Optimization of Porous Titanium Layers for Cementless Implantation Technology

Transcription

1 Optimization of Porous Titanium Layers for Cementless Implantation Technology U. van Osten 1, A. Salito 2, F. Breme 1, M. Aits 1, K. Hufnagel 1 1 GfE Metalle und Materialien GmbH, Nuremberg, Germany 2 Sulzer Metco AG, Wohlen, Switzerland 1

2 1. Introduction Surgeons estimate that approximately 1% of the population of highly industrialized countries suffer from joint damage which usually causes intense pain. More than a half a million total hip prosthetic replacements are implanted around the world each year. Several million people have hip implants which last an average of 10 years. Although surgeons and engineers have been working on this medical and economic problem for almost 100 years, experts have still not been able to agree on how implants should be designed. However, the joint implants should meet the basic requirements listed below. Biocompatibility achieved by use of special materials selected to ensure that the entire system (i.e., also including particles resulting from wear, products of corrosion, anchoring systems, and so on) is totally accepted by the body Optimization of implant bonding within the bone to guarantee early mobilization of the patient Simple, reproducible implantation technique using varied implant design adapted to the individual patient and a set of special instruments Consideration of possible complications (e.g., allergic reactions and loosening of the implant) Bio-mechanically optimized construction to improve freedom of movement and tolerance of exercise High-quality materials with excellent tribological properties, static and dynamic stability, a high degree of toughness and a rigidity adapted to the bone (i.e., Young`s modulus) of the implant In addition to this long list of requirements, availability of materials and production costs must be considered. Among the various demands placed on a cementless implant, surface design engineering plays a particularly significant role since the bone grows onto this surface, thus affecting the quality of the bond between bone and implant. Comprehensive studies have shown that bone ingrowth is not only dependent on the material and biological properties of the surface of the implant but also on its surface design. See table 1. Pore diameter 75 to 350 micron Porosity 20 % to 40 % Surface roughness Ra > 3.5 micron Layer thickness max. 500 micron Table 1: Geometric requirements placed on the surface [1,2,3,4] It has been shown that to be effective, open pores should have a minimum size of 75 micron to ensure ingrowth of mineralized bone. Much faster ingrowth was demonstrated with 300 micron pores. However, the limitations of healing the opening become evident when the size exceeds 400 micron. Pore sizes from 75 micron to 350 micron can be viewed as ideal. Literature offers little information on optimization of porosity. Hahn and Palich [4] specify a value of 20% to 40%. A minimum roughness Ra of 3.5 micron has proven to be favorable for the promotion of sufficient primary fixation and simultaneous growth-stimulating micromovements. In addition, it has been shown that the thickness of the porous structure does affect anchorage strength but a layer thicker than 500 micron does not provide further significant improvement. When considering the surface structure, one must remember that natural movement places dynamic stress on the joint prostheses, and hip endoprostheses in particular. However, endurance limit is reduced by the processes used to create surface structures. 2

3 For example, when titanium spheres are sinterfused on the shaft, thermal stress reduces the fatigue endurance limit of titanium alloys by approximately one-third [5]. During corundum blasting, damaging of the surface reduces the fatigue endurance limit by up to two-thirds [3]. In contrast, when applying porous structures with the VPS process (i.e., Vacuum Plasma Spraying) presented here, losses can be kept within a justifiable range (i.e., < 15%) since the coating process is performed at low substrate temperatures. 2. The Vacuum Plasma Spraying Process (VPS) During plasma spraying, an electric arc is generated between two water-cooled electrodes in a gun. The arc heats a gas to an extremely high temperature, partially ionizing it and igniting a plasma jet. See figure 1. Figure 1: Plasma spraying process Since gas temperatures of up to 20,000 C can be achieved, the gases are accelerated by the tremendous expansion in volume and pass through the jet-shaped anode at a high speed. The powder for the coating is injected using a carrier gas. In the plasma gas stream the powder particles accelerate to a high speed, are melted and impact the surface of the substrate with high kinetic energy. Porous to dense layers are created on the substrate by adjusting the spraying parameters. The dynamics of the coating process keep the temperature of the implant surface below 500 C [6,7]. Atmosphere composition and pressure are important variables during plasma spraying. Various processes can be obtained by adjusting these parameters. See table 2. Gas Pressure Composition of the spraying atmospere Inert Gas Reactive Gas P < 1 bar Vacuum PS - - P = 1 bar Inert PS Reactive PS Atmos- P > 1 bar - Table 2: The different plasma spraying (PS) processes Air pheric PS When vacuum or inert-gas plasma spraying is used, interactions between the plasma jet, powder, substrate and the surrounding atmosphere are reduced significantly. This is indispensable for coating with titanium, which is a sensitive material to oxidation and nitrogen absorption. In contrast, during reactive plasma spraying, a reaction between plasma gas and powder is desired. During VPS, the sprayed particles have more kinetic energy than when under normal pressure since the particles are subjected to less gas friction and impact loss during their flight through the reduced atmosphere. In addition, the enthalpy density of the plasma stream is a function of the pressure. This permits controlled formation of layers, either high density or controlled porosity. Adhesion and cohesion are also improved [6]. The VPS coating process has been used for many years for medical purposes and has been validated as a reliable industrial process. Figure 2 shows a VPS system during which the plasma flame generated in the plasma spray gun is controlled by a robot. The implants are placed in batches on a rotary table. 3

4 To ensure the high reproducibility of the coating process, all procedures performed by the robot and the substrate carrier are co-ordinated by a computer and monitored. As already mentioned in section 1, it is important to ensure that the roughening procedure does not generate too much stress concentration during corundum blasting since this would have a negative effect on the fatigue limit of the substrate. Careful adjustment of the blasting parameters has made it possible to maintain an adhesion strength of the titanium layer on the substrate of > 70 N/mm 2 in accordance with ASTM C despite a reduction in corundum grain size from < 1400 micron to < 150 micron. During adhesion tests, the adhesive failed whereas the bonded substrate layer did not. At the same time, by decreasing the size of the corundum grain, the surface roughness of the substrate was reduced from 5 micron to 1.4 micron thus improving the fatigue endurance limit of hip shafts. Figure 2: VPS Coating plant 3. Porous VPS Titanium Coatings for Hip Endoprostheses Use of biocompatible materials is a prerequisite for the production of implants. This is particularly true of the powder used for the coating since the bone has to grow into the porous layer with a large surface. Titanium stands out from other biocompatible materials (e.g., niobium and tantalum) not only because of its superior mechanical properties but also because it is less expensive [9]. 3.2 Titanium Powder for the VPS Process A number of processes can be used to produce titanium powders. Figures 3 to 6 show scanning electron microscope images of powders produced by different processes. 3.1 Substrate Pretreatment As with other coating processes, the VPS process also requires pre-treatment of the substrate to provide sufficient adhesion strength of the coating material. Physical roughening with corundum blasting is used to improve spraying powder bonding and thus adhesion. Figure 3: Magnesium-reduced titanium sponge powder 4

5 The different particle shapes of the powders are clear to see. Sponge powder has a very irregular grain shape, high porosity and a large specific surface. Due to the production process and the high porosity, sponge powder contains a relatively high number of impurities. See table 3. In addition, the irregular shape of the particle makes it very difficult to control the plasma spraying process. Ele- Ti- Sponge Ti-Sponge REP HDH ment Mg-Red. Na-Red. Ti-Powder Ti-Powder Figure 4: Sodium-reduced titanium sponge powder Al C Cl Fe Na Mg H N O Table 3: Typical impurities in titanium powders in wt% Figure 5: Titanium REP-Powder Figure 3 shows a magnesium-reduced powder and figure 4 a sodium-reduced sponge powder. The very irregular shape of the particles is quite obvious. This lack of uniform shape causes excessive gas content (e.g., in the form of oxides) and an irregular layer structure which is extremely difficult to reproduce. In contrast, the particles produced by the rotation electrode process (i.e., REP) have a uniform spherical shape as shown in figure 5. This permits precise definition of the layer structures. The disadvantage of this powder is its high cost since the yield of the small grains required for the spraying process is very low. In addition, the low specific surface of the spheres hinders the transfer of heat from the plasma to the powder particles. Bonding surfaces in the layer made with spherical particles are much smaller than with other powders. Figure 6: GfE HDH-Powder 5

6 The three types of powder described above are not very suitable for use in the optimized porous VPS titanium layer applied to cementless implants to permit bone ingrowth. For this reason, a special powder was developed by GfE based on the HDH method (i.e., hydrogenization-dehydrogenization) which not only meets the requirements placed on the coating itself and provides the reproducibility necessary for spraying but also does not exceed reasonable production costs. Figure 7 shows a diagram of powder production starting with the initial titanium sponge. The decisive factor of this method of production is that a uniform powder with a high degree of purity and no secondary porosity is obtained by melting and mechanical crushing. See figure 6. and the coating process are relatively low. This is due to the fact that production yield of the required powder fraction and shape during the HDH process is high and reproducibility of the coating process is excellent. This means that the spraying parameters must only be modified slightly from one batch of powder to another. Typical distribution of particle size of the powder is shown in figure 8. The coating process can be controlled exactly, and it is possible to obtain the precise layer structure desired. In addition, costs for both powder production Figure 8: Typical distribution of particle size Figure 7: Production of HDH titanium powder 6

7 3.3 Optimized Titanium Layer with the VPS Process When coating hip shafts with titanium, the primary goal is to achieve the ideal surface structure without reducing the fatigue limit of the implant. The ultimately desired porous layer is obtained by a coarse spraying powder, but coarse particles create large indentations when they hit the substrate directly with high kinetic energy. The cross section of the treated implant surface shown in figure 9 has been so damaged by assignment of unsuitable parameters that the fatigue endurance limit has been reduced by more than 45%. However, use of slightly modified parameters and particle grain sizes < 45 micron for the coating produces a compact layer. The smaller particles with their lower kinetic energy cause less damage to the substrate, and subsequent deformation by the larger particles of the porous layer does not occur on the substrate directly but in the intermediate layer which handles deformation well. Damaging stress concentration is significantly reduced by this intermediate layer. In addition, when a porosity of 5 to 15% is used for the intermediate layer, any tears are "lost" in small pores. Figure 10 shows the interface of such an intermediate layer system. The more compact intermediate layer which was sprayed on the implant without damaging is clearly seen. A second layer with optimized morphology is applied to this intermediate layer to promote bone infiltration. This second layer is characterized by an open porosity of various sizes which are bonded to additional pores within the layer. This permits bone to grow in and establish a bond deep within the metallic layer. Figure 11 shows the surface of this layer demonstrating that the ideal geometrical requirements were achieved precisely with a pore size of 75 to 350 micron and a porosity of approximately 30%. Figure 10: Cross section of the intermediate layer system with GfE HDH titanium powder Figure 9: Damage to the surface of the implant caused by use of unsuitable coating parameters (i.e., stress concentration) 7 Figure 11: Surface of the coating with open pores

8 Maximum Load (N) Figure 14 presents a summary of strength values for uncoated hip shafts and shafts coated by various methods tested according to ISO Since the type of shaft was tested with the smallest load cross section, higher strength values could be calculated for larger shafts. Optimization of the spraying process and use of the intermediate-layer system increase component stability up to 5400 N (i.e., corresponds to a stability loss of < 15%) without compromising on the geometrical requirements of the layer necessary for ideal bone infiltration. Figure 12: Cross section of a porous coated dental implant. Blue areas indicate mineralized bone. Figures 12 and 13 show the infiltration of mineralized bone using a dental implant porously coated with HDH titanium powder as an example. New bone growth shown in blue is clearly evident. The example clearly shows that this technique permits optimal bone infiltration and quick and permanent primary fixation of an implant. Particularly the caverns formed in the porous layer gives the bone an opportunity to anchor itself to the coating. This ensures permanent long term secondary fixation uncoated 3700 conventional VPS coating 5400 optimized VPS coating Figure 14: Fatigue limit of hip joint GSS-CL 1. No fractures after 5 x 106 load cycles (Standard: ISO ) Figure 13: Cross section of a porous coated dental implant. Blue areas indicate mineralized bone. 8

9 Figure 16: Surface of the fracture Figure 15: The GSS-CL 1 hip joint Figure 15 shows the GSS-CL 1-type shaft. The size of this type of shaft is of very filigree design and is suitable for treatment of hip malformations. The small shaft cross section requires optimal coating and a high degree of fatigue of the implant. Figure 16 shows the surface of a fractured implant. One can see that the layer has not become detached even when the shaft is broken. As expected, the failure has occurred in the lateral area of the implant. The failure in coated and uncoated implants is identical if an intermediate layer has been applied as shown in figure Summary Implants can be firmly anchored in the body by using porous coated surfaces. Until now, a significant loss in fatigue strength caused by stress concentration during coating could not be ruled out for coatings made of titanium or titanium alloys. Comprehensive studies have been done to determine how to use the vacuum plasma spraying technique with specially optimized GfE- HDH titanium powder for medical applications to produce porous layers with a high degree of reproducibility and without weakening the implant material. The entire production process was examined in detail and critical points identified. These efforts have brought about a significant increase in mechanical stability and coating reproducibility (e.g., particularly for very filigree hip endoprostheses) which is far beyond the minimum requirements of international standards placed on hip implants. 9

10 5. References [1] Bobyn J.D., Pillar R.M., Cameron H.U. und Weatherly G.C.: The optimum pore size for the surface of porous surfaced metal implants by the undergrowth of bone; Chemical Orthopeadics and Related Research, No. 150 Juli/August (1980) [2] Cameron H.U., Pillar R.M. und Macnab I.: The rate of bone ingrowth into porous metal; J. Biomed. Mat. Res., Vol. 10 (1976), [3] Pillar R.M., Lee J.M. und Maniatopulus C.: Observations on the effect of movement on bone ingrowth into porous surfaced implants; Clinical Orthopaedics and Related Research, No. 208 Juli (1986), [4] Hahn H. und Palich W.: Preliminary evaluation of porous metal surfaced titanium for orthopedic implants; J. Biomed. Mater. Res., Vol. 4 (1970) [5] Cook S.D., Georgette F.S. und Skinner H.D: Fatigue properties of carbon and porous coated TiAl6V4 alloy; J. Biomed. Mater. Res., Vol. 10 (1984) 497 ff. [6] Salito A., Barbezat G.: Filmer H., Plasma-Technik AG Wohlen-CH und Trotta F., Plasma-Technik AG Rome-Italy: Plasma Sprayed Techniques for Biomedical Ceramic Coatings. [7] Die Plasmapore-Beschichtung für die zementlose Verankerung von Gelenkendoprothesen, N.N. Hrsg. Aesculap; Wissenschaftliche Information Nr. 22, [8] van Osten U., Sattelberger S.: Glien W.: Maßschneidern poröser Titan-Schichten für die biomedizinische Anwendung; 1995, Biomedizinische Technik Bd 40, Ergänzungsband 1, [9] Clarke E.G. und Hickman J.: An investigation into the correlation between the electrical potential of metals and their behaviour in biological fluids; J. Bone Joint Surg., Vol. 35 B (1953) 467. [10] Sattelberger S.: Die Prothepor-Beschichtung für die zementfreie Gelenkendoprothetik; Keramed Symposium Regensburg,