## 1 Introduction  
Bioresorbable coronary artery stents are clinically desired to provide temporary scaffolding to narrowed blood vessels, thereby allowing the restoration of biological functions to the blood vessels, wherein the stents can be absorbed after vessel restoration [1]. Previous studies on coronary artery stents made of bioresorbable polymers showed promising results from the clinical trials [2]. The bioresorbable poly(L-lactide) (PLLA) vascular scaffold system became the first type of bioresorbable coronary artery stents approved by the U.S. Food and Drug Administration (FDA) in 2016, with proven efficacy in treating narrowed coronary artery lesions that were 2.5-3.75 mm in reference diameter and less than 24 mm in length [3]. However, the low elastic modulus and low strength of the bioresorbable polymers resulted in stent dimensions that were greater than those of the conventional bare metal stents, and this reduced navigability and deliverability, whilst it increased flow disturbances and delayed re-endothelialization [4]. In addition, the acidic degradation products of the lactic acid polymers could trigger inflammatory reactions [2; 5]. In view of the desired mechanical performance of a coronary artery stent, bioresorbable magnesium alloys are considered more attractive, because magnesium alloys have a higher elastic modulus and a higher strength than polymers. Moreover, the electronegative charge during the degradation process provides the magnesium alloys with potential antithrombotic properties [6; 7]. Furthermore, the non-ferromagnetism of magnesium alloys will not cause artifacts when a magnetic resonance imaging (MRI) system is used to guide stent deployment [8]. Among the desirable mechanical requirements of a coronary artery stent are: (i) high radial strength, (ii) low elastic radial and longitudinal recoil, (iii) high flexibility, and (iv) high fatigue strength or long fatigue life. Optimization of the stent geometric design can significantly contribute to the improvement of the mechanical performance of stents [9], as well as to compensate for the stents limitations due to the intrinsic mechanical properties of a given material. In the geometric design of a magnesium alloy stent (MAS) or in the modification of the existing design of a conventional stainless-steel stent, two distinctive features of magnesium alloys should be taken into consideration. First, the majority of magnesium alloys have low ductility at normal body temperature, and thus, the large deformation that the stent experiences during crimping and deployment must be minimized [10]. Second, the elastic modulus and strength of magnesium alloys are higher than those of polymers but are still lower than those of stainless-steel. Thus, the stent geometric redesign should enable the magnesium alloy stent to have an adequate radial scaffolding capacity and low elastic radial recoil. Several studies which focused on the evaluation of the mechanical performance of MASs and the optimization of MAS design [10; 11; 12] have been conducted. [PERSON] et al. [10] applied a shape optimization method to 2-D magnesium alloy stent models. The optimized designs were used to create 3-D stent structures. Improved safety properties and scaffolding ability, as well as extended degradation time, were achieved. [PERSON] et al. [11] took the effect of biocorrosion on the mechanical performance of the MASs into consideration in searching for the optimal stent geometric parameters. The authors studied the long-term radial stiffness and maximum radial forces of the MASs undergoing uniform corrosion. The results showed that the rate of strength loss during biodegradation mostly depended on the hinge circumferential width, rather than on the material properties of the alloy or the cross-section shape of the stent. Thus, the mechanical properties of a magnesium alloy stent before biocorrosion could give a reasonable indication of its post-operative mechanical performance. [PERSON] et al.