figure 5.17 SEM picture of unbent stainless steel wires coated with dilute methanol solution and fired for (a) 10, (b) 30, and (c) 60 minutes at 500°C 66 Figure 5.18 XRD results on thin film coatings fired at 500°C for different time periods. 67 Figure 5.19 (a) SEM picture of the stent surface obtained after de-oxidizing process, (b) EDS of the ticker layer on the stent surface 68 Figure 5.20 (a) and (b) show SEM micrographs of the de-oxidize stent surface after coating with dilute methanol solution and fired at 500°C for 10 minutes 69 Figure 5.21 EDS results of the (a) front and (b) interior surfaces of the coated stent structure 70 Figure 5.22 (a-b) SEM pictures of the coated stent after expansion 71 Figure 5.23 Electropolished stent (a) prior, and (b) after etching in HCI at 70°C for 10 minutes 71 Figure 5.24 Electropolished stent (a) prior, and (b) after etching in HCI at 70°C for 10 minutes 72 V l l l AKNOWLEDGMENTS I would like to thank T o m Troczynsk i for his supervis ion of this project and also Dean -Mo liu for shar ing his knowledge of this field with me. I would also like to thank MIVI technologies and N S E R C for funding the current study. 1 C H A P T E R 1 I N T R O D U C T I O N Ceramics and g lasses have been used in health-care industry for a long time. The main advantages of ceramics in medical appl icat ions are their res istance to microbial attack, pH changes , solvent condit ions and temperature. They have also been appl ied in dentistry and as a hard t issue replacement in musculo-skeleta l sys tems such as bone and teeth. However their main application remains to be as bioactive coat ings on various prostheses in order to add bioactive character ist ics to the otherwise inert implants. Sol -gel p rocess ing s e e m s to be one of the best methods to produce thin film coat ings at low temperatures, with controlled microstructures. Lower temperature process ing is the most attractive characterist ic of the sol-gel technique in b iomedica l appl icat ions, as high temperature process ing of the coating might adversely affect the mechanica l properties of the substrate. The microstructure and phase development of the coat ing, that affect coat ing's bioactivity as well as its long-term per formance in biological environment, are well control led in the sol-gel p rocess . P rocess development and characterizat ion of hydroxyapatite (HAp) b ioceramics through sol-gel has been the main focus of this work. The knowledge thus obtained is appl ied in coat ing severa l medical substrates. S o m e of these attempts, such as HAp coating on coronary stents, are novel appl icat ions of hydroxyapatite coat ings. 1.1 Biomaterials A biomaterial is "any material used to replace or restore function to body t issue and is cont inuously or intermittently in contact with body fluids" [1]. Biomater ia ls are divided in to different categor ies according to the responses they elicit in the body, including biologically inert, porous, resorbable, and bioactive materials. Bioact ive materials "elicit a speci f ic biological response at the interface of the material which results in the formation of a bond between the t issues and the material [2]". 2 1.1.1 Hydroxyapatite Calc ium phosphate materials, and hydroxyapatite (HAp) in particular, are among the most widely used biomaterials in b iomedical appl icat ions. Hydroxyapat i te is the most appropriate ceramic material for artificial teeth or bones due to its excellent biocompatibil i ty and bioactivity. H A p interacts with body fluids through complex dissolution-reprecipitation process . In a simplif ied descript ion [2], a H A p implant, or an implant coated with HAp , will produce a local decrease of pH (an "acidic condition") at the implant a rea upon introduction to the biological environment. Th is acid ic condit ion will promote the dissolution of the HAp , i.e. re lease of C a 2 + , P 0 4 3 " and H P 0 4 2 " , which will saturate the microenvironment with calc ium phosphates. This saturation leads back to deposit ion of a layer of carbonate-containing apatite through s e e d e d growth on the HAp crystals. The rate of formation of this carbonated hydroxyapatite determines the extent of the bioactivity of the implant and is control led by the crystallinity of the synthetic H A p . The dissolution of H A p is a lso governed by chemica l composi t ion, crystal structure and microporosity of the material [2]. Despite excel lent bioactivity and biocompatibil i ty of hydroxyapati te, its poor mechanica l properties restrict its appl icat ions to powders (e.g. bone loss fillers), coat ings on metallic substrates, or implants under very low loads (e.g. inner ear implants [3]). Coa ted implants combine good mechan ica l strength of the metall ic substrate with excellent bioactive properties of the HAp . Uncoated metall ic implants do not integrate with the bone and are encapsu la ted by dense fibrous t issue, which prevents proper stress distribution and will cause implant loosening [4]. In the c a s e of H A p coated prostheses, the bone will integrate with the implant providing a stable fixation [5]. The HAp coat ing will a lso prevent the f ibrous t issue encapsulat ion of the implant while decreas ing the re lease of metal ions from the implant into the body and protecting the metal sur face from environmental attack [6]. Many techniques have been adopted and studied in order to coat a H A p layer on a metallic substrate. P l a s m a spraying [7], hot isostatic press ing [8], ion beam deposit ion [9], f lame spraying [10], and electrochemical deposit ion [11] are among the many. S o l -gel process ing offers an alternative process in coat ing metall ic implants. The 3 advantages of this p rocess are the low temperature synthesis and microstructural control. 1.2 Sol-gel process Sol-gel is a multi-step process in which a col loidal solution is prepared through hydrolysis of a suitable precursor. The starting precursor in the sol-gel process is usually a metall ic alkoxide d isso lved in a lcohol , and then hydrolysed through controlled addition of water [12]. Figure1.1 Schemat ic of a sol-gel p rocesses and ceramic products. The hydrolysis products (complex hydroxides) condense and form growing clusters through polymerizat ion. The clusters join each other and as a result an interlocking 4 phase cal led "gel" is formed, which can be easi ly shaped . The sol-gel process ing of ceramics offers a number of advantages over "traditional" powder process ing of ceramics , such as increased purity and homogenei ty (mixing occurs on the atomic level), reduced sintering temperatures due to smal l particle s ize , and ability to coat complex shapes . D isadvantages of sol-gel process ing are high cost of raw materials, large shr inkage accompany ing drying and sintering, and relatively long process ing time for 3D objects [12]. Thin fi lms produced through sol-gel process ing overcome many of these intrinsic d isadvantages of the technique. The early application of sol-gel coat ings was in optical dev ices. Many new uses of sol-gel fi lms have appeared in electronic components , protective fi lms (e.g. corrosion and wear), membrane and senso r appl icat ions as well as biomaterials. Thin sol-gel fi lms can be produced using smal l amount of raw materials, can be p rocessed relatively quickly without cracking while large substrates can be easi ly accommoda ted . Therefore thin film coat ing has become one of the few commerc ia l appl icat ions of the sol-gel technology [12]. 1.3 Focus of the present study Although many controll ing parameters such as starting materials, aging time and heat treatment time and temperature in the sol-gel synthesis of hydroxyapatite have been investigated (reviewed in sect ion 2.5) very few studies have add ressed the effects of solvents on the H A p phase evolution and coat ing character ist ics. In this work we are looking into the effects of organic solvents on a novel sol-gel sys tem, which produces hydroxyapatite at relatively low temperatures (400-500°C). Employ ing different solvents in preparation of the solution affected the phase formation, purity, microstructure and morphology of the final sol-gel H A p product. The ability to control the microstructure and morphology of the resulting hydroxyapatite translates to a control of its bioactivity. The solut ions made using different solvents a lso had different viscosity and surface tension, which affected HAp coat ing formation on metallic substrates. The knowledge obtained from this fundamental study was then appl ied to deposit H A p coat ings on medical substrates such as titanium and medical grade 5 stainless steel wires. Coat ing of coronary stents with hydroxyapati te through this so l -gel p rocess was also attempted. 6 CHAPTER 2 LITERATURE REVIEW 2.1 Biomaterials G l a s s e s and ceramics have been appl ied in the b iomedical industry due to their compatibil ity with the biological environment and their res istance to environmental attack. The appl icat ions of ceramics include replacements for hip [13], knees [14], teeth [15], tendons and l igaments [16] and repair for periodontal d i sease [17], maxil lofacial reconstruction [18], augmentat ion and stabil ization of jaw bone [2], spinal fusion [19] and bone fillers after tumour surgery [2]. They have also been appl ied in dentistry and as a hard t issue replacement in muscu loske le ta l sys tems such as bone and teeth, but their cl inical s u c c e s s requires a good interface between the implant and the t issue and a lso a good match of mechanica l properties at this interface [2]. Depending on the response an implant elicits in a living t issue different mechan isms of implant- t issue attachments have been categor ized. Accord ing to I. Hench [2] these mechan isms can be summar ized as follow: 1- Nearly inert dense biomaterials. In this category attachment occurs through the growth of the hard t issue into the irregularities of the sur face by cement ing the device into the t issue (morphological fixation). In this category of biomaterials the movement that occurs between the implant and the hard t issue usually leads to deterioration of the function of the implant. A lumina, both in the form of single and polycrystal l ine form, is in this group. 2- Porous inert implants. In this case the attachment is ach ieved by the ingrowth of the bone into the pores of the material (biological fixation). This ingrowth provides a larger area of attachment and consequent ly creates an increased res is tance to movement. In order to keep the t issues in pores healthy, the pores should be 100-150u.m in diameter to provide the t issues enough blood flow to survive. Th is requirement of the porous implant will lead to low mechanica l strength and therefore limits the application of implants of this type to coat ings and unloaded appl icat ions. Hydroxyapat i te-coated porous metals can be cons idered in this category. 7 3- Dense , nonporous (porous) resorbable ceramic. In this c a s e ceramic implant is des igned to be slowly replaced by the bone. The criteria for this type of implants are first, to maintain the strength of the implant throughout the degradat ion process and second to control the degradat ion rate in order to match it with the rate of t issue replacement. Ca lc ium sulfate and tricalcium phosphate are good candidates for resorbable biomaterials. 4- Dense nonporous materials with reactive sur faces. T h e s e implants will chemical ly bond to bone (resulting in bioactive fixation). The term bioactive is used for those categor ies of biomaterials, which "elicit a specif ic biological response at the interface of the material which results in the formation of a bond between the t issues and the material" [1]. Bioact ive g lasses , bioactive g lass-ceramics and hydroxyapatite (HAp) are in this group of implants. 2.2 Calcium phosphate ceramics Different phases of calc ium phosphate ceramics are used in either resorbable or bioactive materials. A slight difference in composi t ion can change a biomaterial from resorbable to bioactive or even inert [2]. The type of stable phases of calc ium phosphate ceramics , depend on temperature and presence of water either during process ing or while in use . Figure 2.1 shows the phase d iagram for C a O and P2O5. A s it can be seen H A p is the stable phase till 1360°C when water partial pressure is - 6 6 K P a , without water C4P and C3P are stable phases . A s it can be seen in F ig . 2 .1 , by increasing the water pressure the temperature range of H A p stability increases [2]. Two phases of ca lc ium phosphate materials are stable at the body temperature and in aqueous media . For pH < 4.2, C a H P 0 4 . 2 H 2 0 (dicalcium phosphate, D C P ) and for pH >4.2, Ca io ( P 0 4 ) 6 (OH) 2 (HAp) are the stable phases . At higher temperatures C a 3 ( P 0 4 ) 2 ((3-tricalcium phosphate, (3-TCP) and C a 4 P 2 0 g (tetracalcium phosphate, T T C P ) are stable. The high-temperature phases will interact with water or body fluids at 37°C to form H A p [2]. The reaction of formation of H A p on the sur face of tricalcium phosphate (TCP) is proposed as fol lows [2]: 8 4 C a 3 ( P 0 4 ) 2 ( s ) + 2 H 2 0 ^ C a i 0 ( P O 4 ) 6 ( O H ) 2 ( s u r f a c e ) + 2 C a 2 + + 2 H P 0 4 2 " (2.1) Accord ing to the above reaction the solubility of T C P approaches the solubility of HAp . Micropores present in the sintered implants will increase the solubility of T C P as well as the increase in pH [2]. 1700-1 1600 H Z, 1500 s Z3 2. E 1200H T 70 0 r'CaP+liquid \ liquid / \ C 4 P \ i + \ J liquid \7 1570° \ 1 5 5 0 ° \ a'C3P + CJP C a O + C « P u C P HA \ PI 1300 l l l l l liquid (-CP HA C a O + H A C 4P l l l l ! oCaP + CaP I 65 60 H A C3P — C a O (wt%) 50 Figure 2.1 P h a s e diagram of C a O and P 2 0 5 binary sys tem. From [2]. 2.2.1 Mechanical properties Mechan ica l propert ies of ca lc ium phosphate ceramics are the main factor affecting their b iomedical appl icat ions. The presence of porosity either in the form of micropores (<1um) or macropores (>100u.m) dec reases both the compress ive and tensi le strength [2]. The following equat ions elaborate on this dependence : o c =700 exp(-5V p ) (2.2) 9 a t =220 exp( -20V m ) (2.3) Where a c , V P i a t a n d V m are the compress ive strength, total pore vo lume fraction (V p=0-0.5), tensi le strength and volume fraction of microporosity, respectively [2]. The compress ive strength of dense H A p ceramics falls in the range of 120-900 M P a and the range for tensi le strength is 38-300 M P a [6]. The low mechan ica l properties of ca lc ium phosphate ceramics restricted their appl icat ions to (1) powders, (2) dental implants with reinforcing metal posts, (3) coat ings, (4) smal l un loaded implants, (5) low loaded porous implants and (6) polymer-bioactive ceramic compos i tes [20]. 2.2.2 Bioresorption and biodegradation Calc ium phosphate (Ca-P) materials undergo a bioresorption and biodegradation process after being exposed to biological environment. "Biodegradat ion is the process caused by the action of living sys tems (e.g. microorganis ims, cells) when material breaks down into its s impler components ; reduces the complexi ty of a chemica l compound or wears away by eros ion" [20]. The C a - P implant material undergoes both physical and chemica l changes when introduced to the biological environment. Phys ica l changes include disintegration, changes in micro and macroporosi ty of the implant and also change in the s ize and/or weight of the implant. Chemica l changes primarily include reduction of pH in the implant environment thus changing its solubility. Consequent ly concentrat ion of C a and P ions increases in the implant microenvironment beyond the saturation level, resulting in deposit ion of a layer of C a - P material on the sur face of the implant or their incorporation into the new bone formed at implant/bone interface. Bioactivity of a material is def ined as its ability to form a C a - P compound on the implant surface when introduced to biological environment saturated with C a and P ions, and therefore is related to the biodegradat ion of the material [20]. The C a / P ratio of the starting precursor materials is an important factor in determining the C a - P phases occurring in the final product, and thus affecting the solubility, biodegradation and finally bioactivity of the implant material. If the C a / P ratio of the 10 starting material is less than 1.67 (i.e. the stoichiometric value for HAp) , (3-TCP forms along with H A p . If C a / P ratio is higher than 1.67, the product will contain C a O along with H A p [20]. Al though sintering temperature does not affect the [3-TCP/HAp ratio of the final product, it changes the crystallinity of the phases , which will affect the biodegradat ion or biodissolut ion of the C a - P materials. It is reasonab le to bel ieve that lower crystallinity will lead to higher dissolution rates [20]. Accord ing to L e G e r o s [20], the parameters affecting the solubility of C a - P materials include (1) physical parameters (e.g. density), (2) composi t ion, i.e. different composi t ion and mixtures of C a - P materials can greatly affect the extent and rate of the dissolut ion of the implant, (3) crystal structure and (4) crystallinity. The orders of relative solubility of s o m e C a - P compounds are as fol lows: A C P > D C P > T T C P > a - T C P > p - T C P > H A p Where A C P is "amorphous ca lc ium phosphate". Bone cel ls will readily attach to C a - P materials [20]. Ce l l proliferation and increased D N A synthesis were also observed due to increase in intracellular C a + 2 ion concentrat ion [21]. The increase in the concentrat ion of ca lc ium and phosphate ions resulting from dissolut ion of C a - P materials a lso affects bone-cel l activity [20]. One of the most important effects is the increased inhibition of bone resorption due to reduced osteoclast 1 formation and dec reased activity of mature osteoclasts [20]. 2.2.3 Mechanism of new bone formation The dissolut ion of the C a - P in acidic environments will lead to an increase in the concentrat ion of ca lc ium and phosphate ions in the surrounding solut ion. This will eventually lead to precipitation of apatite microcrystals, usual ly incorporating other ions such as M g 2 + , CO32", etc. a long with the organic molecu les. Other non-apatite C a - P phases such as dicalc ium phosphate dihydrate ( D C P D ) and octaca lc ium phosphate ( O C P ) , which are more stable toward dissolut ion, may a lso hydrolyse to C 0 3 -containing apatite [20]. The ser ies of events which will finally lead to the formation of 1 Osteoclast: any of the large multinucleate cells closely associated with areas of bone resorption [22] 11 a biological bond between the implant and the bone can be presented as fol lows [20]: (1) biodegradation/bioresorpt ion of bioactive material, (2) formation of adhes ive proteins and col lagen fibril due to the differentiation of osteoblast 2 , (3) formation of C 0 3 - a p a t i t e microcrystals on the degrading C a - P implant material, and (4) s imul taneous mineral ization of the col lagen fibrils and incorporation of the new apatite crystals. 2.3 Calcium phosphate bone cements Calc ium phosphate bone cements are mixtures of var ious ca lc ium phosphate powders, such as C a H P 0 4 . 2 H 2 0 , Ca4 (P0 4 )20, C a H P 0 4 , C a 8 H 2 ( P 0 4 ) 6 . 5 H 2 0 , C a ( H 2 P 0 4 ) 2 . H 2 0 , or T C P and water or another liquid ( H 3 P 0 4 or N a 2 H P 0 4 ) . The mixture transforms into HAp during setting, forming a porous body even at 37°C [6]. The process of dissolution-precipitation of C a - P phases constitutes the setting p rocess and imparts mechan ica l strength to the final product. The main advantages of ca lc ium phosphate bone cements are their high biocompatibil ity, bioactivity and osteoconductivi ty, while their main d isadvantage is their relatively poor mechan ica l strength. The wet compress ive strength of ca lc ium phosphate cements range from 21 M P a to 51 M P a , which is affected by factors such as powder-to-l iquid ratio, and porosity of the cement [23]. Ca lc ium phosphate bone cements are currently used as bone fil lers, filling of the teeth root cana l and as drug delivery sys tems [6]. 2.4 Hydroxyapatite-based biomaterials 2.4.1 HAp powders There are severa l techniques to produce HAp powder, general ly c lassi f ied as wet methods and solid-state reactions. The wet methods can be divided into three groups: precipitation, hydrothermal technique, and hydrolysis of other ca lc ium phosphates [6]. Depending on the technique, materials with var ious morphology, stoichiometry, and level of crystallinity can be obtained. There are a lso other alternative techniques for preparation of H A p powders, such as sol-gel , flux method [24], electrocrystal l ization 2 Osteoblast: a bone-forming cell [22] 12 [25], spray pyrolysis [26], f reeze-drying [27], microwave irradiation [28], mechano-chemica l method [29], or emuls ion process ing [30]. 2.4.2 Dense HAp ceramics Dense H A p ceramics are obtained through shaping and pressure less sintering of HAp powders at moderately low temperatures (1000-1200°C). However, hot pressing (HP), hot isostatic press ing (HIP) or HIP-post-sintering makes it poss ib le to dec rease the densif icat ion temperature and grain s ize , and ach ieve higher densi t ies. Unfortunately, due to poor mechan ica l properties, appl icat ions of H A p dense ceramics have been limited to low-loaded implants. Due to their good compatibil i ty with human skin and t issues they are presently a lso used as dev ices for monitoring of b lood pressure and blood sugar, or optical observat ion of inner body t issues [6]. 2.4.3 Porous HAp ceramics Porous H A p materials have been appl ied as bone substitutes as they show a strong bonding to the bone. Bone will grow into the pores providing mechan ica l interlocking and thus increasing the mechan ica l strength of the H A p implant/bone assembly . A minimum pore s ize of 100pm is necessary to enable the ingrowth of bone and blood supply to the bone. This large s ize of the pores will in turn dec rease the mechanica l properties of the implants. The traditional way to produce porous H A p ceramics is to mix the powder with a pore-creating agent, or cast the powder into C a C 0 3 skeleton, which will be removed later by dissolut ion. There are s o m e low temperature process ing alternatives, including convers ion of a porous skeleton of C a C 0 3 into H A p under hydrothermal condit ions. Porous HAp materials have been used in medical appl icat ions in the form of b locks or granules for filling bone defects, drug delivery sys tems, and alveolar ridge augmentat ion [6]. 2.4.4 HAp-based ceramic composites HAp ceramic compos i tes attempt to address the problem of low mechan ica l reliability of pure HAp ceramics . Many reinforcements, including particles [31], whiskers [32], long fibers [33], partially stabi l ized z i rconia particles [34], and metal d ispersoids [35] 13 have been used in producing H A p compos i tes . Al though these may improve mechanica l reliability of HAp ceramics , at the s a m e time introduce new limitations such as lower biocompatibil i ty, stress shielding of the hard t issue, carc inogenic effect especia l ly when fibrous materials are used as reinforcements, and more complex and difficult p rocess ing. H A p / T C P ceramic compos i tes have been a lso deve loped, not to improve mechan ica l properties but to affect biological per formance, e.g. by controll ing the biodegradat ion rate of the composi te implant [6]. 2.4.5 HAP/bioactive glass composites Combinat ion of bioactive g lasses with H A p results in biomaterial with improved mechanica l properties without degradat ion of biocompatibil i ty or bioactivity [6]. In spite of high bioactivity, high biocompatibil ity and superior mechan ica l propert ies to HAp ceramics , HAp/bioact ive g lass compos i tes are used as coat ings or smal l , non- loaded implants, as their appl icat ion has not been success fu l in hard t issue replacement. 2.4.6 HAp coatings The concept of H A p coat ings on the metal implants comb ines the mechanica l advantages of metal and bioactive properties of HAp . The H A p coat ing prevents the encapsulat ion of bioinert metal implant with f ibrous t issues, improves the integration of bone into the implant and provides a stronger bond between the implant and bone. HAp coat ing also dec reases the re lease of metal ions from the implant into the physiological surrounding. Different methods have been proposed to apply H A p coat ing on metall ic implant, such as hot isostatic press ing [8,36], oxy-fuel combust ion spraying [37], magnetron sputtering [38], f lame spraying [37], ion beam deposit ion [9,39], chemica l deposit ion under hydrothermal condit ions [40], e lectrochemical deposit ion [11,41], pu lsed laser deposit ion [42], and sol-gel p rocess . In addition to all the methods ment ioned, p lasma-spraying [7] technique has become the most popular method to fabricate HAp coat ings. Th ickness of H A p coat ings can vary in a wide range, from fraction of micrometer for ion beam / sol gel techniques, to hundreds of micrometer for thermal spray techniques. 14 Increase in the th ickness of the coat ing will general ly dec rease the amount of metal ions introduced to body from the implant and extend the coat ing resorption time (the resorption rate can be as much as 15-30 urn per year). P r e s e n c e of porosity in the coat ing layer will improve bone ingrowth and thus bonding. Main problem regarding the HAp coat ings is delaminat ion of the coat ing due to fatigue or thermal coefficient mismatch at the meta l /HAp interface. To increase bonding strength between metal and coat ing, buffer layers such as bioactive g lass or C a 2 S i 0 4 have been cons idered [6]. A layer of dense uniform H A p can also be deposi ted on the metal sur face through a biomimetic p rocess of soak ing the implant in s imulated body fluid at 37°C, but the growth rate is slow, in the range of severa l micrometers per day [6]. 2.4.7 HAP/polymer composites HAp/po lymer compos i tes improve reliability and dec rease stif fness of the HAp biomaterials. HAp/polyethylene and HAp/poly lact ide compos i tes have been studied in depth [43]. HAp/polyethylene compos i tes have good mechan ica l properties such as high fracture toughness and a Young 's Modu lus c lose to that of the bone (1-8 G P a , depending on orientation). Unfortunately these compos i tes are not b iodegradable and also due to the p resence of the bioinert polymer their ability to bond to bone is low. HAp/poly lact ide compos i tes are both b iodegradable and bioactive but there are several reports on their toxic effect in body. The new approach is to precipitate H A p crystals on col lagen fibers to produce the composi te [6]. in spite of poor mechan ica l properties, HAp/co l lagen compos i tes exhibit superior osteoconduct ion (as compared to HAp or col lagen alone) and control led biodegradabil i ty, which makes them a good candidate for bone filling appl icat ions. 2.5 Sol-Gel HAp 2.5.1 The Sol-Gel process Sol-gel is a p rocess in which an inorganic structure evo lves from a solution (sol) following a chemica l multi-step pathway. A "so l " is a col loidal suspens ion of ~1-100nm large sol id particles in a liquid. T h e s e particles are under strong inf luence of short-1 5 range forces such as van der W a a l s and surface charges. Transit ion metal a lkoxides, M ( O R ) x are widely used as molecular sol-gel precursors to g lasses and ceramics . Due to the high electronegativity of the O R group, M is a lways in its highest oxidation state and therefore is suscept ib le to nucleophi l ic attack [12]. T h e alkoxide will therefore undergo a hydrolysis upon introducing to the solvent. The hydrolysis p rocess consis ts of a nucleophi l ic attack in which a proton is transferred from the attacking molecule to the alkoxide or hydroxo-l igand. H H \ / H 2 0 + M - O R 0 : M - ° R H O - M ^ O ^ -» M - O H + R O H (2.4) H H Hydrolysis The protonated spec ies will be then removed as either a lcohol (alcoxolation) or water (oxolation): R / M _ O + M - O R -> M — O: -» M - O R -> M - O - M <r O -> M - O - M + R O H I \ \ H H H Alcoxolat ion (2.5) H / M _ O + M - O H -» M — O: -> M - O H -> M - O - M <r O -» M - O - M + H 2 0 I \ \ H H H oxolation ( 2 . 6 ) W h e n the hydrolysis product is not yet saturated in coordinat ion condensat ion can occur through olation: H H / I M - O H + M ^ O M - O _ M + R O H (2.7) \ R 16 H H / I M - O H + M <r O -» M - O - M + H 2 0 (2.8) \ H The thermodynamics and kinetics of the hydrolysis, alcoxolat ion and oxolation are inf luenced by the nucleophi l ic strength of the attacking molecule, electrophilicity of the metal, charge stability of the leaving group, extent of coordinat ion unsaturation of the core metal and molecular complexity of the metal alkoxide which is a lso dependent on the alkoxide l igand [12]. Catalysts are usually used to inf luence the rate of the hydrolysis and condensat ion process . For example , acid catalysts can protonate the negatively charged alkoxide groups thus enhancing the reaction kinetics [12]: H / M- O R + H 3 0 + M + ^ : 0 ^ + H 2 0 (2.9) R The structure of the condensed products depends on the relative rate of the four reactions involved in the sol -gel : hydrolysis, oxolat ion, alcoxolat ion and olation. The extent of the contribution of each of these steps is a lso affected by parameters such as nature of the metal ion (M) and alkyl groups, molecular complexity, catalysts presence, concentrat ion, type of solvent and temperature [12]. A s a result of hydrolysis and condensat ion reactions inorganic polymers will form, and finally grow into clusters. The clusters then coll ide and link, leading to formation of a single giant cluster cal led the gel. At the time of gel formation many clusters are present al though not attached to the spann ing cluster. With t ime, the clusters would connect and as a result, the stiffness of the gel increases. The moment in which the last link is formed is cal led the gel point [12]. Fol lowing gelat ion, the steps of drying (solvent removal), calc inat ions (decomposit ion of hydrated oxide to oxide) and thermal consol idat ion (sintering) of the gel must be performed to produce a ceramic . At the first step of drying, the body would shrink due to removal of liquid from the gel pores through evaporat ion from the sur face. A s gel 17 pore s ize may be of nanometer range, surface tension strongly affects the drying process and may lead to gel cracking. Whi le drying, the gel body becomes stiff and shrinks as the liquid recedes into the pores. Eventual ly liquid would be trapped inside the pores and evaporat ion cont inues by vapour diffusion through the body to the outside sur face [12]. The calc inat ions and sintering of thus formed gels is performed in a single thermal treatment s tage. Densif icat ion process is driven by the tendency of materials to reduce the interfacial sur face energy between the sol id and vapour by reduction of sur face area. In c a s e of gels due to enormous surface area of the complex porous structure (100-1000m 2 /g) , sintering can occur at exceptional ly low temperatures [12], starting at several hundred degrees cent igrade. For example , the H A p gels in this work were consol idated at 400-500°C. The main advantages of sol gel over convent ional methods of ceramic process ing are as follows [44]: 1- High purity of starting materials 2- Homogenei ty and excel lent control of microstructure 3- Flexibility of shape in process ing. 4- Low temperature formation and crystall ization of the material, leading to lower cost and higher purity. The d isadvantages are [44]: 1 - High cost of raw materials 2- Shr inkage/deformat ion due to solvent removal 3- Multiple process ing step 2.5.2 Ceramic coatings by sol-gel The most technological ly important aspect of sol-gel process ing is its ability to produce thin fi lms prior to gelat ion. Compar ing to convent ional methods of coat ing such as chemica l vapour deposi t ion, evaporat ion, p lasma spraying or sputtering, sol-gel film formation requires less, and low cost equipment. Sol -ge l p rocess has a lso the unique 18 ability to precisely control the microstructure of the thin film i.e. pore vo lume, pore s ize and surface area . Sol -gel thin fi lms can be deposi ted on the sur face of the substrate using the following techniques [12]: Spin coat ing. The main advantage of this technique is its ability to produce a uniform thin layer. Sp in coat ing process can be divided into the s teps of deposi t ion, spin-up, spin-off and evaporat ion, while evaporat ion accompan ies other s tages, F ig . 2.3. During the deposit ion an excess amount of sol is deposi ted on the sur face, which will f low radially outwards at the start-up stage as a result of the centrifugal force. Droplets of excess liquids would leave the surface at the perimeters at the spin-off stage leaving behind a uniform, thin film [12]. 1 SDin-off Figure 2.3 S tages of the spin-coat ing process . From [12]. 2- Dip Coat ing . Dip coat ing can be divided into five s tages: immers ion, start-up, deposi t ion, dra inage, and evaporat ion, Fig 2.2. In case of volatile solvents evaporat ion accompan ies the start-up, deposit ion and drainage steps. Cont inuous dip coat ing is simpler as it el iminates the start-up and hides the drainage of the deposi ted film. A competit ion between six forces governs the film th ickness: (1) v iscous drag upward on 19 the liquid by the moving substrate, (2) force of gravity, (3) resultant force of surface tension in the concave ly curved men iscus , (4) interfacial force of the boundary layer liquid arriving at the deposit ion region, (5) surface tension gradient, and (6) the disjoining or conjoining forces. The main advantage of dip coat ing is its ability to coat complexly shaped substrates [12]. Figure 2.2 (A) s tages of dip coat ing process: (a-e) batch, (f) cont inuous. (B) Detail of the liquid f low patterns in a rea 3 of the cont inuous process . U is the withdrawal speed , S is the stagnation point, 8 is the boundary layer, and , h is the th ickness of the fluid film. From [12]. Other coat ing techniques appl ied to produce a thin film in sol-gel p rocess ing are: 20 3- Electrophoresis. In this method the conduct ive substrate is instal led as an anode or cathode and the coat ing forms due to the deposit ion of charged particles mobi l ized due to an external current [12]. 4- Settl ing: in which the particulate sols are deposi ted on the sur face due to the force of gravity accompan ied by the connect ive motion resulting from solvent evaporat ion [12]. 2.5.3 Sol-gel HAp materials Many studies have been conducted on the sol-gel synthesis of hydroxyapatite. Through these investigations different combinat ions of alkoxide precursors for calc ium and phosphorous have been studied, which have resulted in a range of products with different hydrolysis and aging time, sintering temperature and a variety of by-products accompany ing hydroxyapatite. J i l lavenkatesa et. a l . [45] have used calc ium acetate and triethyl phosphate as a source for ca lc ium and phosphorous, respectively. They have appl ied X-ray diffractometry (XRD) and infrared spect roscopy (IR) as their main character izat ion methods. In this study they have also investigated the effect of organic solvents such as methanol , ethanol and propanol on their so l gel sys tem. They have conc luded that hydroxyapatite can be obtained in this system at temperatures as low as 500°C, though it is usual ly accompan ied by calc ium carbonates, which will decompose into calc ium oxide at higher temperatures. They did not observe any significant effect induced by var ious organic solvents. Layrolle et. a l . [46] have produced amorphous ca lc ium phosphate ( A C P ) through a so l -gel p rocess , using ca lc ium diethoxide and phosphor ic ac id as precursors and ethanol as solvent. They have shown that the A C P powder crystal l izes to a carbonated hydroxyapatite and a trace of pMricalcium phosphate before convert ing to pure hydroxyapatite at 900°C. They have also obtained a microporous structure through the decomposi t ion of carbonated hydroxyapatite with the pore s ize of 0.2| im. Y . M a s u d a et. a l . [47], have used, calc ium diethoxide and triethyl phosphite as starting metal alkoxide. They have investigated the effect of pH on the final phase formation of hydroxyapatite in their sol-gel sys tem. They have been able to obtain pure, plate-like 21 hydroxyapatite using solut ions with pH range of 6-8, water content of 60%, and calcining temperature above 600°C. It was shown that alkal ine solut ions with pH>9 result in precipitates, which will turn into hydroxyapatite, ca lc ium oxide and tri-calcium phosphate when ca lc ined at 900°C. The acidic solid-free solution produces a single-phase hydroxyapatite when solidif ied and heated to 900°C. Takahash i et. a l . [48], have synthes ized stoichiometric hydroxyapatite through a sol-gel route of ca lc ium nitrate and phosphonoacet ic acid in aqueous citric ac id solution. A brown powder was obtained by heating the solution , which turned into hydroxyapatite after heating at 980°C for 7 hours. X R D experiment did not show any carbonate ions in the final hydroxyapatite. G . Kordas et. a l . [49], have investigated the feasibility of sol-gel p rocess for the synthesis of hydroxyapatite. Ca lc ium acetate and P O ( O C 2 H 5 ) 3 were used as precursors while methyl, ethyl and propyle alcohol were used as solvents. They investigated the evolution of the structure using X-ray diffraction, IR and F T - E P R . It was conc luded that sol-gel method yields satisfactory results for H A p preparation. They have detected smal l amounts of C a O along with hydroxyapatite after heat treatment at 930°C, which can be removed upon washing with acet ic ac id and water. Us ing F T - E P R the authors have also showed that the network develops in one direction during hydrolysis and complexat ion. They have also ment ioned that one-ethyl group remains unreacted after drying at 75°C. J . L ivage et. a l . [50], have done a fundamental study on the sol-gel synthesis of phosphates, showing that polyphosphates cannot be synthes ized under ambient condit ions from P O ( O H ) 3 or P O ( O R ) 3 . The hydrolysis of P O ( O H ) 3 is difficult while P O ( O R ) 3 leads to the formation of [Hx(P0 4) ] ( 3" x )" spec ies , which have high reactivity toward complex ing and therefore do not condense . This effect will result in precipitation rather than gelat ion. They have also shown that dissolv ing P2O5 in a lcohols results in P O ( O H ) 3 . x ( O R ) x spec ies , which have reactivities intermediate between P O ( O H ) 3 and P O ( O R ) 3 . Accord ing to their research Alkyl phosphates, P ( O R ) 3 , are a lso suitable precursors in the sol-gel p rocess of phosphates. 22 C. S . Cha i et. a l . [51] have shown that a critical aging of at least 24 hours is necessary to obtain pure hydroxyapatite. Shorter aging periods result in impure phases such as carbonates, which further decomposed to calc ium oxide at 600°C. It is suggested that the solution should be aged till no abrupt weight loss is observed between 680°C and 750°C upon heating. It is conc luded that the aging process is dependent on the calc ium and phosphorous precursors (calcium diethoxide and triethyl phosphite in this research). Ag ing time increases with increase in the concentrat ion of these reactants while it reduces with the reduction of solvent used in the sol-gel process. A compromise should be obtained with aging time and the amount of solvent used as it might adversely affect the chemistry and quality of the sol-gel H A p coat ing. J i lavenkatesa et. a l . [52] have conducted an electron microscopy study on the different s tages of hydroxyapatite formation through a sol-gel p rocess . Ca l c i um acetate and triethyl phosphate have been used as calc ium and phosphorous precursors. They have conc luded that a nucleation-growth process controls formation of hydroxyapatite in this sys tem. Increasing the temperature provides enough activation energy for crystal l ization. The fine structured matrix is bel ieved to be of a ca lc ium apatite nature as it is c o n s u m e d by the crystall ization and growth process . Hydroxyapati te crystals are observes to be equiaxial at temperatures below 1000°C. Higher temperatures induce hexagona l -shaped crystals, which is consistent with the hexagona l unit cell in which hydroxyapatite crystal l izes. W e n g et. al . [53] have used calc ium glycolate, phosphor ic ac id , and P ( O H ) x ( O E t ) 3 . x as precursors while acet ic acid was used as reagent to modify the ca lc ium glycolate and also change the acidity of the mixture. They have been able to obtain a transparent gel depending on the amount of the acetic acid and the extent of stirring. They have attributed this behaviour to the large molecular s ize of the ethylene glycol solvent, which makes the reactions dependent on diffusion. They have a lso been able to synthesize ca lc ium phosphates with different C a / P ratios by changing the ratio of acetic a c i d / C a . Deptula et. a l . [54] have produced spher ical powders of hydroxyapatite with diameters of 100 urn using water extraction variant of the sol-gel p rocess . The solution is prepared with ca lc ium acetate and phosphor ic acid as precursors. The solution is then 23 emulsi f ied in dehydrated 2-ethyl-1-hexanol. Drops of the emuls ion were solidif ied by extraction of water with this solvent. They have observed that the formation of hydroxyapatite starts above 400°C while formation of carbonated hydroxyapatite happens at 580°C. W e n g et. a l . [55] have used an ethylene glycol solution of C a ( O A C ) 2 . x H 2 0 and butanol solution of P 2 0 5 as precursors for sol-gel hydroxyapatite. Acet ic ac id and ammonium nitrate were used as stabil izer and oxidizer, respectively. At 500°C poorly crystal l ized hydroxyapatite has been obtained through this p rocess . They have been ab le to obtain well-crystal l ized H A p and a smal l amount of (3-tricalcium phosphate when N H 4 N 0 3 was used as a stabil izer. A new water based sol-gel process has recently been deve loped by D e a n - M o et. al . [56] at U B C e r a m . In this process calc ium nitrate tetrahydrate and triethyl phosphite are used as starting precursors and water and ethanol as diluting med ia for the calc ium precursor. They have observed the formation of H A p crystals at temperatures as low as 350°C. A more pure HAp phase has been deve loped in powders from ethanol based solut ions while other calc ium compounds such as tr icalcium phosphate s e e m e d to accompany H A p in merely aqueous-der ived powders. In a fol low up study, the effect of hydrolysis on this sys tem has been investigated by the s a m e group [57]. They have introduced a chemica l pathway for hydrolysis reactions in this sys tem. Accord ing to their speculat ions triethyl phosphite undergoes hydrolysis in the p resence of water to form diethyl phosphorous esters: P ( O C 2 H 5 ) 3 + H 2 0 - * H P O ( O C 2 H 5 ) 2 + C 2 H 5 O H (2.10) Diethyl phosphorous ester may undergo further hydrolysis and as a result more O R groups will be replaced by O H group from aqueous environment: H P O ( O C 2 H 5 ) 2 + H 2 0 H P O ( O C 2 H 5 ) 2 - x ( O H ) x + x C 2 H 5 O H (2.11) The hydrolysed phosphates will further interact with C a through a condensat ion polymerization reaction: 24 2 H P O ( O C 2 H 5) 2.x ( O H ) x + C a + + ^ [ ( O C 2 H 5 ) 2 - x O ( H ) P - 0 ] 2 C a + 2 H + (2.12) They have attributed the formation of T C P to condensat ion of hydrolysed phosphite molecules, which will further interact with ca lc ium to form ca lc ium phosphate derivatives [57]. The U B C e r a m group has also observed that the addit ion of acid catalyst will accelerate the rate of hydrolysis and condensat ion in this sys tem. The work presented in this thesis originated from this p rocess deve loped at U B C e r a m [56,57]. 2.5.4 Sol-gel H A p coatings H A p coat ings on metall ic implants combine good mechan ica l character ist ics of metals with bioactive properties of hydroxyapatite. Many coat ing techniques have been deve loped to deposit a layer of calc ium phosphate on metall ic substrate. Al l these coat ing methods have the problem of a week mechan ica l bonding between the metal oxide layer and the ca lc ium phosphate coat ing. The sol gel p rocess offers a relatively low coat ing temperature, results in stoichiometric and homogeneous coat ings, and can yield both amorphous and crystal l ine fi lms (this is important in order to control the film resorption process) . The low temperature process ing is an advantage when the heat treatment can adversely affect the properties of the substrate, such as phase transformation occurr ing in titanium at 883°C. A l so , oxidation of metall ic substrates at e levated process temperatures could be a problem. Spinning or dipping the sample in the solution results in thin sol-gel coat ings. The fi lms obtained with this method are usual ly thin (<1 urn) and therefore can easi ly undergo heat treatment process without cracking or substant ial shr inkage, characterist ic of the sol-gel method for bulk materials. It is a lso poss ib le to uniformly coat both s ides of planar and axially symmetr ic substrates such as p ipes, rods, fibres and wires, not easi ly handled by more convent ional coat ing methods. Major process ing s tages involve producing a solution from suitable a lkoxides or salts, hydrolysis and ageing of the solut ion, deposit ion of the coat ing, and heat treatment. 25 The heating p rocess involves the removal of the organics and provides enough thermal energy for chemica l and physical reactions necessary to form and density hydroxyapatite. Many studies have been done on different starting materials, process ing s tages and characterizat ion of hydroxyapatite thin fi lms on different substrates, mainly titanium and stainless steel al loys. Piveteau et. a l . [58] have deposi ted a layer of mixed C a - P and titanium oxide (T i0 2 ) on titanium substrate to improve the bonding between the substrate and the coat ing. A so l of titanium dioxide was mixed with a solution of ca lc ium nitrate and phosphorous esters. The composi te was then deposi ted on the sur face of the titanium substrate, which has undergone various pre-heat treatments and has been blasted with aluminium balls to improve mechan ica l interlocking between the substrate and the coat ing. Sp in coat ing was used to apply the solution on the substrates, which were then fired to 850°C. They have observed that the viscosity of the precursors and also surface topography of the substrate directly affected the sur face topography of the coated samp les . Accord ing to their results more v iscous starting precursors produce rougher sur faces. Haddow et. a l . [59] have deposi ted a biphasic calc ium phosphate (HAp and T C P ) on quartz g lass substrate through a sol-gel process. The coat ing has been deposi ted using dipping of the substrate in the solution and then firing the substrate to 1000°C. They have conc luded that the coat ing th ickness after firing at 900°C were significantly less that 1 um for solution containing 1 M ethanol, with the thinnest coat ing being 0.15 urn thick. They have also observed that the time and a tmosphere of heating affects the phases evolved in the coat ing, e.g. water molecule in the heat ing a tmosphere will promote the formation of hydroxyapatite while in a dry a tmosphere T C P and tetracalcium phosphate (TTCP) are more stable. It was reasoned that a b iphasic calc ium phosphate coat ing on a titanium substrate will increase the bioactivity of the coat ing as a more soluble C a - P phase (TCP) along with an insoluble phase (HAp) will promote a rapid bone response. W e n g et. a l . [60] have prepared hydroxyapatite coat ing on an a lumina substrate using C a ( N 0 3 ) 2 - 4 H 2 0 and P2Os as starting precursors and ethanol as the diluting med ia . The coat ing showed good adhes ion strength of about 10 M P a when fired to 500°C. They 26 have observed that increasing the temperature to 750°C will produce (3-tricalcium phosphate as a by-product while it introduces more porosity to the morphology of the coating. Vary ing the process ing condit ions such as heat treatment time and temperature produced hydroxyapatite with various degrees of crystallinity and porosity. A nanocrystal l ine layer of hydroxyapatite has been deposi ted through a sol-gel process on various substrates by C h a i et. al . [61]. The solution was prepared using calc ium diethoxide and triethylphosphite diluted in ethanol and then aged for seven days. The sol was then spin coated on var ious substrates such as Vycor g lass , polycrystall ine a lumina, partially stabi l ized zirconia and single crystal M g O . Multiple layers of the coating were deposi ted with the prefiring temperature of 500°C and then 1000°C as the final treatment. The coating thus prepared was crystall ine with the crystal s ize of 200-800 nm. They have a lso observed good coat ing/substrate adhes ion . A s it cou ld be expected the relatively high firing temperature has produced ca lc ium oxide as a by-product. A hydroxyapatite thin layer has been coated on the surface of a single crystal S i (001) substrate through a sol-gel process by Hwang et. a l . [62]. They have used phosphoric acid and calc ium nitrate as phosphorous and calc ium precursors, respectively. The solution has been spin coated on the surface, prefired at 300°C and then heat treated at 500°C and 700°C. The authors observed the formation of hydroxyapati te structure at 500°C while the formation of [3-tricalcium phosphate started at 700°C. They have attributed the formation of (3-tricalcium phosphate to the decomposi t ion of carbonated hydroxyapatite during the final heat treatment at 700°C. Gross et. a l . [63] have investigated the changes occurr ing in the so l during the aging period and also the behaviour of the xeroge ls 3 upon heat treatment using 3 1 P nuclear magnet ic resonance spect roscopy and thermal gravimetric analys is . The solution was prepared with ca lc ium diethoxide and triethylphosphite as starting precursors. The solution was then aged for 24 hours and spin coated on the sur face of titanium coupons and finally fired to 800°C. Accord ing to their exper imental results, hydroxyapatite crystall ization starts at 550°C and further heat treatment will remove the organic res idues. The authors have also conc luded that a min imum of 24 hours of 3 Xerogel: dried gel, when drying has taken place in ambient conditions [12] 27 aging period is necessary to obtain pure hydroxyapatite, otherwise ca lc ium oxide will be present as an impurity phase. Accord ing to their f indings during the aging period the two alkoxide will react to form calc ium and phosphorous diethoxide ( O E t - C a - O -P(OEt ) 2 ) , which will then condense to form [ C a - 0 - P 0 3 ] n through a rapid reversible process. Hwang et. a l . [64] have coated porous a lumina substrate with hydroxyapatite through sol-gel p rocess . Their p rocess uses ca lc ium nitrate tetrahydrate and phosphor ic acid as the starting precursors. The coated substrate was subjected to a heat treatment of 700°C. They have observed a shift in the peak posit ions of hydroxyapati te, which was attributed to the incorporation of residual carbon into the structure of hydroxyapatite. Lopatin et. a l . [65] have looked at the effect of drying and firing s tages in a hydroxyapatite sol-gel using N-butyl acid phosphate and ca lc ium nitrate tetrahydrate as precursors. Increasing the drying temperature supp ressed the formation of hydroxyapatite. The authors have attributed this effect to the difficult rearrangements of M - O - M links formed during the drying stage. Accord ing to their hypothesis increasing the drying temperature will result in higher number of M - O - M bonds . It was a lso observed that hydroxyapatite crystal l izes from a medium, consist ing of hydroxyapatite nuclei and an amorphous matrix of organic compounds . Longer soak ing time during drying will promote nucleation of hydroxyapatite crystals and therefore will lead to lower H A p formation temperatures. Many of the results from the studies summar ized above are similar regardless of the starting materials and the sol-gel p rocess features. A critical aging period (24 hours according to most of the authors) s e e m s to be necessary in order to obtain monophas ic hydroxyapatite. In the case of insufficient aging, ca lc ium oxide would be the impurity phase accompany ing H A p . Formation of ca lc ium oxide has been attributed to the unreacted ca lc ium precursor, which will further react with water to form calc ium hydroxide. The calc ium hydroxide thus formed, will react with C 0 2 (a by-product of decomposi t ion of organic compounds) to form calc ium carbonate, which will finally decompose to ca lc ium oxide and C 0 2 . pMricalcium phosphate is a lso observed to be present a long with hydroxyapatite at higher p rocess ing temperatures (>600°C). This effect has been attributed to decomposi t ion of the carbonated hydroxyapatite. 28 Carbonated hydroxyapatite forms due to incorporation of carbon into the structure of hydroxyapatite. Th is effect has been verif ied by a slight shift in the position of hydroxyapatite characterist ic X R D peaks [64]. Many authors have observed that hydroxyapatite crystal l izes from an amorphous C a - P matrix so the crystal growth of the final H A p would depend on the decomposi t ion rate of this amorphous phase. Due to the complexity of the chemistry involved in sol-gel p rocess ing of H A p no c lear pathways for different s tages of this process has been suggested as yet. 2.6 Stents A stent is "a short narrow metal or plastic tube that is inserted into the lumen of an anatomical vesse l (as an artery or a bile duct) especia l ly to keep a formerly b locked passageway open" [22]. A n Engl ish dentist, Char les T. Stent, who lived from 1807 to 1885, first used a stent. The first dental stent was descr ibed as "a new dental impression material of wax and gutta percha, a latex relative consist ing of the milky sap from the Pa laqu ium gutta tree in south-eastern A s i a " [66]. During the First Wor ld W a r J . F. E s s e r of Hol land used the dental stent as a urologic tool, which marked the first appl icat ion of the dental impression in medical surgery [66]. The new technologies of the 2 0 t h century such as d iscovery of antibiotics and biocompatible materials have provided the platform for the appl icat ion of stents in endovascu lar operat ions. The new stents are wire mesh tubes common ly made of a medical grade stain less steel 316L, L indicating the low carbon content (0.03%). This alloy is c o m p o s e d of iron (60% to 65%), chromium (17% to 18%) and nickel (12% to 14%). Sta in less steel provides good mechan ica l , chemica l and physical properties but its biocompatibil i ty remains an issue. Nickel (~55%)-titanium (-45%) (Nitinol) stents may also be used as they offer some better early biocompatibil i ty al though concerns regarding nickel leakage have limited their appl icat ion. Tanta lum is another option for materials used as stent material. Nitinol stents offer better radio-opacity, biocompatibil i ty and mechan ica l properties although its cl inical superiority to stainless steel stents has not been yet proved [67]. Stents are widely used to keep the blood vesse l open fol lowing a bal loon angioplasty. During this procedure the stent is p laced on a bal loon and manoeuvred into b locked 2 9 area. After the bal loon is inflated, stent will lock in its p lace and will hold the artery open. Upon introducing to the blood vesse l , stents permanent ly stay in the vesse l and cannot be removed. Figure 2.4 shows the schemat ic picture of a stent p laced inside the artery. Res tenos i s 4 , th rombos is 5 , inflammation and neo in t ima 6 formation are the early biocompatibil ity problems regarding stent implantation in the coronary arteries [69]. Al though the mechan ism of restenosis is not yet clear, it is bel ieved to be the result of hyperproliferation of vascu lar smooth musc le cel ls which will finally lead to scar formation inside the blood vesse l and therefore blocking the blood flow [70]. Late problems with the stents can be divided into two broad categor ies: mechan ica l failure due to material fatigue under the stress of cardiac contractions and chemica l corrosion leading to the leakage of toxic subs tances , degradat ion products or contaminants [67]. The materials used as stents must have certain physica l , chemica l and mechanica l properties. A n expandab le metal stents should have enough plasticity to retain the required s ize after expans ion. O n the other hand sel f -expanding stents should posses enough elasticity so that they can be compressed and then expanded while 4 Restenosis: the re-occlusion of the vessels [68] 5 Thrombosis: the formation or presence of a blood clot within a blood vessel [22] 6 Intima: the innermost coat of an organ (as a blood vessel) consisting usually of an endothelial layer backed by connective tissue and elastic tissue [22] Figure 2.4 Stent P laced inside the artery. 30 maintaining sufficient radial loop strength. Corros ion resistance is the most important chemica l characterist ic regarding stents. The oxide layer formed on the sur face of the metallic stent will retard corrosion and thus provide the metal with highly pass ive surface [67]. Mult idiscipl inary studies are now being conducted on stent research and as a result, new des igns and different materials and coat ings are being studied in order to enhance the quality of these prostheses. Exper imental data suggests that the stent design and surface propert ies will strongly inf luence its performance [67]. In order to evaluate the effect of the coat ing on the performance of the stent severa l different coat ing materials have been studied. The results of these studies can be summar ized as fol lows: 1- Nonbiodegradable synthetic polymers. Many polymer materials such as polyurethane, s i loxane (silicon), polyethylene terephtalate, and cross- l inked phosphorychol ine have been coated on the sur face of different stents. The t issue response, such as inf lammation, neoint ima formation and thrombosis were then studied after implantation the stents in the animal model [71,72,73,74]. The experimental results of these exper iments do not provide any final conc lus ion whether any polymer coat ing may improve the stents' biocompatibil i ty, as the results vary from no apparent dif ference between the coated sample with the bare stent [74] to intense inflammatory responses in the presence of the coat ing [71] or better biocompatibil i ty due to the coated surface [73]. However, polymers such as polyethylene and phosphorychol ine s e e m to be suitable candidates for drug delivery sys tems on stent surface, in order to pharmacological ly address the negative response of the t issue to stent p resence [67]. 2- B iodegradab le synthetic polymers. Occ lus ion and a wide inf lammatory response were observed upon implantation of polyglycoli/polylactic ac id copolymer, polycaprolactone, polyhydroxy-butyrate/valerate copolymer ( P H B V ) , polyorthoester, polyethyleneoxide/polybuthlene terephtalate copolymer ( P E O / P B T P ) , and low molecular weight poly-l-lactic acid ( L M W P L L A ) coated stents [71, 75]. High molecular weight poly-l-lactic ac id ( H M W P L L A ) coat ings show no ev idence of chronic inf lammation. A m o n g biodegradable polymer coat ings P L L A is a lso regarded as a good matrix for temporary drug re lease. 31 3- Natural coated stents. In this c a s e the surface of the stent (either bare or pre coated) is e n c a s e d with fibrin [67]. Theoret ical ly the natural coat ing should minimize inflammation and other negative t issue responses . A significant reduction in platelet deposit ion has been indeed observed by Baker et. a l . [76] after implantation of a titanium sel f -expandable stent coated with fibrin and impregnated with Hepar in. Stefanid et. a l . [77] have covered the convent ional stent with autologous vein or artificial graft which has produced significantly enhanced results but the process is yet too complex to be widely adopted. 4- Heparin coated stents. Hepar in 7 coat ings have shown to be effective in the reduction of thrombosis but their effect on the neoint ima formation needs to be establ ished [67]. 5- Drug-eluting stents. Much interest has been focused on loading a drug onto a stent to limit the early thrombogenicity and late neoint ima formation. Drugs maybe re leased by diffusion mechan isms or during polymer break down [67]. 6- Polymer ic stents. Polymer ic stents have been deve loped using Dacron [78], polyglycolic acid [79], and biodegradable stent in col lagen [79]. T h e s e stents are appl ied through angioplasty, which will then degrade inside the vein. The efforts to develop b iodegradable stents have recently s lowed down due to the complexity of the process and also acceptable performance of the metall ic stents, while the idea of a biodegradable stent loaded with drug is still quite appeal ing [67]. 7- Radioact ive stents. Certa in amount of radioactive doses del ivered at the site of stent implantation may decrease the negative responses of the t issue. Radioactivity can be produced by either particle bombardment or ion implantation of the metall ic surface, which on the other hand may affect the corrosion propert ies of the sur face. Delivering radioactive isotopes through an eluting system can be another alternative [67]. To enhance the effect iveness of stents and broaden their appl icat ions to more complex lesions, better biocompatibil i ty needs to be ach ieved. The new generat ion of stents should be less thrombogenic coronary endopros theses and more tolerable by the body 7 Heparin1 a mucopolysaccharide sulfuric acid ester that is found specially in liver, that prolongs the clotting time of blood [22] environment. This provides the rationale for studying the possibil i ty of H A p coat ings stents, as presented in this thesis. 33 CHAPTER 3 SCOPE AND OBJECTIVES 3.1 Scope of investigation Ceramics have been used in medical appl icat ions due to their biocompatibil i ty while ceramics like hydroxyapatite and bioglass® have bioactive character ist ics in addition to their compatibil i ty with body environment. Al though low mechan ica l strength of bulk b ioceramics has limited their application in areas where mechan ica l durability is critical, they are still used in forms of compos i tes or as coat ings on a wide range of prostheses. B ioceramic coat ing provides enough biocompatibil i ty and bioactivity for the implant which otherwise would be biologically inert or even elicit negative response in the surrounding t issues. Sol -gel p rocess al lows fabrication of thin film ceramic coat ings on var ious substrates. The main advantages of this s imple process ing method are the homogenei ty of the final coat ing, control of its microstructure, ability to produce thin fi lms (<1uxn), and low temperature of heat treatment. The low temperature factor is of great importance in biomedical appl icat ions as the process ing temperature of the coat ing might adversely affect the mechan ica l properties of the substrate. Thus , sol-gel synthesis of hydroxyapatite has been studied in the present thesis. The resulting coat ings should have good adhes ion to the substrate such that they are not d a m a g e d (or removed) during surgical implantation. In the c a s e of stents, the appl icat ion cons idered in this thesis, the coat ings should resist delaminat ion upon expans ion of the stent in the arteries. Crystal l izat ion of the coat ing a lso needs to be control led as it will further affect the rate of dissolution of the thin film into the blood st ream while affecting the plasticity of the coat ing. In order to produce a high quality HAp coat ing through a sol-gel sys tem comprehens ive understanding of the particular sys tem as well as sol-gel technique in general s e e m s necessary . Understanding the effect of controll ing parameters in the sol-gel p rocess , such as solvents and heat treatment temperatures, provides valuable knowledge to fulfill speci f ic qualif ications necessary for an individual appl icat ion. In this respect, the scope of this work is to study the effect of organic solvents on the phase 34 evolution of hydroxyapatite in a novel , relatively low temperature sol-gel p rocess . Very few previous studies have addressed the effect of the solvents on H A p sol-gel sys tems. The knowledge gained through this fundamental study is then appl ied to develop an opt imum coating process for titanium substrates (flat) as well as stainless steel coronary stents (-0.1 mm thin wires). The latter is the first attempt to apply hydroxyapatite as a biocompatible coat ing on such implants. Severa l character izat ion techniques such as S E M , X R D , and IR spect roscopy are used in order to evaluate the obtained coat ings. 3.2 Objectives The general objective is to investigate the effects of organic solvents (methanol, ethanol, propanol), on the phase evolution of hydroxyapatite in triethyl phosphite-calc ium nitrate sol-gel sys tem. The study objectives can be broken down as fol lows: 1- Character ize the phases present in the heat-treated H A p powders at various temperatures to monitor phase evolutions, and crystal l ization. 2- Investigate the effect of organic solvents on the morphology of the sol-gel H A p powder. 3- Investigate properties of H A p coat ings from sols obtained using var ious solvents. 4- Apply and character ize the HAp sol-gel coat ings on different model metall ic substrates to opt imized process parameters. 5- Apply and character ize the opt imum H A p sol-gel coat ings on commerc ia l coronary stents. These investigations were carr ied out using the following techniques: 1- P h a s e identification of the dry and sintered powders was carr ied out by qualitative X-ray diffraction (XRD) . 2- Thermal gravimetric analysis (TGA) and differential thermal analys is (DTA) were used to investigate thermal behaviour of the dry gels upon heat ing. 3- Scann ing electron microscopy (SEM) was used to observe powder morphology and also coated sur faces. Electron dispersive spectrometry is used in order to study the elemental composi t ion of the surface of coated and uncoated substrates. 35 4- Sol id-state infrared (IR) analys is is carr ied out to character ize the coat ings, obtained from different solvent sys tems, on the model titanium substrates. 36 CHAPTER 4 EXPERIMENTAL METHODOLOGY The sol-gel precursors consist of triethyl phosphite and ca lc ium nitrate tetrahydrate, d issolved in organic solvent, i.e. methanol , ethanol, or propanol . Powde r process ing, character izat ion, and coat ing process ing and character izat ion follow the sol preparation. Therefore, this experimental and methodology chapter is divided into three sect ions: (1) general procedures, (2) sol-gel powder preparation and character izat ion, and (3) sol-gel coating preparation and character izat ion. Samp le preparation procedures for character izat ion exper iments and instrumental information are a lso included. The second sect ion of this chapter descr ibes the preparation and heat treatment of sol-gel hydroxyapatite powder. Character izat ion p rocesses of the sol-gel powder are a lso provided. In this sect ion phase evolution upon heat treatment, and microstructure of the sol-gel powders are investigated. The third sect ion details the coat ing and characterizat ion procedures for the H A p coat ings on various substrates. Coat ing procedures usually follow the sequence of (a) substrate surface treatment (preparation); (b) coat ing, (c) heat treatment, and (d) character izat ion. The samp les were character ized using X-ray diffractometry, FTIR spect roscopy, scann ing electron microscopy, and energy dispersive spectrograph. 4.1 General procedures 4.1.1 Materials Unless otherwise stated, all materials were used as received. Triethyl phosphite (P(OEt) 3 ) and ca lc ium nitrate tetrahydrate (Ca(N0 3 )2' 4 H 2 0 ) were obtained from Fisher Scientif ic and Aldr ich, respectively. Sod ium carbonate and Borax were used as neutralizing agents for acid treated stain less steel wires and were obtained from Fisher scientif ic and Greenbarn Potters Supply , respectively. Ti tanium substrates had a th ickness of 0 .127mm and were manufactured by A l fa Aesar . Med ica l grade stainless steel (316L) wires and coronary stents were provided by MIVI technologies of Vancouver , B C , and were used as coat ing substrates. 37 4.1.2 Analysis Thermal gravimetric analysis was conducted on powder samp les obtained from solution prepared from different organic solvents using a Perk in Elmer, 7 ser ies, Thermal Ana lys is Sys tem, in order to monitor the weight loss of the powder upon heating. The tests were done at a heating rate of 10°C/min, from ambient to 1200°C. These exper iments were conducted at the Institute of Mater ia ls Sc ience and Manufactur ing, Ch inese Culture University, Y a n g Ming S h a n , Ta ipe i , Ta iwan, our col laborators on this project. P h a s e identification of ca lc ined powder and thin film coat ings was done using X-ray diffractometer R igaku Rotaf lex 200, Tokyo, J a p a n , using C u K a (1.54 nm) with 29 of 25-40° with the step s ize of 5°/min. Powder samp les were packed in an aluminium sample holder and then exposed to radiation. To investigate thin coat ings, and especial ly coat ings on wires and stents, powders were obtained from the coating solut ion, fol lowing the s a m e process as for the coat ings. The powder was then ground with a smal l amount of ethanol and appl ied as a thin layer on a g lass substrate. The thin layer on the g lass substrate was then dried and heat-treated a long with the actual coated substrates to simulate the coat ing heat treatment condit ions. After heat treatment, the g lass substrates were directly subjected to further X R D tests. Scann ing electron microscopy was conducted for microstructural examinat ion, using a Hitachi S E M . Due to the smal l s ize of the coated samp les and dry gels , they could be easi ly p laced on the stub and be examined. Except for the bare metall ic substrates, all the samp les were gold coated prior to S E M examinat ion. Energy dispersive X-ray analysis was carr ied out to investigate the elemental composi t ion of the sur face of coated substrates. Infrared spect roscopy was carr ied out using a Perkin E lmer sys tem 2000 FTIR. So l id-state analys is was conducted on coated samp les using K B r as background. 4.2 Powder preparation and characterization Figure 4.1 schemat ical ly shows the so-gel H A p synthesis procedure. Phosphorous sol was prepared by hydrolysing smal l amount of triethyl phosphite (5.08 g), with distilled water (the molar ratio of phosphite/water was kept at 3) under v igorous stirring, with 38 the use of nitric acid as a catalyst. A 2.1 N acid solution (10.2 g) was first added to the phosphorous alkoxide, aged for 5 minutes, and fol lowed by the addit ion of a second solution. The second solution was prepared by dissolv ing ca lc ium nitrate (11.80 g) into either one of three organic solvents (25 ml of methanol , ethanol or 1-propanol). The mixture was then stirred for 30 minutes fol lowed by an aging period of 24 hours. The aged sol was then oven dried for 24 hours at 80°C until a white, dried gel was obtained. Th is procedure was first proposed by D e a n - M o et. a l . [56]. Triethyl phosphite + Acid water Calcium Nitrate + Alcohol — — i M i x i n g Aging Figure 4.1 F low chart of so-gel H A p synthesis procedure. The dry gel was ground and heat-treated at different temperatures (375, 425 , 445 , 465, 500 and 545°C). The heat treatment temperatures were chosen based on D T A results, to monitor phase evolution in the gel upon heat treatment. Powde r samp les were p laced in the furnace in a lumina crucibles. The temperature was then increased to the desired level, at which the sample was instantly removed from the furnace. X R D and S E M tests fol lowed heat treatment procedure to investigate the phase composi t ion and morphology of the ca lc ined powder. 39 4.3 Coating preparation and characterization 4.3.1 Titanium substrates In order to obtain a good adhes ion between the coat ing and substrate, and also to increase the wetting of the substrate by the solut ion, the sur face of the substrate should be roughened. This can be accompl ished either through a chemica l treatment, in which the substrate is exposed to a corroding environment, e.g. strong ac id , or mechan ica l roughening such as S i C ball blasting or sand blast ing. Mechan ica l methods of sur face roughening tend to be more destructive and are not suitable for substrates with complex shapes , concave sur faces, or fine structures such as wires. In this work titanium substrates were chemical ly treated (i.e. partially corroded) in phosphor ic acid (85%) at 50-60°C for 30 minutes. A number of t itanium substrates were also subjected to sandblast ing prior to coat ing. Pre-treated titanium substrates were spin coated with the sol-gel hydroxyapatite precursor solut ion, prepared using different organic solvents as diluting med ia as descr ibed above. The spin coat ing was carr ied out using a commerc ia l spin coater (Headway research Inc. E C 1 0 1 ) with the speed of 3000 rpm for 20 seconds . IR spect roscopy, S E M , and E D S tests were done on the coated sur faces to investigate the chemica l composi t ion and surface morphology of the coat ing layer. Coa ted samp les were then p laced in simulated body fluid (SBF) for a period of seven days after which they were removed and rinsed with distil led water and air dried for three days. The S B F composi t ion is presented in appendix I. Format ion of apatite crystals on the surface of the coat ing, which is an indication of the bioactivity of the coat ing [80], was investigated with S E M . 4.3.2 Stainless steel substrates 4.3.2.a Wires It is bel ieved that a rough surface improves the adhes ion of ceramic coat ings to metallic substrates, e.g. sta in less steel substrate, as the adhes ion is mainly governed by mechan ica l interlocking [81]. In order to find the best chemica l treatment for the wire's sur face and produce a sufficient degree of roughness on the sur face, the effect of phosphor ic acid (85%), nitric ac id (70%), and hydrochloric ac id (37%) on the 40 stainless steel wires was tested. The acid solut ions were diluted in water by of 50 v o l % (equivalent to 8 N for the phosphor ic ac id , 20 N for the nitric ac id , and 86 N for the hydrochloric acid) . The stain less steel wires were exposed to the acidic solut ions for 30 minutes at 70°C. The wires were then neutral ized in an aqueous solution of 1.3 g/l N a 2 C 0 3 and 0.4 g/l Borax ( N a 2 B 4 0 7 ) at 70°C for 4 minutes fol lowing a process proposed by shieu et. al . [81]. To investigate the effect of exposure time on the extent of etching, s ta in less steel wires were etched in 2 .4N HCI solution for 10, 20, 30, 40, 50, and 60 minutes. The samp les were then neutral ized and studied using S E M . The surface treated wires were shaped (bent) to resemble the shape of an actual stent and dip coated using an in-house made dip coater, with dipping rate of 14 cm/min. The coating solut ions were the hydroxyapatite so ls prepared using water, methanol , ethanol and propanol as diluting med ia for calc ium nitrate, as descr ibed above. To investigate the effect of the solution concentrat ion on the coat ing propert ies, methanol and ethanol so ls were also prepared with a dilute concentrat ion of ca lc ium nitrate in the organic solvent by increasing the amount of the solvent to 50 ml (1M), while C a / P ratio was kept at 1.67 (hereafter referred to as 1M solution). The coated wires were dried at 80°C and fired at 500°C in air. S E M was used to investigate the surface of the wires after coat ing. In order to qualitatively study the effect of stress on the coat ing integrity, such as microcracking or de-bonding of the coat ing from the surface, the wires were bent back after the heat treatment. The unbending s imulated the deformation of real stents during p lacement in blood vesse l . The surface morphology of coat ings fired at 500°C for different t ime per iods (10, 20, 30, 40 , 50, and 60 minutes) was studied using S E M . To investigate the effect of the firing time on the coat ings, X R D was done on thin layers of powder heat treated on glass substrates at 500°C for various time periods (10, 30, and 60 minutes). 4.3.2.b Stents Stents obtained from two different s tages of preparation (de-oxidizing and electropolishing) were coated and studied in this sect ion. As- rece ived de-oxid ized coronary stents were coated with H A p solution (1M) using a dip coater with withdrawal speed of 14 cm/min. Coa ted de-ox id ized stents were then 41 heat treated at 500°C for 10 minutes and studied under S E M . Stents were then inflated using a commerc ia l rapid exchange catheter with a semi-compl iant bal loon. The effect of stent deformation on the coat ing quality was investigated with S E M . Electropol ished stents were sur face treated using a 2 .4N hydrochloric ac id aqueous solution at 70°C for 10 minutes and were neutral ized following the s a m e procedure as for the sta in less steel wires. S a m p l e s were then coated with methano l -based sol prepared using 1M calc ium nitrate solut ion, and a dip coater with a withdrawal speed of 14 cm/min. Stents were manual ly shaken after dip coat ing in order to remove the extra solution droplets on the sur face, dried at 80°C and fired at 500°C for a period of 10 minutes. The coat ings were studied under S E M . 42 CHAPTER 5 EXPERIMENTAL RESULTS AND DISCUSSION 5.1 Phase evolution in sol-gel hydroxyapatite 5.1.1 Results 5.1.1.a X-ray diffraction Figures 5.1, 5.2 and 5.3 show the X R D patterns of ca lc ined sol-gel hydroxyapatite (SG-HAp) powders derived from methanol , ethanol and propanol based gels, respectively, heat treated at different temperatures. The methano l -based sample , calc ined at 375°C (Figure 5.1), consis ts of a mixture of ca lc ium phosphates C a 3 ( P 0 4 ) 2 (TCP) and C a 2 P 2 0 7 (PP) , and ca lc ium carbonate. At 425°C smal l peaks appear at 29 = 37.5° and 29 =41.2°, characterist ic of ca lc ium oxide and calc ium nitrate, respectively. At 445°C, a well-crystal l ized ca lc ium nitrate is present a long with T C P and C a 2 P 2 0 7 . Broad H A p peaks also start to appear, at 29 between 31° and 33° (reference: diffraction pattern J C P D C #9-432 (presented in appendix II)). At 460°C amount of amorphous 8 H A p increases, while calc ium nitrate peaks start losing their intensity due to its decomposi t ion to ca lc ium oxide and nitrogen dioxide. At 500°C, a well-crystal l ized hydroxyapatite is present together with T C P and a smal l amount of C a 2 P 2 0 7 . The reflection intensity of HAp peaks has grown at the expense of the T C P , C a 2 P 2 0 7 and ca lc ium oxide. Therefore it is reasonable to bel ieve that reactions occur between calc ium phosphate compounds , such as C a 2 P 2 0 7 and T C P , and the remaining ca lc ium nitrates, carbonates and calc ium oxide to form H A p . At 545°C, HAp is a major phase , along with trace amount of T C P . For the ethanol-der ived gel, a smal l amount of poorly crystal l ized hydroxyapatite is already present at 375°C (Fig. 5.2). No significant change in the X R D pattern occurs by increasing calcinat ion temperature to 425°C or 445°C. At 460°C, poorly crystal l ized HAp and amorphous T C P are present. Crystall inity of H A p (as indicated by the peak width) significantly improves by increasing calcination temperature to 500°C (traces of 8The broad XRD peaks are considered as indicative of amorphous, partially amorphous, poorly crystalline or nanocrystalline character of the compounds; these descriptive words are used interchangeably; this is qualitative assessment and no attempt has been made to determine the degree of crystallization of the powders. 43 T C P are still present). Further increase in temperature to 545°C results in formation of phase-pure H A p (Fig. 5.2). 40 35 30 25 2H Figure (c) 445' 5.1 X R D patterns for methanol -based gels ca lc ined at (a) 375°C (b) 425°C, i°C, (d) 460°C, (e) 500°C and (f) 545°C. 44 m A C a ° A J (b) r a ) ^ ^ ^ ^ ^ 40 35 2ft 30 25 Figure 5.2 X R D patterns for e thanol -based gels ca lc ined at (a) 375°C (b) 425°C, (c) 445°C, (d) 460°C, (e) 500°C and (f) 545°C. Wel l -crystal l ized ca lc ium nitrate (CN) and amorphous apatite phases are present in propanol -based gel powder ca lc ined at 375°C, F ig . 5.3. Upon increasing the temperature to 425°C the amorphous apatite crystal l ized into ca lc ium phosphate compounds such as T C P and C a 2 P 2 0 7 . This trend cont inues till 445°C. At 460°C a poorly crystal l ine H A p develops while ca lc ium nitrate T C P and C a 2 P 2 0 7 are still 45 present. The crystallinity of H A p improves by increasing the temperature to 500°C, while trace amount of T C P is still present. The characterist ic peaks of ca lc ium nitrate are not present at this temperature. Calc inat ion at 545°C results in a well-crystal l ized hydroxyapatite as ev idenced by increased intensity and smal l half-intensity width of the characterist ic reflection peaks . T race amounts of T C P and ca lc ium oxide are also present (calcium oxide has probably formed as a result of decomposi t ion of calc ium nitrate and other carbonated impurities). 40 35 26 30 25 Figure 5.3 X R D patterns for propanol -based gels ca lc ined at (a) 375°C (b) 425°C, (c) 445°C, (d) 460°C, (e) 500°C and (f) 545°C. 46 5.1.1.D Thermal Analysis The thermal gravimetric analys is (TGA) and differential thermal ana lyses (DTA) traces of the three samp les are presented in F igs. 5.4 and 5.5, respectively. Al l the T G A curves start with a rapid weight loss between 100°C and 115°C, attributed to the evaporat ion and desorpt ion of water and organic residues and are accompan ied by endothermic effects in the D T A results. The s lope of T G A curves changes at ~115°C (i.e. when weight loss stabil izes) which stays constant up to 420°C for both methanol and propanol-der ived gels, and 460°C for the e thanol -based gel . Th is weight loss region bears smal l exothermic peaks , which are bel ieved to be due to ol igomerizat ion and polycondensat ion of the hydrolysed spec ies [51]. ProDanol Methanol Ethanol 0 200 400 600 800 1000 1200 1400 Temperature Figure 5.4 Thermal gravimetric analys is of the dried gels . A significant dif ference between the D T A curves for different solvent sys tems lies between 420 and 460°C, where an endothermic peak was detected for the methanol-47 derived gel, whilst no speci f ic peak was detectable for the e thano l -based gel. For the propanol -based gel , an exothermic peak was detected in this temperature region. The exothermic peak in the propanol-dr ied gel at 425°C is attributed to decomposi t ion of the amorphous apatite to T C P and C a 2 P20 7 . Th is effect can be a lso seen in the X R D patterns as the peaks related to T C P and C a 2 P 2 0 7 appear in the X R D patterns after heating to 425°C. The T G A results a lso show a weight loss in this temperature range (420-460°C). The endothermic peak at 445°C in c a s e of the methanol-der ived gel can be attributed to crystall ization of hydroxyapatite together with decomposi t ion of calc ium nitrate. This effect is a lso accompan ied by a weight loss in T G A data. 0 200 400 600 800 1000 1200 1400 Temperature Figure 5.5 Differential thermal analys is of the (a) methanol , (b) ethanol , and (c) Propanol-der ived dried gels. For all solvents, temperatures between 460°C and 510°C bear an exothermic drift in the D T A curves, which is represented by a s low weight loss. Th is exothermic drift can be ass igned to crystall ization of hydroxyapatite. This fol lows by an immediate endothermic peak with a minimum at around 530°C. Th is endothermic peak is 48 represented by rapid weight loss of the gels, which can be cons idered as a result of reaction between C a 3 ( P 0 4 ) 2 / C a 2 P 2 0 7 and ca lc ium nitrate/carbonate to form hydroxyapatite. It is suggested therefore that hydroxyapatite forms from both the transformation of amorphous apatite phase and reactions among C a 2 P 2 0 7 , C a 3 ( P 0 4 ) 2 , C a C 0 3 , and C a ( N 0 3 ) 2 , when the calcinat ions temperature reaches about 500°C. Although the exact pathway for the reaction of T C P and C a 2 P 2 0 7 with other calc ium compounds is not yet clear, the re lease of gaseous by-products of these reactions, i.e. oxides of carbon and nitrogen, is probably responsible for the weight loss observed. 5.1.1 .c Electron microscopy Figures 7 a , 7b, and 7c are the scann ing electron micrographs of methanol , ethanol and propanol -based gels calc ined at 500°C, respectively. Figure 5.6 Scann ing electron micrographs of (a) methano l -based, (b) e thanol -based, (c) propanol -based gels ca lc ined at 500°C. 49 The micrograph of methanol -based gel after calcinat ion at 500°C shows a porous structure with pore s izes in 5-20 urn range. A three-dimensional , amorphous- l ike, and highly porous structure is observed in the micrograph. S o m e crystal growth is observed on the background matrix. Fig 5.6b shows a porous morphology with the pore s ize ranging from 2 to 5 urn. The microstructure of the e thanol -based gel s e e m s to consist of large agglomerates rather than a cont inuous amorphous- l ike structure, as observed in that of the methano l -based sys tem. A porous, cont inuous structure is observed for the propanol -based gel ca lc ined at 500°C (Fig. 5.6c) with the pore s ize ranging between 4 to 16 urn. 5.1.2 Discussion It is obvious from the X R D data that the type of the alcohol (i.e. reflected through its molecular size) plays an important role on the phase evolution of hydroxyapatite in this sol-gel p rocess . A s the organic solvents and calc ium nitrate have polar molecu les, the solubility of ca lc ium nitrate increases with the increase in the polarity of the alcohol solvent, which has been well recognized as a function of dielectric constant. The dielectric constants of propanol , ethanol and methanol are 20 .1 , 24.3 and 32.6, respectively, i.e., dec rease with increasing molecular chain length. Therefore, the dif ferences observed in the phase evolution of hydroxyapatite in different solvent sys tems can be attributed to the solubility of ca lc ium nitrate in different organic solvents. The dissolut ion of ca lc ium nitrate in the alcohol can be s imply expressed as : C a ( N 0 3 ) 2 + x R O H C a ( O R ) x ( N 0 3 ) 2 - x + x N 0 3 " (5.1) Better solubility of ca lc ium nitrate in an a lcohol , should result in larger x value in Eqn (1), due to higher dielectric constant of the organic solvent. It has been proposed previously that it would be more difficult for a calc ium ion chemica l ly bonded to a long-chain alcohol to interact with the hydrolysed alkoxide in the so l , due to, for instance, steric h indrance produced by the alkyl cha ins attached to ca lc ium ion [64]. Liu et. al [50] have recently suggested that the hydrolysed phosphite can react with C a precursor accord ing to the following chemica l pathway: 50 H P O ( O C 2 H 5 ) 2 - x ( O H ) x + C a + + + N 0 3 -> C a - P intermediate + H + (5.2) where the C a - P intermediate represents an amorphous phase , which is able to transform to apatite at e levated temperatures. They have sugges ted that if enough calc ium ions react with hydrolysed phosphite then the C a - P intermediate phase will form and will finally thermally transform into hydroxyapatite. In this case lower activation energy may be required for apatite formation. Apatite can a lso be synthes ized through the interactions among impure phases , such as C a 2 P 2 0 7 , C a 3 ( P 0 4 ) 2 , C a C 0 3 and C a ( N 0 3 ) 2 , as observed in the X R D results, where higher activation energy for the synthesis would be required. It is reasonable to bel ieve that larger ca lc ium-alcohol complex spec ies C a ( O R ) x ( N 0 3 ) 2 . x retard the interaction with phosphorus alkoxide, necessary to form hydroxyapatite. The interactions among impure phases in order to form hydroxyapatite can be descr ibed by the following equation [50]: 1 /2Ca 2 P 2 0 7 +Ca 3 (P0 4 )2+CaC0 3 - i -Ca(N0 3 ) 2 - i -1 /2H 2 0^ C a 5 ( P 0 4 ) 3 ( O H ) + C a O + 2 N 0 2 + C 0 2 (5.3) In the c a s e of methano l -based gels ca lc ined at 375°C, the p resence of impure phases such as C a 2 P 2 0 7 , C a 3 ( P 0 4 ) 2 , C a C 0 3 and C a ( N 0 3 ) 2 is expected due to incomplete reactions represented by equation (5.2). Therefore, better solubility of ca lc ium nitrate in methanol , leading to large value of x in equation (5.1) could be the reason for more difficult interaction between C a ions and hydrolysed phosphi te, which will consequent ly leave equation (5.2) incomplete and will cause formation of impure phases at lower temperatures. However, these impure phases will further interact to form apatite at higher temperatures (Fig. 5.1). It is hypothes ized that intermediate solubility of calc ium nitrate in ethanol compared to methanol leaves enough calc ium ions to interact with hydrolysed phosphite and 51 therefore moves E q n . 5.2 to the right. The presence of amorphous apatite phase in X R D patterns at low temperature supports this hypothesis. Th is amorphous, intermediate phase transforms into H A p upon increase in temperature as ev idenced in F ig. 5.2. The low solubility of calc ium nitrate in propanol results in unreacted ca lc ium nitrate, which in turn leads to incomplete reaction of ca lc ium nitrate and hydrolysed phosphite, Eqn . 5.2. Th is will alter the mechan ism of apatite formation toward the interaction among "impurity" phases ( T C P , C a 2 P 2 0 7 ) rather than crystal l ization from the amorphous phase . A s it is ev idenced from X R D results (Fig. 5.3), the well-crystal l ized calc ium nitrate present at lower temperature will d e c o m p o s e into ca lc ium oxide and N 0 2 with further increase in temperature. A smal l amount of amorphous phase is a lso present at low temperatures (375°C), which d e c o m p o s e s into T C P and C a 2 P 2 0 7 at 425°C. T h e s e "impurity" phases will further react at higher temperatures to form hydroxyapatite. Decomposi t ion of ca lc ium nitrate provides ca lc ium for calc ium phosphate compounds such as T C P and C a 2 P 2 0 7 to form hydroxyapati te (Fig. 5.3). The endothermic peak observed in the range of 500-550°C in D T A curves of all the samp les (regardless of the solvent) is representative of final H A p formation either through crystall ization process or chemica l interactions among different phases . A s shown in F ig . 5.5, there is less energy required to form H A p in an ethanol -base system as compared to methanol- or propanol -based sys tems, as smal ler endothermic peak appears at this temperature range. This effect is produced due to different formation pathways which is consistent with the results of Liu et. a l . [82]. In their work they have speculated that crystall ization of the amorphous intermediate phase in order to form HAp requires less activation energy compared to interaction of impure phases . The crystal s ize of the gels ca lc ined at 500°C was calculated using the Scherrer equat ion: A(2e)=O.9A7Dcos(0) (5.4) W h e r e A(20) represents the peak width at half max imum intensity of the reflection (002), X is the wavelength for C u K a (A.=0.15418 nm), and D is the crystal s ize (nm). 52 This approximate calculat ion reveals that the crystal s ize of H A p formed in different solvents and after calcinat ions at 500°C fol lows the order of methanol (~50nm) > propanol (~40nm) > ethanol (~30nm). Al though, due to the large number of parameters that inf luence different s tages of this complex sol-gel p rocess , suggest ing a conclus ive reason for the observed phenomena , using the characterizat ion methods appl ied here, is impossib le but it s e e m s that phase evolution and crystal s ize in different solvent sys tems is probably affected by the ability of the alkyl group from alcohol solvent to substitute the O E t groups of the phosphorous precursor. The interchange between two alkyl groups can be expressed as fol lows: H P O ( O H ) x (OEt) 2 - x + y R O H ^ H P O ( O H ) x (OEt)2-x-y(OR)y + y E t O H (5.5) where R O H can be methanol , ethanol or propanol . Accord ing to L ivage et. al . [83], both the entropy and enthalpy of the new molecular spec ies will change upon such substitution, leading to a change in chemica l reactivity towards hydrolysis and condensat ion reactions, thus affecting the morphology and the of the final gel [83]. It is therefore expected that in the process studied in this thesis O C H 3 - substitution for O C 2 H 5 - group and a lso substitution of O C 3 H 7 - for O C 2 H 5 - changes the rate of hydrolysis and condensat ion reactions and therefore results in different H A p formation pathways as well as different gel morphology and crystal s i ze . S E M observat ions confirm this hypothesis. Scann ing electron microscopy (Fig. 5.6a, 5.6b and 5.6c) shows a porous, cont inuous three-dimensional structure for both methanol and propanol-der ived gels. However, the ethanol-derived gel shows somewhat different particulate morphology consist ing of smal ler pores and agglomerate particulates rather than a cont inuous, three-d imensional structure. S u c h difference can be attributed to different chemica l pathways during phase evolut ion. In the c a s e of methanol and propanol sys tems, formation of crystall ine intermediate phases such as T C P and C a 2 P 2 0 7 and ca lc ium carbonates, their interaction and re lease of gaseous by-products (especial ly C 0 2 ) results in larger pore s ize , as ev idenced in F igs. 5.6a and 5.6c Moreover , s imi lar phase evolution 53 pathway in methanol and propanol -based systems observed in this study might be the cause for similar morphologies of ca lc ined gels observed in these two systems. For the ethanol-derived gel the main mechan ism of H A p formation appears to be via crystall ization from an amorphous intermediate phase , which results in less porosity in final ca lc ined gel as well as a particulate structure as obvious in F ig . 5.6b. 5.2 HAp coating characterization 5.2.1 Coatings on titanium substrates 5.2.1.a Electron microscopy examination The substrate surface preparation techniques included chemica l treatment and sand blasting. Figure 5.7a is the electron micrograph of the surface of titanium plate, chemical ly treated with phosphor ic ac id . The surface dissolution reaction has resulted in formation of a roughness in a range of 2-3 um on the surface. Figure 5.7b presents the titanium surface after the alternative surface preparation method of sandblast ing. The surface thus produced has a high roughness of about 10 um. Figure 5.7 Scann ing electron micrograph of (a) surface of titanium plate, treated with phosphor ic ac id , and (b) titanium surface after sandblast ing. Figure 5.8a, 5.8b, and 5.8c show the chemical ly treated titanium substrates after spin coat ing with methanol , ethanol, and propanol -based so ls , respectively. The coating surface finish fol lows the morphology of the substrate's sur face regardless of the 54 solvent employed. The coat ing is transparent to the electron beam and therefore the substrate sur face features can be easi ly seen under the coat ing layer. The coatings based on all three solvent sys tems provide a good coverage of the surface, while some microcracks (1-2 um) are observed in the fi lms. T h e s e cracks are much less extensive as compared to those produced by p lasma spraying [80,84] and are probably formed due to drying and sintering strain of thicker sect ions of the coat ing filling the rough surface features. ( c ) Figure5.8 Sur face of chemical ly treated in phosphor ic acid (85%) at 50-60°C for 30 minutes titanium substrates after spin coat ing with (a) methanol , (b) ethanol, and (c) propanol -based solut ions. 55 Figures 5.9a-d show the E D S results of the titanium etched sur face and titanium surface after coat ing with methanol , ethanol and propano l -based so ls , respectively. The appearance of ca lc ium and phosphorous peaks in the E D S spect ra after the coat ing deposi t ion is an indication of the presence of an apatite layer on the sur face. ( a ) ( c ) ( d ) Figure 5.9 Energy dispersive spectrometry of (a) titanium etched surface and titanium sur face after coat ing with (b) methanol , (c) ethanol and (d) propanol -based solut ions. Figure 5.10a-c show the sandblasted titanium sur faces after coat ing with solut ions obtained from methanol , ethanol and propanol solvent sys tems, respectively. It is 56 obvious from the picture that the extent of microcracking increased with the s ize of the alcohol molecule used as the diluting med ia for solution preparat ion. T h e s e cracks are mainly observed in the thicker pockets in cavit ies formed by sur face irregularities. The solution's flow from the top of these irregularities toward the substrate surface produced a reservoir for coat ing solution inside these cavit ies and thus a ticker layer of coating formed in these areas. Features resulting from this down-f low of the solution are clearly observed in the propanol -based coated sur face (Fig 5.10c). Thicker coating translates to more drying and sintering strain in direction parallel to the substrate surface, which will finally results in microcracking. With increase in the molecule s ize of the alcohol solvent, the viscosity and molecular weight of the solution increases. This effect will further increase the solution flow toward the above ment ioned cavit ies and therefore a thicker layer of coat ing will form, which is more suscept ib le to cracking after heat treatment, due to larger strain. Figure 5.10 Sandb las ted titanium sur faces after coat ing with solut ions obtained from (a) methanol , (b) ethanol and (c) propanol solvent sys tems. 57 Both chemica l sur face preparation and sandblast ing, appl ied to improve coating adhes ion to the substrate by increasing mechan ica l interlocking, y ie lded good coat ing coverage on the substrates ' sur faces. Chemica l treatment produces sur face features with less roughness compared to sandblast ing, and results in a more uniform coating th ickness and finally leads to less microcracking. Pract ical ly, chemica l treatment is the only method viable for smal ler objects, such as stents. 5.2.1.b Infrared spectroscopy Figure 5.11 shows the FTIR spect ra of the surface of sandb las ted Ti substrates coated with methanol , ethanol and propanol -based solut ions and fired at 500°C for 10 minutes. The characterist ic v 4 P 0 4 bands at 560 cm" 1 and 600 cm" 1 , v i band at 1000 c m ' 1 and a v 3 P 0 4 absorpt ion band at 1070 cm" 1 are observed , which are typical of apatitic structure [64]. Short-range atomic arrangement in the apatite structure is responsible for the diffuse shape of the peaks [64], although all the peaks are clearly dist inguishable. This effect is produced due to the short dwell ing t ime of the samp les at the heat treatment temperature, i.e. poor crystallinity of the structure. Increasing the soak ing time at 500°C would result in an improved molecular arrangement and higher crystallinity H A p . A smal l peak at 620 cm" 1 can be ass igned to O H group, indicating the p resence of bonded water in the film structure [45]. The broad bands at 1429 cm" 1 and 1494 cm" 1 are attributed to carbonate groups. The broad nature indicates the possib le p resence of organic spec ies such as - C H = C H 2 or - C O O - groups formed due to reactions between the evolved organics and calc ium [28]. The smal l peak observed at 860 cm" 1 on the methanol and ethanol spect ra is a lso related to C 0 3 groups, suggest ing carbonates substitution for P 0 4 group in the apatite structure. Therefore, the coat ings can be cons idered as carbonated apatite, similar to the structure of natural bone. Compar ing P 0 4 bands in the range of 500-600 cm" 1 , the sha rpness of characterist ic peaks dec reases with the s ize of alcohol molecule, i.e. methanol>ethanol>propanol. Cons ider ing the short dwell ing time of the samp les at the firing temperature (10 minutes) the rate of organic residue desorpt ion and decomposi t ion plays an important role in the rate of apatite formation. Therefore, faster desorpt ion of smal ler organic 58 molecules (i.e. in methanol solvent) might be a reason for faster arrangement of apatitic structure. Sur face features (e.g. roughness) are responsib le for the background noise observed in the FTIR spectra. i , , , r 2000 1600 1200 800 400 Wavenumber (cm-1) Figure 5.11 FTIR spect ra of the surface of sandblasted titanium substrates after coat ing with (a) methanol , (b) ethanol and (c) propanol -based solut ions. 5.2.1.c In-vitro bioactivity test The H A p coated titanium substrates were submerged in a s imulated body fluid (SBF) , which has the s a m e ionic components as those found in human blood p lasma (detailed in Append ix I), for 7 days at ambient temperature. Precipitat ion of H A p from the S B F in such a test is cons idered as indicative of bioactivity of the exposed surface [80]. Figures 5 .12a, 5.12b and 5.12c show the S E M micrographs of sur face morphology of titanium substrates coated with methanol , ethanol , and propanol -based so ls , respectively, after incubation in S B F . Scat tered ball-like apatite grains appeared on the surface after the incubation. Microcracks and disintegration of the coat ings are a lso observed in the vicinity of the newly formed grains. The occurrence of these cracks and ball-like apatite deposi ts is bel ieved to be a result of dissolution-reprecipitat ion process between the H A p coat ing sur face and the S B F solut ion. The similar effect and 59 microstructure of the ball-like apatite grains have been observed for p lasma sprayed calc ium phosphate coat ings studied by Liu et. al . [80], under similar S B F incubation condit ions, and cons idered as an ev idence of formation of apatite structure. The newly formed apatite layer is scattered on the sur face in the form of clusters of ball-like grains. The grains are equiaxial , with an average diameter of 2.5 j im, in all the samp les (Figs. 5.12a-c). The HAp "balls" s e e m to have a porous structure, consist ing of agglomerates of smal ler particles. No apparent difference in apatite formation was observed for titanium substrates coated with solut ions made with different organic solvents. This in-vitro test and formation of the apatite layer on the sur face of the HAp coated substrate conf irms the bioactive nature of the coat ings. Figure 5.12 S E M micrographs of surface morphology of titanium substrates coated with (a) methanol , (b) ethanol, and (c) propanol -based solut ions, after incubation in S B F . 60 5.2.2 Stainless steel substrates 5.2.2.1 Stainless steel wires A s no previous attempts of coat ing sta in less steel wires have been reported, the coat ing p rocess has been opt imized on trial-and-error bas is . In this sect ion the results obtained through this p rocess are presented in the sequence of (a) sur face treatment, (b) solution effect and coat ing process , and (c) firing condit ions. 5.2.2.1.a surface treatment The wires were ac id-etched to produce surface roughness to enhance adhes ion of HAp coat ings to the wires. Figures 5.13a-d show the sta in less steel wires after exposure to different acidic reagents in order to produce surface roughening. All the etched samp les were treated in acidic solution of 50 v o l % of hydrochloric, phosphor ic and nitric ac ids in water, for 40 minutes at a temperature between 60-70°C. Figure 5.13a shows the sur face of the as- rece ived stain less steel wire. Sur face features resulted from rolling and greasy spots remained from finishing p rocesses can be observed on the sur face. The greasy spots are removed from the sur face although no sign of etching is observed after exposure to nitric ac id solution as shown in F ig . 5.13b. Phosphor ic ac id produces slight sur face roughening in the form of sur face grooves (Fig 5.13c) while hydrochloric ac id (HCI) results in extended sur face roughening in form of grooves and grain boundary corrosion (Fig. 5.13d) after similar etching procedure. Nitric and phosphor ic ac ids do not s e e m suitable for sur face treatment of sta in less steel wires due to their slight etching effect consider ing the long t ime and relatively high temperature of exposure. The corrosion effects produced by HCI offers the possibil ity of controll ing the etching parameters such as t ime and temperature in order to produce an opt imized surface roughness suitable for the subsequent coat ing procedure. 61 ( c ) ( d ) Figure 5.13 S E M micrographs of (a) as received stain less steel wire and wires after exposure to (b) nitric ac id , (c) phosphor ic ac id, and (d) hydrochloric ac id . In order to investigate the effect of t ime of HCI exposure on the sur face roughness, stainless steel wires were etched for periods of 10, 20, 30, 40, 50 and 60 minutes. The S E M micrographs of these samp les are presented in Figures 5.14a-f, respectively. The effect of HCI on the sur face starts with attacking the grain boundar ies, which cont inues up to 30 minutes of exposure, when grain corrosion starts producing a rougher sur face (Figs. 5.14a-c). Further corrosion of the sur face occurs on the grain sur faces rather than the grain boundar ies, as shown in F igs. 5.14d-f. ( a ) ( b ) 6 2 ( e ) ( f ) Figure 5.14 S E M micrographs of stainless steel wires treated with HCI (2.4 N) for (a) 10, (b) 20, (c) 30, (d) 40, (e) 50, and (f) 60 minutes at 75°C and neutral ized for 4 minutes. A s only slight degree of sur face corrosion is bel ieved to be necessary for coating purposes, and in order to reduce the damage produced by chemica l treatment, 10 minutes etching in HCI at 70°C was chosen as the pre-treatment p rocess . Immediately afterwards the samp les were neutral ized in a process descr ibed earlier. 5.2.2.1.b Effect of coating solution The effect of the coat ing solution has been examined in the two fol lowing sect ions. First the effect of organic solvents on the coat ing properties of the hydroxyapatite sol was investigated. Study of the effects of solution concentrat ion fol lows as the second sect ion. 63 Effect of the organic solvents: It has been pointed out by Br inker and Sherer [12] that the deposi ted film th ickness (h), obtained through dip coat ing in low viscosity solut ions is dependent on the substrate withdrawal s p e e d (U), liquid viscosity (rj), l iquid-vapour sur face tension ( y L v ) , and gravity forces (pgh): r 7 = 0 . 9 4 ( r ) U ) 2 / 3 / y L V 1 / 6 ( p g ) 1 / 2 (5.6) In coat ing sta in less steel wires with so ls prepared with different so lvents, parameters such as L /and pg are kept constant in all the exper iments. The remaining parameters, such as YLV and r\, are affected by the solvent and therefore are var iables. Due to the high complexi ty of the sys tem, caused by p resence of var ious components in the solution (water, organic solvent, ca lc ium nitrate and triethyl phosphite), quantitative analys is of the effect of the solvent on the film th ickness and coat ing propert ies is very difficult. Therefore, the effect of theses variable parameters are investigated qualitatively using S E M images. Figures 5 .15a-c show the picture of bent sta in less steel wire samp les , dip coated with hydroxyapatite sols prepared from methanol , ethanol and propano l -based solut ions, respectively. The wire samp les are unbent before microscopic investigation, to observe the coat ing behaviour in s imulated stent deformation experiment. Crack ing and delaminat ion is observed at the bending point of the methano l -based coated sample (Fig. 5.15a) while only slight microcracking is observed in the vicinity of the bending point. The coat ing remains intact in all the other areas on the sur face of the wire. Accumulat ion of the coat ing solution in the bent sect ion has resulted in an excess ive ly thick coat ing in this a rea . Th is thicker layer is more prone to crack ing due to residual drying and sintering strain and the resulting microcracks pre-exist ing in this layer, which will finally lead to cracking and delaminat ion during unbending. For the ethanol-based coat ing, delaminat ion, cracking and disintegration is observed in the bent sect ion, probably due to a thicker accumulated layer, F ig . 5.15b. Larger extent of cracking and disintegration is observed in the propanol -based coat ing (Fig 5.15c). The cracks are present in the vicinity of the bending point as a result of s t ress in the coating layer. 64 ( c ) Figure 5.15 S E M micrographs of bent stainless steel wire samp les , dip coated in (a) methanol , (b) ethanol, and (c) propanol -based solutions after unbending. A s observed in f igures 5.15a-c the extent of cracking and coat ing disintegration after unbending increases with the s ize of alcohol molecule. Th is can be expla ined by the effect of the solvent on the solution viscosity and surface tension, which will finally affect the film th ickness, especia l ly in a V -shape substrate such as a bent wire. A thicker layer is more prone to cracking due to residual s t resses of drying and sintering. A solvent with a larger molecule s ize increases the th ickness of the coat ing at the bending point, resulting in more cracking after unbending. Consequent ly only methanol and ethanol -based solut ions are chosen for further examinat ions. 65 Effect of solution concentration: Figures 5.16a and 5.16b show the pictures of bent wire samp les coated with dilute (1M) methanol and ethanol -based solutions, respectively. C o m p a r e d to f igures 5.15a and 5.15b, the effect of solution accumulat ion at the bending point and the extent of microcracking are significantly reduced. By increasing the amount of solvent used to prepare the sol (i.e. diluting the sol), the viscosity of the solution dec reases , which will a lso dec rease the th ickness of the coating accord ing to equation 5.6. Reduced th ickness of the coat ing layer results in less strain c a u s e d by shr inkage during drying and sintering, in direction parallel to the substrate (for sub micron sol-gel fi lms most of the drying plus sintering strains appears to be in direction normal to the substrate, and this component of strain does not produce stress in the film). Therefore less cracking is observed in the final fi lm. ( a ) ( b ) Figure 5.16 S E M picture of bent wire samp les coated with dilute (1M) (a) methanol , and (b) e thano l -based solut ions after unbending. 5.2.2.1.c Effect of heat treatment time Figures 5.17.a-c show the bent portion of stainless steel wires coated with dilute (1M) methanol sol after firing for 10, 30 and 60 minutes at 500°C, respectively. The samples were unbent before S E M investigations, in order to simulate system deformation during stent p lacement. After 10 minutes firing, some crystall ization of the thicker layer at the 6 6 bending point is observed (as descr ibed earlier the thicker layer is a result of excess amount of solution trapped in the bent area during dip coating). However, no excess ive cracking or disintegration is observed, F ig . 5 .17a. By increasing the firing time to 20 minutes, enhanced cracking results in the disintegration and removal of the coat ing. Th is damaged a rea has a s ize of - 5 0 (im. No further cracking is observed in the adjacent thinner coat ing, Fig 5.17b. Signif icant cracking is observed on the bent a rea of the sample fired for 60 minutes. The damaged (cracked and delaminated) a rea is about 100um in length. The degree of damaged produced after unbending coated wire samp les increases with the increase in firing t ime. Figure 5.17 S E M picture of unbent stainless steel wires coated with dilute methanol solution and fired for (a) 10, (b) 30, and (c) 60 minutes at 500°C. 67 To examine the extent of crystall ization of the coat ings after var ious heat treatment t imes, X R D was conducted on similarly heat-treated thin fi lms of the so ls on g lass substrates. Figure 5.18 shows the phase evolution of hydroxyapati te after different firing time per iods at 500°C. 10 minutes of heat treatment produces a highly amorphous apatite thin film. Increasing the firing time to 30 and 60 minutes, significantly improves the crystallinity of the coat ing layer as ev idenced by increase in the sharpness (i.e. dec rease of half-intensity width) of hydroxyapati te characterist ic peaks . HAD 60 minutes 30 minutes Jt^^U iu 10 minutes • ^ ^ ^ ^ 40 35 30 25 26 Figure 5.18 X R D results on thin film coat ings fired at 500°C for different t ime periods. It appears that the extent of damage of longer-fired fi lms correlates with the sintering time, which in turn increases degree of crystall ization, density, st i f fness and britt leness of the fi lms. The increased sintering shr inkage produces higher residual stress, and therefore cause further cracking and damage after unbending of coated wires. At the s a m e time, increase in the firing time may result in a better adhes ion between the coat ing and the substrate, as a result of interfacial inter-diffusion between the substrate and the coat ing [85]. However, due to the relatively low temperature of firing appl ied here (500°C), this effect is unlikely. 68 5.2.2.2 S t e n t s The stents evaluated in this work were produced in proprietary process, including the two final steps of sur face c leaning (termed "de-oxidation") and edge rounding ("electropolishing"). Figure 5.19 illustrates surface of a stent after the de-oxidation step. 2-8 urn wide grooves are observed on the surface, and a 2-3 urn thick nickel rich layer (evidenced by E D S ) covers the interior s ide of the stent structure after the "de-oxidation" step . Counts -10000 J 8000 6000 4000 2000 J Nl F igu re 5.19 (a) S E M picture of the stent surface obtained after de-oxidiz ing process, (b) E D S of the ticker layer on the stent surface. 69 Su ch stents were dip coated with a withdrawal speed of 14 cm/min with diluted (1M) methanol solution and fired at 500°C for 10 minutes. The colour of the stent surface changed from silver (before coating) to gold (after coat ing and firing). Figures 5.20a and 5.20b show S E M micrographs of the "de-oxid ized" stent sur face after coating and firing. F i gu re 5.20 (a) and (b) show S E M micrographs of the de-oxidize stent surface after coating with dilute methanol solution and fired at 500°C for 10 minutes. A s it can be observed in these two pictures, the coat ing covers the surface entirely. No microcracking and disintegration or accumulat ion of coat ing solution in the U-shaped areas is observed. A thicker coat ing is observed on the Ni-rich layer present on the interior s ide of the stent. Th is effect is due to the rougher sur face of these areas, which traps the solution during dip coat ing. Figures 5.21a and 5.21b show the E D S results of the front and interior sur faces, respectively. The appearance of calc ium and phosphorous peaks proves the existence of the coat ing while higher intensities of these peaks in F ig . 5.21b demonstrates the thicker layer of the coat ing on the pre-existing Ni-rich layer on the interior surface. Figures 5.22a-b show the coated stent after expans ion with a rapid exchange catheter with a semi-compl iant bal loon at 10-bar pressure. A s it can be observed in F ig . 5.22a no cracking or disintegration is observed on the sur faces of the bar -shaped parts of the ( a ) ( b ) 70 stent structure. S o m e microcracking and delaminat ion of smal l a reas are observed on the coat ing as ev idenced in F ig . 5.22b. The damaged areas are Ni-rich layers delaminat ing from the surface due to their loose bonding to the sur face and also cracking on the coat ing due to its larger th ickness on the Ni-rich layer. Counts 4000 _J 2000 J ( a ) Ni keV Counts 4000 J 3000 2000 _J 1000 _J ( b ) Figure 5.21 E D S results of the (a) front and (b) interior sur faces of the coated stent structure. 71 ( a ) ( b ) F i g u r e 5.22 (a-b) S E M pictures of the coated stent after expans ion . Figure 5.23a and 5.23b show the electropol ished stent prior and after etching in HCI solution for 10 minutes at 70°C, respectively. A s it is ev idenced in F ig . 5 .23a the outer sur faces of the stent is significantly smoothened after electropol ishing, while some surface roughness is observed on the interior sur faces. After etching the samp les in HCI for 10 minutes no significant surface roughening on the sur faces is observed. minutes. Figure 5.24 shows the picture of the etched sample after coat ing. Coat ing shr inkage is observed and therefore a good coverage was not ach ieved . Th is effect is due to 72 inadequate interlocking between the coating and the sur face, which results in the accumulat ion of the solution in the form of droplets and is usually observed on very smooth sur faces. A better coverage was obtained in the interior sur faces due to larger degree of roughness present. F i gu re 5.24 Electropol ished stent (a) prior, and (b) after etching in HCI at 70°C for 10 minutes. 73 CHAPTER 6 SUMMARY AND CONCLUSIONS 6.1 Summary In this work the effect of organic (alcohol) solvents on the synthesis and coating properties of a sol-gel hydroxyapatite is investigated. The novel sol -gel technique has been used to process and deposit the coat ings. X- ray diffraction analys is , thermal gravimetric and differential thermal analys is, scann ing electron microscopy, energy dispersive spectrography and infrared spect roscopy were employed as characterizat ion techniques. Us ing these techniques, the phase evolution of hydroxyapatite in different solvent sys tems (methanol, ethanol and propanol) and their effect on microstructure of the calc ined gels was investigated. T h e coat ing properties are evaluated and d i scussed . Titanium plates and sta in less steel wires were chosen as coating substrates due to their re levance to biomedical industry. Us ing the knowledge obtained through the course of this study, commerc ia l coronary stents were coated with the sol-gel hydroxyapatite. Deformation character ist ics of the coated stents were qualitatively a s s e s s e d and correlated with the process parameters. In this sol-gel sys tem triethyl phosphite and calc ium nitrate were used as phosphorous and calc ium precursors, respectively. In the first step of solution preparat ion triethyl phosphite was hydrolysed in water, and then mixed with a solution of ca lc ium nitrate in an organic solvent. The mixture of the two solut ions was stirred for 30 minutes and aged for 24 hours. G e l s were obtained by drying thus prepared solution at 80°C for 24 hours, which were then ca lc ined at 375 -545C. Thermal behaviour of the gels and phase evolution of H A p upon heat treatment was monitored. This study suggests that different organic solvents induce different pathways for H A p formation in the sol-gel sys tem under considerat ion. The H A p phase evolution depends on the interaction between the solvent and the precursors. Speci f ical ly, it is hypothesized that the interaction of the organic molecule of a lcohol with ca lc ium can affect the subsequent interaction of calc ium-alcohol spec ies with hydrolysed phosphite molecu les. Accord ing to our observat ions, impurity phases of ca lc ium phosphate compounds (such as T C P and P P ) form in the methanol and propano l -based gels. 74 T h e s e phases further react with other ca lc ium bearing compounds (such as calc ium nitrate, ca lc ium oxide and calc ium carbonates) to form hydroxyapati te. Due to the good solubility of calc ium nitrate in methanol , methyl l igands will easi ly attach to calc ium ion. B e c a u s e of the large s ize of thus formed molecu les , and their steric hindrance, reaction between calc ium and hydrolysed phosphi te, necessary to form HAp , would not easi ly complete. A s a result of the incomplete react ion, other (non-HAp) apatite compounds , such as T C P and P P , form along with amorphous apatite intermediate phase . Format ion of H A p depends on the reaction between the impurity calc ium phosphate compounds , and the remnants of the other impurity phases , such as calc ium oxide and calc ium carbonates. Due to the low solubility of ca lc ium nitrate in propanol , unreacted ca lc ium nitrate remains in the sys tem, which results in the formation of ca lc ium deficient apatitic compounds . Therefore H A p formation in propanol fol lows a pathway similar to that in methanol sys tem. In the ethanol sys tem, enough calc ium ions are avai lable to interact with hydrolysed phosphite to form an amorphous intermediate apatite phase , which eventual ly crystal l ises into HAp . This effect can be attributed to the intermediate solubility of ca lc ium nitrate in ethanol. Therefore it is reasonable to bel ieve that the solubility of ca lc ium nitrate in the organic solvent, and the mechan ism of their interaction, will affect the H A p phase evolution pathway in this sol-gel sys tem. The endothermic D T A peak appeared at the range of 520-550°C in all D T A traces, is ass igned to the final H A p formation. The s ize of this peak was observed to be different the different organic solvents used . A relatively smal l endothermic peak was observed in the D T A trace at 530°C for e thanol -based so ls , compared to the larger respective peaks in D T A results for methanol and propanol -based gels. B a s e d on this observat ion, it can be conc luded that the energy required for formation of hydroxyapatite is different in each solvent sys tem. There is less energy required to form H A p in an ethanol -based system as compared to methanol or propanol -based sys tems, possibly due to different formation pathways. Alkyl exchange between solvent and phosphorous alkoxide will take p lace in this sys tem. Th is substitution will affect the rate of hydrolysis and condensat ion reactions leading to different H A p formation pathways in each of the solvent sys tems while 75 affecting the morphology of the gel. A cont inuous, porous structure was observed for the methanol and propanol sys tems. The large pore s ize in these sys tems (-12 pm, on average) can be attributed to evolution of gaseous by-products during the chemica l reactions to form H A p . Ethano l -based calc ined gel has a particulate, porous structure with pores ranging from 2 to 5 pm in s ize . The pore s ize in ethanol sys tem appeared to be smal ler compared to those in methanol and propanol sys tems due to its different HAp formation pathway (crystallization from amorphous phase rather than reaction among impure phases) . Titanium plates were coated with hydroxyapatite through this sol-gel p rocess . In order to increase the adhes ion between the coat ing and the substrate the sur faces of the titanium substrates were roughened using either chemica l corrosion or mechanica l sandblast ing prior to spin coat ing. The coat ings were fired at 500°C for 10 minutes, and character ized using S E M and FTIR. Incubating the substrates in the Simulated Body Fluid (SBF) solution and examining the surface after 7 days of incubation tested the bioactivity of the coat ings. Al l solut ions prepared from different organic solvents provided good coverage on the titanium substrate. In the sandb las ted sur faces, due to the deep surface irregularities (~10pm) the solution would flow into the cavit ies and therefore a thicker layer of coat ing will form which is more suscept ib le to cracking. This effect increases with the molecular weight of the solvent used , i.e. propanol>ethanol>methanol. Faster desorpt ion of solvents with smal ler molecule s ize and faster decomposi t ion of organic compounds made of these molecu les affects the rate of formation and crystall ization of HAp . For the similar heat treatment condit ions, the H A p rate of formation and crystall ization increases with dec rease in the alcohol molecule s ize i.e. methanol>ethanol>propanol, as determined by FTIR. However, no obvious difference between the bioactivity of these coat ings was observed after the S B F incubation test. Medica l grade stain less steel wires (316L) were coated with hydroxyapati te using this sol-gel technique. Short period etching in HCI (10 minutes at 70°C) was found as the most suitable procedure for surface roughening of the wires. The wires were then bent, dip coated into the so l , dried and fired at 500°C for 10 minutes. The amount of excess solution trapped in the bent a rea s e e m s to increase with the organic solvent 76 molecule s ize (me thanoke thanokpropano l ) , especia l ly for the concentrated solut ions (2M). The entrapment of the excess solution produced a thicker coat ing at the bent a rea , which then c racked and delaminated from the substrate 's sur face after unbending. Lower concentrat ion of the solution used for coat ing (1M) dec reases this effect. Increasing the firing time of the wires at 500°C to 30 and 60 minutes, promotes crystall ization and densif icat ion of the coat ing, while increasing its britt leness. Sta in less steel coronary stents were also used as substrates for this H A p sol-gel p rocess . Stents were obtained after de-oxidizing and electropol ishing s teps. The rough surface of the deoxid ized stents provided enough mechan ica l interlocking for the coating purposes. A good coverage of the coat ing was obtained although thicker layers were observed on the nickel-rich layer pre-existing on s o m e sur faces of the stent structure. Crack ing and delaminat ion a lso occurred on this (Ni rich) layer after expans ion of the stent while the coat ing remained intact on other sur faces. The electropol ished stent was etched prior to coat ing in order to produce enough roughness on the sur face to provide mechan ica l interlocking between the coat ing and the substrate. 10 minutes etching in HCI solution at 70°C did not provide sufficient surface roughness and as a result the coat ing was accumula ted on the sur face in the form of droplets, an effect observed in coat ing very smooth sur faces. 6.2 Conclusions From the data resulting from the present investigation, analys is of the literature data and previous studies, the following conc lus ions are drawn: 1. Hydroxyapati te (HAp) can be obtained using different organic solvents, such as methanol , ethanol and propanol , in the sol-gel sys tem of triethyl phosphite and calc ium nitrate as precursors. 2. Analyt ical studies of the materials resulting from the p rocess , in particular X-ray diffraction revealed that these different organic solvents induce different phase evolution pathways for the hydroxyapatite, during this sol -gel synthes is . Th is results in di f ferences in phase composi t ion and crystallinity of the final ca lc ium phosphates. In particular: 77 2.1. In the methanol sys tem "impurity" ca lc ium phosphate compounds , such as tri-ca lc ium phosphate (TCP) and C a 2 P 2 0 7 form that will further react with other impurity phases such as ca lc ium oxide and calc ium carbonates, to eventual ly form the hydroxyapatite. 2.2. Propanol sys tem a lso promotes the formation of impure apatite compounds due to the low solubility of ca lc ium nitrate in propanol , leaving behind unreacted calc ium nitrate. Therefore H A p formation fol lows a mechan ism similar to that of the methanol sys tem. 2.3. Due to the complete reaction between ca lc ium ions and hydrolysed phosphite, an amorphous intermediate apatite phase forms in the ethanol -based system that will further crystall ize into hydroxyapatite. 2.4. The crystal s ize of the final hydroxyapatite formed in each of these sol-gel sys tems fol lows the order of methanol>propanol>ethanol after heat treatment at 500°C when the samp les were removed instantly upon reaching this temperature which is bel ieved to be affected by the HAp formation pathways. 3. A s observed in D T A data, the energy required to form hydroxyapati te in different solvent sys tems fol lows the order of methanol>propanol>ethanol. 4. S E M observat ions showed different morphologies for the ca lc ined gels obtained from the different solvent sys tems. A three d imensional and highly porous structure was observed for the methanol and propanol-der ived gels, with the pore s ize ranging between 2-20 um. Ethano l -based gel has a particulate structure with much smal ler pores (2-5 urn). 5. Ti tanium plates were spin coated with this sol-gel sys tem. 5.1. Sma l l microcracks (1-2 um) were observed in the coat ings deposi ted on chemical ly treated (in 8 5 % phosphor ic acid at 50-60°C for 10 minutes) titanium sur faces. The cracks were however much less extensive as compared to p lasma sprayed H A p coat ings. 5.2. The extent of microcracking increased with the s ize of the solvent (alcohol) molecule used for sol preparation, for the coat ings on sandb las ted titanium sur faces. This effect is due to sol accumulat ion in the cavit ies of sur face irregularities produced by sandblast ing, which caused thicker layer formation in these a reas . 78 5.3. Infrared spect roscopy (FTIR) on the surface of coated titanium plates showed the presence of carbonated hydroxyapatite, similar to that of the natural bone. 5.4. FTIR study also revealed that the apatite crystallinity (for the s a m e heat treatment schedule) dec reases with the s ize of the a lcohol molecu le i.e. methanol> ethanol> propanol , possibly due to the faster resorption of smal ler organic spec ies . 5.5. Al l the obtained coat ings were proved to be bioactive through the in-vitro Simulated Body Fluid (SBF) incubation test. Coat ings obtained from different solvent sys tems showed no apparent difference in bioactivity, when tested in S B F (simulated body fluid). 6. Sta in less steel wires were bent (to simulate geometry of the stent) and coated with hydroxyapatite through this sol-gel process, with the fol lowing results: 6.1. Hydrochlor ic acid is a suitable agent to roughen the sur face of s ta in less steel wires to enhance its mechan ica l interlocking with the coat ing. Both dilute and concentrated HCI are suitable for this purpose. 6.2. Organic solvents with smal ler molecular s ize result in less entrapment of the sol in the bent a rea of the wire and therefore less cracking and delaminat ion is produced after unbending, as ev idenced by S E M observat ions. 6.3. S E M micrographs show that lower concentrat ion of the solvent (i.e. 1M) reduces the amount of excess solution in the bent a rea . 6.4. X R D results show that increase in the firing time promotes crystall ization of the coat ing layer. However, more crystall ine layer is more prone to cracking under appl ied st ress due to the higher amount of residual sintering strain, and the expected higher sti f fness of the coat ing. 7. Coronary stents were a lso coated in this so-gel sys tem. 7.1. De-ox id ized stents were successfu l ly coated with this sol-gel sys tem. S E M observat ions conf i rmed a good coverage of the coat ing on the stent sur face. 7.2. Both E D S and S E M examinat ions revealed the p resence of a thicker coat ing on nickel-rich layers, pre-existing on some sur faces. This effect is p roduced due to the larger degree of roughness present on these layers. 7.3. Slight cracking and delaminat ion of the coat ing was observed after expans ion of the stent. Th is effect was mostly observed on thicker coat ing present on Ni-rich 79 layers. Delaminat ion was also occurred due to de-bonding of theses Ni-rich layers from the surface. 7.4. 10 minutes etching of the electropol ished stent in HCI solution at 70°C was not sufficient to produce enough surface roughness and as a result the coat ing solution was accumula ted on the surface in the form of droplets, an effect usual ly observed on very smooth sur faces. 80 CHAPTER 7 RECOMMENDATION FOR FUTURE WORK The future work should focus on the optimization of the H A p sol-gel coat ings on various medica l grade substrates such as stain less steel stents. T h e main chal lenge in the coat ing of the stents is to reduce the surface roughening necessa ry for coat ing adhes ion while maintaining enough adhes ion to prevent delaminat ion of the coat ing after expans ion of the stent. The extent of chemica l sur face preparat ion (involving surface corrosion), enough to provide mechan ica l interlocking, should be optimized through examining the coat ing adhes ion to the stent sur face. Th is parameter can also be ba lanced with coat ing th ickness. The effects of different dip coat ing parameters, such as dipping s p e e d , on the quality and th ickness of the final coat ing remains to be investigated. A s mechan ica l properties of the stent should not be affected by the coat ing and surface preparat ion procedures, the effect of sur face treatment and firing time and temperature on the mechanica l properties of stent such as fatigue and wear resistance should be studied. The rate of coat ing dissolution into the physiological environment determines the bioactivity of the coat ing while affecting the coat ing integrity throughout the stent operation inside the body. Therefore parameters affecting the rate of coat ing dissolut ion such as the degree of crystallinity of the coat ing and a lso coat ing th ickness should be investigated. Finally, in-vivo and in-vitro bioactivity tests shou ld follow the success fu l coat ing procedure. 81 NOMENCLATURE Latin Symbols D crystal s ize (Eqn 5.4) h film th ickness (Fig 2.2, Eqn 5.6) S stagnation point (Fig 2.2) U withdrawal speed (Fig 2.2, Eqn 5.6) Greek Symbols Y i v l iquid-vapor sur face tension (Eqn 5.6) 5 boundary layer (Fig 2.2) q v iscosity (Eqn 5.6) A wave length (Eqn 5.4) A(20) peak width at half max imum intensity (Eqn 5.4) Abbreviation ACP amorphous ca lc ium phosphate CN ca lc ium nitrate DCP monetite, C a H P 0 4 DCPD d icalc ium phosphate dihydrate HAp hydroxyapatite, C a 1 0 ( P O 4 ) 6 ( O H ) 2 OCP octacalc ium phosphate, C a 8 H 2 ( P 0 4 ) 6 . 5 H 2 0 PP C a 2 P 2 0 7 SEM scann ing electron mic roscope TCP tr icalcium phosphate, C a 3 ( P 0 4 ) 2 T T C P tetra ca lc ium p h o s p h a t e , C a 4 P 2 0 9 XRD x-ray diffraction 82 REFERENCES [I] H. R. Piehler," The Future of Medic ine: Biomater ials," M R S bulletin, p. 67-70, Aug . 2000. [2] L. L. 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In H B P (mM) Na+ 142.0 142.0 K+ 5.0 5.0 Mg+ 1.5 1.5 Ca + 2.5 2.5 c r 147.0 103.0 HC03" 4.0 27.0 H P O 4 2 1.0 1.0 SO2" 0.5 0.5 A P P E N D I X II Standard X-ray diffraction data for hydroxyapati te. 9- 432 JCPBS-ICDD Copyright (c) 1991 Rad.= 1.54056 Quality: i Ca (PO ) (OH) 5 4 3 Calciui Phosphate Hydroxide Hydroxylapatite, syn Rad: CuKal Laibda: 1.54056 Fi l ter : d-sp: D.S. -114.6 Cutoff: Int: I/Icor: Ref: de Holff, Techniscfc Physische Dienst, Delft, Netherlands, JCPDS Grant-in-Aid Report Sys: Hexagonal a: 9.418 A: Ref: Ibid. Dx: 3.15 Di: 3.08 S.8.: P63/I (176) c: 6.884 C: SS/FDH: F30=54(.016.35) A: 7: C: .7309 •p: ea: nwB: 1.651, ey: 1.644, Sign: - 2V: Ref: Dana's Systei of Mineralogy, 7th Ed., 2 879 Color: Sta-green. asparagus-green, bluish green, grayish green also, violet, violet-blue, aiethystine, sotetites colorless, pale greenish white, gray, brown, flesh red, rose-red. clear blue Saiple obtained following the procedure indicated by Hodge et a l . . Ind. Eng. Chei. Anal. Ed., 10 156 (1938). CAS no.: 1306-06-5. I / f l are peak values froi a pattern which shows slight broadening of prist reflections. Validated by calculated data 24-33. Apatite group, apatite subgroup. PSC: hP44. To replace 34-10. Hvt: 502.32. VoluteiCOl: 528.80. 2-theta Int. h k 1 : 10.820 12 1 0 0 16.841 6 1 0 1 ; 18.785 4 1 1 o : 21.819 10 2 0 0 ! 22.902 10 1 1 1 25.354 2 2 0 1 ' 25.879 40 0 0 2 ! 28.126 12 1 0 2 ! 28.966 18 2 1 0 i 31.773 too 2 1 i : 32.196 60 1 1 2 ! 32.902 60 3 0 0 ! 34.048 25 2 0 2 : 35.480 6 3 0 i : 39.204 8 2 1 2 ! 39.818 20 3 1 0 ! 40.452 ! 2 2 2 i : 42.029 10 3 1 1 ! 42.318 ! 4 3 0 t i & i 43.804 8 1 1 3 i 44.369 2 4 0 o ': 45.305 ! 6 2 0 3 : 46.711 30 2 2 2 : 48.103 : 16 3 1 2 ! 48.623 6 3 2 o : 2-theta Int.! h k 1 t 2-theta Int.! h k 1 '. 2-theta Int.! h k 1 49.468 40 : 2 1 3 ! 60.457 6 : 3 3 1 ! 73.995 7 : 4 2 3 50.493 20 ! 3 2 1 ! 61.660 1 i o : 2 1 4 ! 75.022 3 : 3 2 4, 6 51.283 12 ! 4 1 0 ! 63.011 12: 5 0 2 ! 75.583 9 : 2 1 5 52.100 16 : 4 0 2 ! 63.443 ' 4 : 5 1 0 ! 76.154 ' l : 4 3 2 53.143 20: 1 0 0 4 ! 64.078 13 : i 3 0 4, 3 2 3 ! 77.175 I t u : • 5 1 3 54.440 4 ! 1 0 4 ! 65.031 9': 5 1 1 i 78.227 9 j 5 2 2 55.879 io : 3 2 2 ! 66.386 : 4 : 4 2 2, 4 1 3 • 57.128 8 : 3 1 3 ! 69.699 3 : 5 1 2 • 1 58.073 ' 4 : 5 0 1 ! 71.651 ! 5 : 4 3 1, 4 0 4 i i • i 1 1 53.938 6 : 4 2 0 ! 72.286 4 : 5 2 0 2 0 5 : 1 0 2 Strong lines: 2.81/1 2.78/6 2.72/6 3.44/4 1.84/4 1.94/3 2.63/3 2.26/2 

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