## 3 Magnesium-Based Devices
### Magnesium Alloy Biocompatibility and Resorption  
A metal stent biocompatibility study reported a significant dielectric spectroscopy interaction between the metallic stents and cell membrane of erythrocytes, albeit this phenomenon was not related to unfavorable clinical outcomes [28]. For this, magnesium alloy devices were developed aiming to increase the biocompatibility between device versus blood and vessel wall. Magnesium has a high biocompatibility but low tensile force and accelerated corrosion rate (ACR) with an increased hydrogen (H\({}_{2}\)) evolution rate [9,20], which are H\({}_{2}\) pockets formations next to metal surface and the alkalization (OH\({}^{-}\)) of the surrounding solution mostly due to the ACR. Biologically, ACR delays the intimal healing process inducing local necrosis and separating the tissue layers [9,29,30]. The ACR issue solution came from magnesium alloys extensively used in aerospace since the addition of low levels of rare-earth metal elements decelerates the ACR process, thus minimizing the H\({}_{2}\) evolution rate issue. After many attempts, two alloys remained as valuable options for BRS application, the AE21 (SHL-Alubin Ltd., Kiryat Biolik, Israel) and WE43 (Magnesium Elektron Ltd., Swinton, Manchester, UK) [29]. Magnesium AE21 alloy is a composite of 97.0% of Mg, 2.0% of Aluminum (Al), and 1.0% of rare-earth metals[Cerium (Ce), Praseodymium (Pr), and Neodymium (Nd)] [8], and the WE43 alloy has 93.6% of Mg, 3.7-4.3% of Yttrium (Y), 2.4-4.4% of rare-earth metals (predominantly Neodymium, Nd), and 0.4% of Zirconium (Zr) [9; 30]. From the biologic point of view, these alloys considerably improved the predictability of the degradation process providing longer platform integrity [29] and optimizing the radial strength-to-weight ratio of metallic Magnesium BRS platforms, which are now comparable to aluminum and steel alloys with a lighter platform [29; 30; 31].  
The complete resorption time of magnesium-based platforms vary depending on magnesium alloy, platform characteristics (strut thickness and width), and the presence or not of polymeric coating. In porcine models implanted with AE21 alloy, uncoated platform device struts were observed at 56 days and an expected complete resorption time \(\leq\)89.4 days [8]. In fact, an objective resorption observation data came from a WE43 alloy uncoated platform Lekton Magic (Biotronik AG, Bulach, Switzerland) presenting a complete in-vivo resorption in \(<\)2 months [9; 32; 33]. The addition of polymeric coating in the platform like Magmaris (DREAMS 2G--Biotronik AG, Bulach, Switzerland) significantly delayed the in-vivo resorption time up to 12 months when 95% of platform struts were fully resorbed as a consequence providing effective 3 months of scaffolding vessel support time [34].  
The magnesium alloy resorption main chemical process is hydrolysis, where Mg\({}^{2+}\) and two H\({}_{2}\)O molecules are converted to magnesium hydroxide [Mg\({}^{2+}\)(OH\({}^{-}\))\({}_{2}\)] and hydrogen gas (H\({}_{2}\)) followed by a progressive replacement of magnesium by calcium (Ca\({}^{2+}\)) and phosphate (PO\({}_{4}\)\({}^{3}\)-) into a soft amorphous hydroxyapatite imprint [Ca\({}_{5}\)\({}^{2+}\)(PO\({}_{4}\)\({}^{3}\)-)\({}_{3}\)OH\({}^{-}\)]. Based on the in vitro and in-vivo observations, these processes occur simultaneously but in a multi-phased fashion. First, the initial corrosion phase starts with the hydrolytic process being homogeneously distributed over platform surface. In the second phase, the resorption process is now heterogeneous, deeply advancing to the inner parts of the struts through microfractures in the surface of the platform. Finally, in the third and last phase, the calcium and phosphate are now homogenously distributed over the imprint remains of the platform until the complete substitution of the residual Magnesium for soft amorphous hydroxyapatite (Figure 1) [21; 32].