The role of refractory element based coatings on the tribological and biological behaviour of orthopaedic implants

Journal of

Materials Processing Technology

ELSEVIER

Journal of Materials Processing Technology 56 (1996) 364-374

THE ROLE OF REFRACTORY ELEMENT BASED COATINGS ON THE TRIBOLOGICAL AND BIOLOGICAL BEHAVIOUR OF ORTHOPAEDIC IMPLANTS L.P. Ward, K.N. Strafford, T.P. Wilks, C. Subramanian Surface Engineering Research Group, Advanced Materials & Processing Research Centre, Dept. of Metallurgy, University of South Australia, Australia

ABSTRACT In the area of orthopaedic implants, particularly hip joint replacements, only three groups of metals and alloys are considered suitable as load - bearing implant materials, namely Ti - Ti alloys, Co - Cr alloys and 316L stainless steel. As a result of problems, largely associated with failure of the anchoring poly methyl methacrylate (PMMA) bone cement, the survival of present-day load - bearing prostheses cannot be guaranteed beyond 10 to 12 years. This limited lifespan has essentially limited joint replacements to the elderly. However, the increased number of younger patients requiring replacements has accentuated the need for joint prostheses with improved performance and longevity. The introduction of joint prostheses however with significantly enhanced durability in turn accentuates problems associated with implant loosening, wear, corrosion and, in particular, biological incompatibility. At present, very few monolithic materials are capable of satisfying all service requirements. The solution may lie in the use of coating technology, creating new materials / surfaces with unique combinations of properties that cannot be achieved from the use of monolithic materials alone. The fh-st part of this paper reviews the types of coatings that have been used, or are currently being considered as implant materials, to include porous metallic coatings, "bio-active" (hydroxyapatite, bioglasses) and "bio-inert" (alumina) ceramic coatings. Their advantages and limitations will be discussed. The second part will focus on recent developments in the use of refractory element based coatings, with specific reference to work conducted on sputtered niobium and tantalum. Possible avenues for research to include advanced coating (multilayered / multicomponent) systems, surface treatments and duplex engineering will be reviewed.

1.

INTRODUCTION

Osteoarthritis and rheumatoid arthritis are disabling diseases, due to the degeneration of the articular cartilage and eventual destruction of the human joint, the most common being the hip joint. At present there is no known cure for this disease; relief comes through partial or complete replacement of the affected joints by artificial implants. Total hip joint replacements are constructed from either 316L stainless steel, Co-Cr alloys, Ti alloys or, to a lesser extent A1203, for fabrication of the femoral components, and ultra high molecular weight polyethylene (UHMWPE) for fabrication of the acetabular cup. While prostheses fabricated from these conventional material have a life-span of approximately 10 to 15 years, the demand for implants with long term retention (20 to 30 years) and enhanced durability has stimulated much interest in the use of composite materials, particularly coating composites to provide the necessary requirements associated with longer service life, such as enhanced corrosion resistance and biocompatibility, increased wear resistance and strong anchorage of the implant to the surrounding bones. 0924-0136/96/$15.00 © 1996 Elsevier Science S.A. All rights reserved SSD10924-0136 ( 95 ) 01850-E

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C U R R E N T STATUS OF COATINGS AS I M P L A N T M A T E R I A L S

M e t a l l i c c o a t i n g systems Metallic coatings deposited on metal substrates have been used primarily to provide an effective means for implant fixation by tissue ingrowth. Such systems have included Co-Cr-Mo coatings on Co-Cr-Mo substrates [ 1, 2] and Ti-6AI-4V coatings on Ti-6A1-4V substrates [3]. Studies [4] of porous coated Co-Cr-Mo hip prostheses showed that although with these implants cementless fixation was achieved, histological examination of many retrieved implants revealed problems due to metallic release and failure of the coating caused by the loosening of the implant beads. It was implied that although porous coated prosthetic devices cannot be used to replace cemented prosthetic devices entirely, they nevertheless offer viable means of producing additional bonding. 2.1

B i o - inert c e r a m i c coating systems Brossa et al [5] studied the effect of plasma-sprayed alumina and titania coatings on metal ion release when deposited on 316L stainless steel, molybdenum and titanium substrates. The presence of an overlay TiO2 coating inhibited metal ion release [Fe, Cr, Ni] into fresh human serum, whereas A1203 coating enhanced the metal ion release. This was attributed to the porous nature of the latter coating, leaching of the metal ions occuring through pores. The use of a more dense coating, produced by varying the process deposition parameters (eg. powder size, spraying distance, particle speed and substrate temperature [5]), would certainly be more advantageous. Escudero et al [6] observed a reduction in the corrosion resistance of plasma-sprayed A1203 coatings deposited on 316L stainless steel substrates, when an intermediate coating (Ni-Al, Ni-Cr or 316L stainless steel), normally used to improve the adherence between the substrate and the coating, was deposited on the base material. This was attributed to the formation of a galvanic cell, making the base material more active. "In - vitro" and "in - vivo" studies of plasma sprayed A1203 coatings deposited on 316L stainless steel and Ti-6A1-4V substrates, indicated a substantial loss in coating / substrate adherence, and coating cohesive strength, when exposed to physiological media at 37oc for several weeks. This aging effect, observed both "in - vitro" and "in - vivo", was thought to be associated with the porosity of the coating, leading to the conversion of A1203 to AI(OH)3. Furthermore, the possible release of substantial quantities of Al ions into the body promoting toxic effects "in - vivo" cannot be ruled out. In spite of these chemical and biological problems associated with such coating systems, Trentani et al [7] reported on the clinical and X - ray results of 5 cases of uncemented total hip replacements, with a ceramic acetabular component and an A1203 - coated Ti stem. Here after 10 years implantation no revision of any of the prosthetic components was necessary. A method for the formation of an inexfoliable alpha - A1203 layer on the surface of Fe-Cr-A1 alloy, by heat treatment in air, was suggested to be a possible solution to adhesion problems observed with plasma sprayed A1203 coatings [8]. The superior "in - vitro" corrosion resistance exhibited by the alloy compared to 316L stainless steel, even without the presence of the coating, combined with excellent tissue adherence properties "in - vivo", indicated the potential of this material as a substitute for 316L stainless steel. However, at present there is no indication as to the long term effects of this material "in - vivo". Granular A1203 coatings deposited upon A1203 substrates have been shown to exhibit the highest degree of bonding / strength effects at the bone / implant interface, when compared to granular hydroxyapatite coatings, bulk A1203 and dense hydroxyapatite (HA) [9, 10]. This was attributed to mechanical interlocking effects. The value was even higher than that of bulk hydroxyapatite. 2.2

B i o - a c t i v e c e r a m i c coating systems Bio - active coatings, either calcium phosphate-based ceramics or bioglasses, have been deposited on metallic substrates with a view to achieving enhanced bone / implant fixation by chemical bonding. The disadvantages associated with poor strength properties of "bulk" hydroxyapatite (HA) were overcome by depositing the material as coatings on suitable substrates. Histological examination of HA plasma-sprayed Ti-6A1-4V alloy, after implantation in dogs, revealed mineralisation of interface bone directly onto the coated implant surface, with no fibrous tissue layer interposed between bone and HA, visible at light microscopic level, Furthermore no evidence of HA resorption was observed "in - vivo", even after 32 weeks implantation [ 11]. Lemons [12] from further studies on HA, both in bulk and coating form, confirmed the biological and mechanical stability, but expressed doubts about the long term behaviour of the coating in relation to resorption of the coating and consequent reactions at the substrate / tissue interface. Further uncertainties surround the effect of fatigue and shear conditions on crack propagation and eventual fracture within the coating, or at the two interfaces, and the effect of the various coating process parameters on the strength properties and bioactivity of the coating.

2.3

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The most common method for depositing HA is by plasma / flame spraying. This method of deposition has been shown to produce no adverse bonding effects with respect to the tensile and fatigue strengths of titanium - alloy substrates, yielding typical values of 70, 60 - 100, 55 and 60 MPa for tensile bond strengths of HA coa.ting to Ti, dense HA in bulk form, HA coating to bone, and bone alone, respectively [13]. The methods of deposition have included dipping the substrate into HA powder followed by sintering; dipping the substrate into molten HA; sputter coating and hot isostatic pressing [14]. Sputter coating appeared to produce an extremely dense, well-adherent coating. However, a non - crystalline Ca-P ceramic was obtained in more than 50 % of coating attempts, rather than the desired crystalline HA. Another major disadvantage with sputter deposition is the difficulty in producing thicker coatings due to the relatively low deposition rates using this technique. The poor mechanical properties of bioglasses permits them to be used in the construction of joint replacements only in coating form. Bioglass has been deposited on 316L stainless steel and Ti alloys by plasma spraying and hot dipping [15]. A metal-enriched interfacial glass layer formed by ionic diffusion of metallic constituents [ 16], controlled the coating / matrix adhesion. Improved adhesion has been reported with suitable doping of the bioglass, resulting in diffusion of ions with higher mobility, to favour glass - metal bonding. Fe203 has been added in various concentrations to bioglass, to enhance adhesion to 316L stainless steel [17]. In one study, increased coating-to-substrate adhesion was achieved by cutting channels or depressions into the base metal to provide mechanical interlocking [18]. Cyclic load tests carried out on partially-coated prostheses, using joint simulators, displayed a high degree of resistance as long as the coating remained intact and adhered well. However, cracking and spallation of the coating from the metal caused infiltration of physiological liquids between the two materials, due to capillary action, leading to accelerated corrosion-induced degradation of the coating [18]. The transmission of overall applied stresses to the prosthesis, through the bone / glass bond to the glass / metal interface, is ensured if good adhesion develops between the glass and bone interface. Thus good coating / substrate adhesion is essential for the device to function effectively. Although A1203 has been shown to exhibit excellent mechanical and biological stability in a body environment, it does not promote direct bonding to bone. Consequently attention has focussed on the deposition of bioactive ceramic coatings on A1203 substrates to provide the necessary bone bonding capabilities, without any sacrifice in the mechanical properties. A1203 ceramic coated with bioglass showed very limited bonding with bone, probably due to the increased diffusion of A1203 into the bioglass. However, this problem was overcome by depositing two bioglass layers of different compositions, on to the A1203 surface. The first layer, contaminated with A1203 by diffusion, had physical and chemical properties closer to the ceramic substrate, while the second layer had a nearto-ideal bioglass composition [ 19]. Oonishi et al [20] deposited hydroxyapatite (HA) coatings on A120 3 substrates by plasma spraying and sputtering. Direct anchorage to bone was observed for coatings deposited by both techniques. In the case of sputtered HA, bone was in direct contact with the materials, even when the HA coating disappeared.

2.4

Other coating systems

Shaw and Miller [21] developed a technique whereby bone was rf sputtered onto 316L stainless steel and A1203 prostheses. Sputter targets were manufactured from bone chips obtained from steer bone. Spectrographic analysis of the coating indicated material transferred on sputtering was identical to actual bone, while the chemical bonding was virtually identical to bone and the same molecular bonding was present. However, there is no data relating to the behaviour of such coated prostheses in a body environment, with respect to mechanical, chemical and biological stability. Carbon - PTFE (Poly-tetrafluroethylene) and A1203 - PTFE coating composites have been deposited on metallic hip prostheses to provide cementless fixation [22]. Plasma sprayed C - PTFE is undergoing clinical trials for knee and hip joint replacements. The low elastic modulus of A1203 - PTFE acts as a resilient interface between the metal implant and bone, and hence mediates against persistent postoperative pain. Metallic femoral prostheses, fabricated from Co-Cr alloy, have been coated with porous polysulphone by sintering particles of the neat resin. The results indicated ingrowth of bone into the pores leading to strong direct bone bonding [23]. The development of pyrolytic carbon films deposited on various substrates by the sputtering technique, which has the advantage of producing carbon films at low temperatures, has found use in the manufacture of critical cardiovascular prostheses due to its excellent haemocompatibility [24] . Furthermore, its excellent biocompatibility and high resistance to fatigue and wear, suggests a promising future for the material in the manufacture of load - bearing implants, if deposited on a suitable base material. "In - vitro" tests conducted on diamond-like carbon (DLC) coatings deposited on plastics used in bio - engineering revealed excellent biocompatibility, [25]. Here, two types of mouse cells - primary mouse peritoneal macrophages and mouse fibroblast cells - were allowed to grow in the presence of the coating, without any signs of inflammatory responses or destruction of cell integrity.

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367

R E F R A C T O R Y E L E M E N T BASED C O A T I N G S AS I M P L A N T M A T E R I A L S

In recent years, attention has focussed on the use of Nb and Ta as implant materials as a result of their outstanding bio - compatibility, superior corrosion resistance and excellent fatigue [26-28]. However, in the pure form, Ta and Nb are mechanically weak. This inadequacy in strength has excluded their use for the construction of load - bearing prosthetic materials, although strengthening can be achieved through the application of powder metallurgy techniques. An alternative method to benefit from the biocompatibility of Ta and Nb, without any sacrifice in overall component strength, is to deposit Ta and Nb on a suitable substrate. Here the rationale assumes that the substrate provides the necessary mechanical properties associated with load - bearing implants, while enhanced bio compatibility and corrosion resistance are provided by the Nb or Ta coating. In-depth studies have been conducted by Ward and co - workers on the deposition and characterisation of sputter ion plated Nb and Ta coatings, and evaluation of their physical, mechanical, chemical and biological properties with respect to biomedical applications [29-35]. In this section of the paper, selected results on the structure of Nb coatings deposited on 316L stainless steel as a function of the process deposition parameters are presented, along with the corresponding wear, fatigue, corrosion and biocompatibility properties of the coatings.

3.1

Deposition of Nb coatings

All Nb films were deposited in a Leybold - Hereaus L560 standard coating unit, by the ion - plating technique, using a magnetron sputter target as the vapour source. Of all the deposition variables associated with the coating process (substrate bias voltage, argon gas pressure, film thickness, deposition rate, substrate temperature and surface finish), substrate bias voltage and argon gas pressure were observed to have the strongest influence on the structure and properties of the Nb coatings.

3.2 Morphology of Nb coatings The morphology of Nb coatings deposited on 316L stainless steel as a function of the substrate bias voltage are shown in Fig. 1 (a) - (f). Here, the argon gas pressure was maintained at 1.1 Pa, the deposition rate was 21 to 23 A s-1 and the resulting film thickness was in the range 5 - 9 ~m. At 0 V bias the structure observed was typical of vapour deposited species, consisting of tapered crystallites with domed tops, separated by voided boundaries ( Fig. 1 (a)). This was typical of a zone - I structure as reported by Movchan and Demchishim [36]. The cross section of the coatings revealed a poorly defined columnar structure containing a number of large voids, as shown in Fig.

1 (b). On the application of a bias voltage of - 200 V, a finer-grained, more dense, columnar zone - 1 structure was produced, as a result of ion bombardment at the substrate / coating surface. At high bias voltages of -1000 V, a dramatic change in the microstructure occured, whereby thick columnar grains with highly faceted surfaces separated by distinct intercrystalline boundaries were observed, as shown in Fig. 1 (c) and 1 (d). Such a structure is typical of a zone - 2 structure, normally occuring in the T/T m range where the growth process is dominated by adatom surface diffusion (36). At a higher bias voltages of - 2000 V, a fine - grained equiaxed structure was observed, as shown in Fig. 1 (e) and 1 (f) respectively. This is typical of a zone - 3 structure occurring at the T/Tm ratio where bulk diffusion effects have a dominant influence on the final structure of the coating.

3.3

Wear and Fatigue Results for Nb Coatings

Fig. 2 shows the results of the pin-on-disc experiments for 316L stainless steel and Nb deposited on 316L stainless steel as a function of the sliding distance. Here, ultra high molecular weight polyethylene (UHMWPE) was the counterface material (pin). The tracking speed and load were optimised and held constant at 282 cm s -1 and 40 N respectively for all tests carried out. Zone - 1, zone - 2 and zone - 3 structures were induced by depositing the coatings at -200 V, -1000 V and -2000 V substrate bias voltage respectively [argon gas pressure = 1.1 Pa, deposition rate = 21 to 23 A s-1 and film thickness = 8 ~tm]. Wear results for zone - 1 and zone - 3 structures were very similar, showing an initial running in period, with some scatter in weight change values up to a sliding distance of 5.076 x 105 cm. After this period a steady state wear was achieved showing an almost constant weight loss with sliding distance. In contrast to the results obtained for the zone - 1 and zone - 3 type coating structures, the uncoated 316L stainless steel after approximately 1.3 x 106 cm sliding distance showed an increase in the wear rate [higher weight loss], although a very similar running in period was observed. However, the wear behaviour for coatings having zone - 2 structures showed a completely different trend. Large weight gains were initially observed, followed by very erratic positive and negative weight changes with increasing sliding distance. Furthermore, no steady state wear trend was observed as those exhibited by the zone - 1 and zone - 3 type coating structures.

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L.P. Ward et al. / Journal of Materials Processing Technology 56 (1996) 364-374

Fig. 1. Scanning electron micrographs of Nb films deposited on 316L stainless steel as a function of the bias voltage; (a) OV, as deposited surface; (b) OV, cross section; (c) -1000V, as deposited surface; (d) -1000V, cross section; (e) -2000V, as deposited surface; (f) -2000V, cross section.

L.P. Ward et al. /Journal of Materials Processing Technology 56 (1996) 364-374

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Fig. 3 shows the crack initiation data, in the standard foma-applied stress versus no of cycles (S/N)-for 316L stainless steel and Nb deposited on 316L stainless steel at -200, -1000 and -2000 V, yielding zone - 1, zone - 2 and zone - 3 structures respectively. All coatings were deposited at 1.1 Pa argon gas pressure, 21 to 23 ~ s -1 deposition rate and had a thickness of 8 lam. The results from the S/N plot show some of the characteristics of failure by fatigue. At high stress values of 5.196 X 108 and 4.157 X 108 N/m 2, specimens failed at a relatively low number of cycles - 1000 and 3500 cycles respectively. Further decrease in the applied stress was accompanied by a further increase in the no. of cycles to failure. It was observed at approximately 2.24 X 104 cycles, the curve appeared to start to level off, forming a knee at stress levels of approximately 3.0 + 0.1 X 108 N/m 2. At stress levels of 2.598 X 108 N/m 2 and below, no failure of specimens was observed in 5 X107 cycles. This was taken as the fatigue limit behaviour. Analysis of this region of the curve, yielded a fatigue limit of 2.8 + 0.2 X 108 N/m 2. The general conclusion, taking into consideration the scatter in the results of all specimens tested, was that the presence of an overlay Nb coating, at all zone structures, had no adverse effects on the resulting fatigue life.

3.4 Corrosion and biocompatibility results for Nb and Ta coatings Polarisation curves for Nb and Ta coated stainless steel, together with those of bulk Nb, bulk Ta and 316L stainless steel are depicted in Fig. 4. The plots are presented as standard voltage (E) versus log current (i) curves. The results show that while the 316L stainless steel specimens corroded freely in 4 % NaC1 solution, [typically 108.9 mA cm -2 at +1.0 mV], bulk Nb and Ta exhibited enhanced corrosion resistance, yielding values of 0.13 and 0.20 mA cm -2 respectively at +1.0 mV. Although the presence of a Ta and Nb overlay coating did increase the corrosion resistance of the 316L stainless steel, the resistance of the coatings were poor compared to their respective bulk materials, yielding values of 1.78 and 2.68 mA cm -2 for coated Nb and Ta respectively. A further study of the effect of substrate bias voltage on the corrosion behaviour of Nb coatings (Fig. 5), revealed enhanced corrosion resistance was observed for coatings deposited at -1000 V and -2000 V, corresponding to zone - 2 and zone - 3 type structures, yielding values of 3.36 and 3.95 m A c m -2 at +1.0 V respectively. Films deposited at lower substrate bias voltages [0 to -600 V] with zone - 1 type structures showed, in general a reduction in the corrosion resistance, compared to zone - 2 and zone - 3 type coating structures. The corrosion resistance of the zone - 1 type structures was observed to decrease in the following order; -200 V > 0 V > -600 V, yielding values of 12.74, 19.09 and 46.42 mA cm -2 respectively. Biocompatibility studies of Nb coatings was assessed by examining the effect of coated specimens on the growth and viability of mouse peritoneal macrophage [WEHI] and B - lymphoma [A20] cells. Fig. 6 shows the effect of Nb deposited on 316L stainless steel and that of uncoated 316L stainless steel on the growth of WEHI cells, at an initial concentration of 1 x 105 cells per well. The results were compared with those obtained for standard tissue culture polyethylene, serving as a control. For all materials, an increase in the number of cells per well was observed during the initial 48 h period, reaching a maximum of 2.3 x 105, 1.8 X 105 and 1.75 x 105 cells / well for 316L stainless steel, Nb coated 316L stainless steel and polyethylene control respectively. Prolonged exposure after this period, resulted in a decline in the number of live cells. This was accompanied by a marked increase in the number of dead cells, even after 8 days exposure. The effect of Nb coated stainless steel and 316L stainless steel alone on the growth and viability of A20 cells, at initial concentrations of 1 X 105 cells / well, are depicted in Fig. 7. The maximum growth of these cells was achieved after 72 hours exposure, attaining a maximum of 9.7 x 105, 7.6 X 105 and 6.85 X 105 cells / well for polyethylene, Nb coating and 316L stainless steel respectively. Further exposure time resulted in a decrease in the number of live cells and a rapid rise in the number of dead cells. In contrast to studies conducted on coated and uncoated stainless steel and polyethylene with WEHI cells, a high proportion of A20 cells were still alive after 7 days exposure, yielding values of 4.6 X 105, 3.85 X 105 and 4.2 X 105 cells / well for polyethylene, Nb coating and 316L stainless steel.

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ON T H E P O T E N T I A L R O L E O F R E F R A C T O R Y E L E M E N T N I T R I D E C O A T I N G S F O R ORTHOPAEDIC APPLICATIONS

An important factor which has to be taken into consideration is resistance to wear, particularly wear of the UHMWPE acetabular cup when in contact with the femoral component. It is suggested the deposition of refractory element niwide coatings of the type MX, where M = Ti, Zr, Hf, Ta, Nb, Cr and X = N may reduce degradation of the polymer and provide the necessary corrosion resistance and biocompatibility. Preliminary studies on the

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corrosion behaviour of TiN coated 316L stainless steel substrates in physiological media, showed potential for this coating for use in a biological environment [37]. Here, scope lies in the study of novel, multi - component Ti based coatings such as (TiA1)N, (TiA1V)N, (TiCr)N and (TiZr)N, where the possible formation of a passive oxide layer of the alloying elements might be expected to enhance the performance in the body. The use of a multi - layered coating system may have other significant advantages. An example of such a system may be Ti / TiN / TiA1N, where the Ti acts as a bond coat between the substrate and TiN, and an outer layer of TiA1N to provide a means of passivation in the body environment. A combination of two or more coatings may be required in order to achieve optimum performance of the prosthetic device. For example, a bioactive coating may be deposited on the stem of the femoral component of a hip joint replacement, to achieve the required bonding to the surrounding tissue, while a thin, hard tribological coating may be deposited on the femoral head to reduce friction and wear and release of corrosion / degradation products. Although many advances have been made in the area of bulk biomaterials, particularly in orthopaedic applications, it is becoming increasingly recognised that advanced coating technology may provide the answers to long term problems associated with prosthetic devices, and thus provide materials with a longer term, trouble free life - span. Finally the potential for duplex surface engineering-that is, the use of a suitable coating of the type discussed earlier in combination with a surface hardened, eg plasma nitrided substrate is worthy of consideration in future R and D programmes. The creation of such a coating system may offer even more advantages in the drive to develop prostheses where overall property requirements - fitness-for-purpose-may be more closely tailored to precise requirements.

ACKNOWLEDGEMENT The authors wish to thank Dr. T.Parry of the University of Durham U.K. for advice in the writing of this paper. The authors, and in particular L.P.Ward also wish to acknowledge the useful discussion and advice rendered by Professor P.K.Datta of the University of Northumbria, Newcastle-Upon-Tyne, U.K.

REFERENCES [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] [12] [13] [14] [15] [16] [17] [18]

R.M. Pilliar, Clinic. Orthop. 176 (1983) 95 C.A. Engh, Clinic. Orthop. 176 (1983) 52 A.C. Fraker, A.J. Bailey, H. Hahn and R.H. Rowe Jr., 2 nd World Congress on Biomaterials, Washington D.C. April 27 - May 1 (1984) p.169 M. Spector, J. Arthrop. 2 (2) (1987) 163 F. Brossa, B. Looman, R. Pietra and E. Sabbioni in P. Vincenzini (ed) Ceramics in Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.99 M.L. Escudero, J.A. Gonzalez and J. Ruiz, Br. Corr. J. 22(3) 1987, 182 C. Trentani, A. Montagnani and G. Vincenzini in A. Pizzoferrato, P.G. Marchetti, A. Ravaglioli and A.J.C. Lee (eds) Biomaterials and Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.177 H. Oonishi, H. Okabe, T. Ike and E. Tsuji in P. Ducheyne and J.E. Lemons (eds) Bioceramics: Material Characteristics Versus "In - Vivo" Behaviour, Ann. N. Y. Acad. Sci. 523 (1988) 38 Y. Kakutani, T. Yamamuro, Y.Kotoura and I.Noda in P. Vincenzini (ed) Ceramics in Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.387 Y. Kakutani, T. Yamamuro, Y. Kotoura and I. Noda, J. Biomed. Mater. Res. 23 (1989) 781 S.D. Cooke, K.A. Thomas, J.F. Kay and M. Jarcho, Clin. Orthop. 232 (1988) 225 J.E. Lemons, Clin. Orthop. 235 (1988) 220 K. de Groot in P. Vincenzini (ed) Ceramics in Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.381 W.R. Lacefield in P. Ducheyne and J.E. Lemons (eds) Bioceramics: Material Characteristics Versus "In Vivo" Behaviour, Ann. N. Y. Acad. Sci. 523 (1988) 72 A. Krajewski, A. Ravaglioli and G. Perugini, Science of Ceramics 11 (1981) 85 A. Ravaglioli and A. Krajewski, 2 nd World Congress on Biomaterials, Washington D.C. April 27 - May 1 (1984) p.285 M. Cannas, M.R. Amadeo, A. Krajewski and A. Ravaglioli in A. Pizzoferrato, P.G. Marchetti, A. Ravaglioli and A.J.C. Lee (eds) Biomaterials and Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.563 A. Ravaglioli, A. Krajewski, P. Ponti, V. Valmori, M. Messori and A. Moroni in P. Vincenzini (ed) Ceramics in Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.91

374

[19] [20]

L.P. Ward et al. / Journal of Materials Processing Technology 56 (1996) 364-374

P. Ducheyne, J. Biomed. Mater. Res. 19 (1985) 273 H. Oonishi, S. Miyamoto, A. Kohda, H. Ishimam and E. Tsuji in A. Pizzoferrato, P.G. Marchetti, A. Ravaglioli and A.J.C. Lee (eds) Biomaterials and Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p.69 [21] B.J. Shaw and R.P. Miller, U.S. Patent 3, 918, 100 (1975) [22] J.W. Boretos, Chemtech. (1987) 225 [23] M. Spector, R.J. Davies, E.M. Lunceford and S.L. Harmon, Clin. Orthop. 176 (1983) 34 [24] P. Arra, M. Santi, F. Vallona, G. Majni, G. Ottaviani and A. Paccagnella in P. Vincenzini (ed) Ceramics in Clinical Applications, Elsevier Science Publishers, Amsterdam (1987) p. 117 [25] A.C. Evans, J. Franke and P.J. Revell, Internal Communication, Ion Tech Ltd [26] G. Pfluger, H. Plenk, N. Bohler, F. Grundschober and S. Schider in G.D. Winter, D.F. Gibbons and H. Plenk Jr (eds) Biomaterials 1980, J. Wiley and Sons Ltd (1982) p.45 [27] H. Plenk, G. Pfluger, S. Schider, N. Bohler and F. Grundschober in E. Morscher (ed) The Cementless Fixation of Hip Endoprostheses, Springer- Verlay Berlin (1984) p. 174 [28] H. Plenk Jr, G. Pfluger, N. Bohler, F. Gottsauner - Wolf, F. Grundschober and S. Schider in P. Ducheyne, G. Van der Perre and A.E. Aubert (eds) Biomaterials and Biomechanics 1983, Elsevier Science Publishers, B.V. Amsterdam (1984) p.61 [29] P.K. Datta, K.N. Stratford, D.S. Lin, L.P. Ward, R. Hill and G.J. Russell, Thin Solid Films 168 (1988) [30] P.K. Datta, L.P. Ward, R. Hill and K.N. Stratford in Surface Engineering Practice - Processes Fundamentals and Applications in Corrosion and Wear, K.N. Stratford, P.K. Datta and J.S. Gray (eds) Ellis Horwood, (1990) p. 89 [31] L.P. Ward, Ph.D. Thesis, University of Northumbria, UK. (1991) [32] P.K. Datta, K.N. Stratford and L.P. Ward in "Corrosion - a Tax forever", Conf. Proc. 28 th Australasian Corrosion Association Inc, Nov. (1988) [33] P.K. Datta, K.N. Stratford, L.P. Ward and C. Dick in "The Changing Role of Orthopaedics". Conf. Proc. I. Mech. E, 14 - 15 April, London (1989) [34] P.K. Datta, K.N. Stratford and L.P. Ward in Advanced Materials Engineering, D.M.R. Taplin, D. Taylor (eds) IMF 6, Proc. 6 th Irish Materials Forum, Dublin, Ireland Sept 6 (1989), Parsons Press, Ireland (1989) p. 348 [35] L.P. Ward and P.K. Datta. Paper presented at Irish Materials Forum 7 conference, Limerick, September (1991). Published in proceedings. [36] B.A. Movchan and A.V. Demchishin, Phys. Met. Metallogr. 28 (1969) 83 [37] J. Gluszec, J. Jadrkowiak, J. Masalski and W. Popytak in K.N. Stratford, P.K. Datta and J.S. Gray (eds) Surface Engineering Practice - Processes, Fundamentals and Applications in Corrosion and Wear, Ellis Horwood Publishers, England (1990) p. 233