Balloon-expandable stents are manufactured in the diameter that fits in the catheter sheath and expanded to the diameter of the blood vessel by inflating a balloon. Thus, the stents are plastically deformed. Figure 11.4 shows how the hoop stress in stents (lower part of the diagram) and in the vessel (upper part of the diagram) changes on deployment. As it is expanded from the initial diameter (4 mm in the figure), the stents experience the tensile hoop stress up to point b. As the stent touches the vessel of 7 mm diameter at point a, the vessel now experiences the tensile hoop stress which increases to point b. After deflating the balloon, the compressive hoop stress is applied to stents, and the stents recoil to the diameter of 8.25 mm, which corresponds to the diameter at which the stress on the blood vessel and that on the stent balance (point c).

A variety of metallic materials have been used for balloon-expandable stents as summarized in Table 11.1. Most widely used materials for balloon-expandable stents are 316L stainless steel and Co-Cr alloy. SUS316L is the most commonly used metallic material in the biomedical field. The formation of stable Cr₂O₃ passive film leads to better corrosion resistance. The tensile properties can be varied widely by processing conditions.

In the early 2000s, there are a number of reports on the relation between the restenosis and strut thickness of a stent. Several clinical reports [4, 5] indicated that a reduced strut thickness lowers the restenosis rates as shown in Fig. 11.5. These reports drove the demands for stronger and more rigid materials. Co-Cr alloy has a higher strength and better radiopacity than 316L. It also has good corrosion resistance, has good fatigue resistance, and is nonmagnetic. The alloy has been used for medical implants since 1937. The first stent using Co-20Cr-35Ni-10Mo was developed by Medtronic as “Driver^TM” in 2003. This alloy was originally developed as a heat-resistant alloy known as Elgiloy^TM. The strut thickness was reduced to 80–90 μm compared to 130–140 μm of earlier SUS316L stents.

Strong demands for materials with a higher strength, rigidity, and radiopacity led to a new alloy specifically designed for coronary stents. Boston Scientific developed Pt-Cr steel, Fe – (32.5–33.5 wt%)Pt – (17.5–18.5 wt%)Cr – (2.43–2.83 wt%)Ni [6, 7]. To attain higher radiopacity, a part of Fe and Ni was substituted by Pt. The amount of Pt was designed to give a good balance between strength, ductility, and radiopacity. The alloy was used for OMEGA^TM stents which have a strut thickness of 81 μm and width of 91 μm.

Self-expandable stents are manufactured to be the size of vessel diameter, or slightly larger, and are crimped and constrained in a sheath on a tip of catheter. When it is delivered to the intended site, the sheath is retracted and the stent deploys by itself. Most often, this self-deployment is achieved by superelasticity of shape memory alloy as shown in Fig. 11.6 [8]. Nitinol (TiNi alloy) is almost the only material used for self-expandable stents, due to its superior superelastic properties, good corrosion resistance, and biocompatibility. The basic aspect of nitinol and superelasticity can be found in Chap. 10 or elsewhere [9]. Superelasticity is a kind of pseudoelasticity induced by stress-induced martensitic transformation. As shown in Fig. 11.6, the stress-strain curve is characterized by a stress plateau and large hysteresis. A superelastic stent is in the parent phase when in an unconstrained condition and in the martensite phase when constrained, i.e., inside the sheath. A nitinol stent takes an advantage of the stress plateau and hysteresis. The stents can apply relatively gentle force corresponding to a lower stress plateau to the vessel wall for an extended range of strain. Meanwhile, when an external force is applied to deform or buckle the deployed stent, it resists with the higher stress corresponding to the upper plateau stress. This can be an advantage for the stent which often experiences severe bending or flattering, and thus, the self-expanding stents are most widely used in peripheral as well as biliary arteries. Biliary stents require maneuverability through loops, curvatures, and angulated anatomies and thus high flexibility. A self-expandable stent is also important for carotid and neurovascular vessels since a balloon-expandable stent cannot be used because of a high pressure required for the expansion, which may damage the thin wall of the brain vessels.

For bare-metal coronary stents, a clinical report indicated about 20 % of cases ended in restenosis and thrombosis that arise most typically ~30 days after stenting [10]. To minimize this reaction, most of the current coronary stents are drug-eluting stents (DES). The surface of a metallic stent is coated with non-degradable/degradable polymers containing an immunosuppressant or anticancer drug. The first-generation DESs were “Cypher” by Johnson & Johnson and “Taxus” by Boston Scientific. Both used 316L stent as a platform. Most of the second-generation DESs use Co-Cr (Abbott “Xience,” Medtronic “Resolute Integrity”) or Pt-Cr (Boston Scientific “Promus”) stents as the platform. The use of thin strut stents leads to early reendothelialization and is an advantage also for DES [10].