## 4 Discussion
Second, the displacement at fracture during the stretching tests was related to the deformability of the real stent during expansion. The larger the displacement, the better the deformability during expansion. As stent recoil is attributed mostly to elastic deformation, the transition point in the stretching curve would be related to the radial recoil of the tubular stent after expansion. Based on the understanding of the deformation behavior of the planar specimens, the overall mechanical characteristics of the tubular stents, including crimping behavior, expanding behavior, radial strength, and flexibility, were investigated, which was aided by FEA to reveal the stress and strain distributions. The results obtained from the experimental tests and FEA were compared in terms of the amount of measurable deformation and the deformation trend of the stents. With reliable material models and rational boundary conditions, FEA could deliver results that agreed with experimental measurements to differentiate the stents of different designs. When crimping to the same outside diameter was applied, Stent A experienced the greatest plastic strain, whilst stent B had the least plastic strain and Stent C had a strain that was in between. Strain concentration readily occurred in Stent A, resulting in its earlier failure. Expansion experiments and FEA showed that Pattern B had a better expansion capacity than Pattern C. Fracture occurred at the spots that were exactly the same as those in the planar stretching tests (see Figure 17). Thus, it confirmed that the stretching tests of the planar stent mesh could indeed embody the expansion capacity of the corresponding tubular stent of a given design pattern and indicate the weakest spots in the stent. Obviously, the present design of Stent C should be improved in relation to its expansion capacity. Regarding the radial strength, Stent C was superior to Stent B. Stents B and C were comparable in flexibility, whilst being better than Stent A. The study emphasized the importance of stent design in relation to the mechanical characteristics of the material that was chosen for the stent. Most of the stents currently in clinical use are made of non-biodegradable metallic materials. In the efforts to develop bioresorbable stents made of magnesium alloys, one needs to make use of the existing stent designs and modify the designs with the poor deformability at normal body temperature and the bioresorbable nature of magnesium alloys in mind. Iterative design based on the feedback from experimental tests and FEA is needed before the magnesium alloy stents are to be assessed in vitro and in vivo in biomechanical performance for potential clinical applications. One of the limitations of the present work concerned the limited capability of the DIC method employed to measure the strains of the laser-slit stent structures, leading to large deviations of the measured strains at the edges of the structures. In addition, the difficulties of spraying a speckle pattern with a small size resulted in reduced calculation accuracy. Consequently, quantitative strain analysis could not be reliably performed. Nevertheless, strain distributions could be qualitatively analyzed, and the fracture process could be tracked, which gave insights into the local deformation and fracture phenomena of the stent patterns. Another limitation of the present work concerned the sizes of the materials for making the stent patterns. Vascular stents with wall thickness values varying over a wide range have been reported. The first commercially available ACS MultiLink stent made of 316L stainless-steel (Guidant Corp., Indianapolis, IN, US), for example, had a thickness of 50 \(\upmu\)m [24]. Bioresorbable polymer and magnesium alloy stents usually need to have larger strut sizes, due to a low elastic modulus and a low strength. The magnesium alloy stent DREAMS 2G and PLLA stent Absorb GT1, for example, both have a thickness of 150 \(\upmu\)m [25].