|Biocompatibility evaluation of nickel-titanium shape memory metal alloy:|
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The corrosion resistance of the implant alloy is a very important determinant of its biocompatibility. As pointed out above, the nature of the environment and the surface treatments have a marked influence on corrosion. Most of the knowledge on the corrosion behavior of NiTi is from studies of dental arch wires and in vitro conditions. In fact, the knowledge of the corrosion behavior of NiTi inside the body is very limited.
The good corrosion resistance of NiTi in sea water was first reported by Buehler et al. (1967). The attempts to evaluate the corrosion in a simulated physiological environment and the comparisons with other implant alloys were made much latter.
Speck et al. (1980) found that, in Hank’s solution, titanium materials, including NiTi, have better corrosion resistance than Co-Cr-Mo or 316L stainless steel. The addition of cysteine amino acid to the solution caused a lower breakdown potential for Ti-Ni, but did not affect the breakdown of Ti-6A1-4V.
Edie et al. (1981) compared the corrosion of used and unused NiTi and stainless steel orthodontic wires. They concluded that NiTi wires are no more subject to corrosion than stainless steel.
Sarkar et al. (1983) found NiTi to be more sensitive to corrosion than titanium in 1% NaCl solution. Pitting of the NiTi surface was observed, and they speculated that this pitting could be due to selective dissolution of nickel from NiTi.
When NiTi was tested in artificial saliva, the release rates of nickel from stainless steel and nickel-titanium arch wires were not significantly different (Barrett et al. 1993).
Better resistance to the chemical breakdown of passivity was found for the NiTi alloy compared to AISI 316 LVM (American Iron and Steel Institute, and low vacuum melted) (Wever et al. 1998).
When stainless steel (316L) was coupled with NiTi and subjected to an immersion corrosion test in 37° C, 0.9 wt% sodium chloride solution, 316L was found to suffer from crevice corrosion (Platt et al. 1997).
Contrary to the above studies, Rondelli (1996) found that the Ni ion release was three times higher for NiTi than for austenitic stainless steels when evaluated in physiological simulating fluids. NiTi had good resistance to pitting similarly to Ti6-Al-4V. Tests in which the passive film was abruptly damaged indicated that the characteristics of the passive film formed on NiTi are not so good as those on Ti6-Al-4V, but are comparable or inferior to those on austenitic stainless steels.
Montero-Ocampo et al. (1996) found annealed NiTi to be more corrosion-resistant than cold-worked material. Thus, the heat treatment and mechanical working had a significant influence on corrosion behavior. The same study also indicated that straining of NiTi led to significant improvements in corrosion resistance. This may be due to the development of a single martensite variant during deformation.
Castleman et al. (1976) reported no generalized or localized corrosion on NiTi plates under microscopic examination at magnifications of 50x with a maximum follow-up of 17 months after implantation in dogs. Neutron activation analysis of distant organs in the same study showed no accumulation of trace metals from NiTi.
When Cragg et al. (1993) implanted forty-four NiTi intraluminal stents in the iliac arteries of 22 sheep, only minimal corrosion was seen at 6 months. Pitting was the predominant type of corrosion. They estimated the mean pit penetration rate to be approximately 0.0046 cm per year. Corrosion product analysis around the pit sites indicated that the main product of pitting was a titanium-bearing compound, probably an oxide. The clinical importance of this finding is not known, because no such corrosion studies have been performed on other stent materials in similar conditions.
As there is some dissolution of nickel from NiTi, some surface treatments have been introduced to improve corrosion properties.
The titanium nitride coating of NiTi prepared by the arch ion plating method was found to improve corrosion resistance in 0.9% NaCl solution (Endo et al. 1994).
In the next two pioneering studies by the same author, the corrosion resistance of the NiTi alloy was enhanced by chemical modification with human plasma fibronectin via aminosilane (γ -APS) and glutaraldehyde as coupling agents. It was found that the corrosion rate decreased by approximately 50% with this surface modification in a 0.9% NaCl solution and a cell culture medium containing serum. The reduced corrosion rate was accompanied by a significant reduction in the release of nickel ions from the NiTi alloy substrate (Endo 1995b). The greatest insight of the above treatment was the idea to introduce biofunctional protein precoating to acquire desirable adhesion properties of the NiTi surface (Endo 1995a,b). The stability and durability of this surface remain unassessed.
A plasma-polymerized tetrafluoroethylene (PPFTE) coating has been used to improve the corrosion resistance of NiTi plates and corresponding NiTi stables. A PPTFE coating improved the pitting corrosion resistance. The passivation range increased from 35% to 96% compared to an untreated sample, and the pit diameter decreased from 100 microns to 10 microns (Villermaux et al. 1996). The coating complies with the large deformations induced by the memory effect of the alloy without cracking. However, when the film is damaged, corrosion seems to increase in comparison to untreated samples. A surface of this kind may be suitable to stent applications, but its durability in orthopedic surgery may be insufficient because of the fretting surface loads.
The addition of Cu raises the repassivation potential of NiTi and may improve its corrosion resistance. This was not proved in a study by Wen et al. (1997), but the corrosion potential and corrosion rate of Ti50Ni50-xCux (x = 1, 2, 4, 6, 8) alloys are irrelevant to its Cu content and the values are almost the same as those of NiTi alloys.
The study by Iijima et al. (1998) showed that small amounts of Cr and Cu added to change the super-elastic characteristics do not change the corrosion resistance of the Ni-Ti alloy in simulated physiological environments.
The laser surface treatment of NiTi (i.e. melting of the surface) leads to an increase of the oxide layer, a decrease of superficial Ni and an improvement of the cytocompatibility of NiTi up to the Ti level (Villermaux et al. 1997).
Surface chemistry may be a more important determinant of platelet behavior than surface topography. There were no cytotoxic or hemolytic effects of any of the surface-treated NiTi samples (annealed, polished or shot-peened). However, platelet spreading (size) after attachment showd dependence upon the NiTi surface treatment and was more abundant for NiTi compared to titanium and stainless steel (Armitage et al. 1997).
The covering of NiTi stents with biodegradable PLA material has been shown to be inferior in in vivo conditions. Elastic mismatch of the non-elastic coating and the self-expandable NiTi stent led to misplacement and vessel occlusion, probably due to PLA filaments fraying into the vessel lumen (Schellhammer et al. 1997).
In the most recent study by Trepanier et al. (1998), the effects on the corrosion resistance and surface characteristics of electropolishing, heat treatment, and nitric acid passivation of NiTi stents were studied. The results show that all of these surface treatments improve the corrosion resistance of the alloy. This improvement was attributed to the removal of the plastically deformed native oxide layer and its replacement by a newly grown, more uniform one. The authors concluded that the uniformity of the oxide layer, rather than its thickness and composition, seems to be the predominant factor to explain the improvement of corrosion resistance.