Fatigue Damage of Nitinol Stents in Simulated Physiological Solution

Nowadays, stent-grafts are commonly used in vascular surgery. Stent-graft manufactures are confronted with two basic requirements: stents must have an “infinite’’ life, stents must be made of the “thinnest’’ wires (especially those at the brain). Stent-graft failure or device fatigue remains major concern for stent-graft manufactures and researches. Self-expanding stent-grafts are made of nitinol. The stent-grafts are mechanically loaded, and also the device is placed in very aggressive environment. The corrosion stability of Nitinol is strongly dependent on the surface preparation: grinding, polishing, chemical etching. This article deals with fatigue degradation of stent-grafts in corrosive environment. Introduction Nickel titanium, also known as nitinol, is a metal alloy of nickel and titanium, where the two elements are present in roughly equal atomic percentages. The term nitinol is derived from its composition and its place of discovery: (Nickel Titanium Naval Ordnance Laboratory). William J. Buehler and F. Wang [1,2], discovered its properties during research at the Naval Ordnance Laboratory in 1962. Nitinol alloys exhibit two closely related and unique properties: shape memory and superelasticity (also called pseudoelasticity). As a biomedical material, Nitinol has to meet several requirements: high corrosion resistance in chloride-rich medium and biocompatibility combined with suitable mechanical properties [3]. Nitinol is used for orthodontic treatments, and in cardiovascular surgery for stents, guide wires, filters, etc., in orthopaedic surgery for various staples and rods, and in maxillofacial and reconstructive surgery [4,5]. Probably the most common medical use of nitinol is construction of stent-grafts. Another example of the use of these materials in medicine are dental applications or use in prosthesis [6,7]. Cardiovascular stents are small cylindrical devices introduced in stenosed arteries to reopen the lumen and restore blood flow. if the stent fracture, free end of the broken wire penetrates into and injures the walls of the patient's artery. There then follows a proliferation of cells and the formation of scar tissue around the injury, similar to the scarring of other organic tissues. This reaction to the trauma subjects the artery to close,this constitutes the major mechanism of in-stent restenosis. Nitinol contains about 50 % of nickel, which is a known allergen and carcinogen. Nickel release occurs in a patient's body [4], this process is associated with corrosion of the stent-grafts. In order to diminish the release of nickel from Nitinol wires in the patient's body, numerous treatments have been used. Further untreated Nitinol (without a specially prepared protective coating) is protected by a layer comprised mainly of TiO2, with a small amount of NiO in the outermost surface layers [8], but this natural protective coating is insufficient. Corosion resistence of Nitinol is a function of surface preparation, namely mechanical polishing, electropolishing, electropolishing,chemical passivation, straw-coloured or blue-coloured oxide deposition [9]. The corrosion rate is dependent on the surface preparation, being the lowest for the specimens treated by chemical passivation, and highest for the mechanically polished specimens [3,9]. A finite element analysis of stent grafts under representative cyclic loading conditions was presented in some works [10, 11], but these studies do not consider corrosion processes in the overall degradation of the stent. In this paper, a computational analysis of different stent–graft combinations and their impact on mechanical characteristics while undergoing cyclic pressure loads was carried out, employing the finite element solver ANSYS. We can assume that stents rupture first occur in regions with highest portion of deformation, but a fracture of the stent may occur due to corrosion in places where the protective layer is damaged. Protective layer can be damaged during stents manufacturing or surface layer can be damaged due to friction of individual wires of stents structure. Fig.1. Stents-graft geometry. Experimental details Forty commercially available diamond-shaped stent-grafts made from knitted wires were examined. Knitted stent grafts were made from two different Nitinol alloys. Laser cutted stent-grafts with two different types of cut geometry (diamond-shaped, peak-to-peak bridges, see Fig.1 ) were used for fatigue experiments (Fifty diamond-shaped and Fifty stents with “peak-to-peak“ bridges geometry). Laser cutting is a typically method for stent production. Most of stents in the world are cutted. Both types of laser cutted stents were made from the same stock ( NiTi tube with same diameter 25 mm and same chemical composition). Nominaly diameter of cutted stents was 12 mm and nominaly length 65 mm. Wall thickness of stock tube was 0.62 mm. These stents differ only in the cut geometry. The knited stents were made from wires with diameter 0.64 mm. Dimensions of the individual elements of all stents were approximately the same (compare diameter of wires 0.64 mm to the thickness 0.62 mm in he case of cutted stents ). The examined stent-grafts were treated with four different surface treatments. The specimens were refined by mechanical polishing and conventional ion implantation. In this which ions are extracted from plasma, accelerated, and bombarded into a device. In this study were surfaces refined with molybdenum (Mo) and carbon (C). The pulsed electron-beam modified the surface as deep as 10 nm for Mo and 30 nm for C. As example of bioactive surface was chosen surface refined with Heparin coating and passivation in HNO3. . The Laser surface melting (LSM) of Nitinol using either argon as a shielding gas was used for 20 laser cuted stents. The Laser surface melting (LSM) of Nitinol using either argon as a shielding gas was used for 20 laser cuted stents. After LSM in argon, Nitinol revealed new phases, such as Ti2Ni, TiNi3 and martensite B19. Overview of applied surfaces and types of stents is shown the Table. 1. The examined stents were strung on “artificial blood vessels”. Fatigue loading was simulated by the bending of “artificial blood vessels”.Electrochemical measurements or corrosion tests were performed in performed in a three-electrode corrosion cell. Measurements were performed in simulated physiological solution (This solution is known as Hanks salt solution.), with the following composition: 9 g/L NaCl, 0.5 g/L KCl,0.25 g/L NaH2PO4·2H2O, 0.35 g/L NaHCO3, 0.05 g/L Na2HPO4·2H2O, 0.19 g/L CaCl2·2H2O, 0.4 g/L MgCl2·6H2O, 0.06 g/L MgSO4·7H2O, 1 g/L glucose. The pH was adjusted to 5, 8 and 9 by the addition of HCl or NaOH solutions. The specimens were embedded in a teflon holder and exposed to Hanks simulated physiological solution at temperature 37.3±1 o C. Table 1. Types of stents used in the experiments Surface treatment conventional ion implantation Laser surface melting (LSM) mechanical polishing (MP) bioactive surface (BS) Geometry of atent-graft Mo O „diamondshape“ 10 10 10 10 10