Over recent years, there has been an exponential rise in the usage of biomaterials in medicine-based applications [1]. Healthcare professionals and commissioners have been emphasizing early diagnosis and treatment of diseases and enhancing the efficiency of the available treatment strategies. This has been due to the more active lifestyle of the population, which significantly transformed the treatment of fatal diseases and conditions. However, there exists a requirement of affordable healthcare resources for the active aging generation. The longevity of an implanted biomaterial depends on the anchorage and integration between the implant and the living bone [2]. There are various factors that determine the success of an implanted biomaterial, namely the form and structural characterization, stability, mechanical-loading, material property, location of the implanted site, and host response [3,4]. The key outcome is to finally derive a matrix that matches the bone in composition, structure, and properties [5]. Biocompatibility and mechanical strength are the essential pre-requisite characteristics of any implanted biomaterial to be employed for biomedical, orthopedic, and dental applications [6,7,8,9,10].

The inert metals, such as stainless steel, titanium, and cobalt alloys, having superior corrosion resistance and mechanical properties (concerning biological tissues) have been traditionally used as implant materials for cardiovascular and orthopedic applications [11,12,13]. Liu et al. [14] constructed a novel stromal cell-derived factor-1α (SDF-1α)/laminin-loaded nanocoating on the 316 L stainless steel (SS) surface to provide improved function in the modulation of vascular remodeling. The modified surface was found to control the delivery of biomolecules and exhibit promising potential to provide stage-adjusted treatment after injury. Furthermore, an in vitro biocompatibility study suggested that the constructed layer may effectively prevent thrombosis formation by inhibiting platelet adhesion and activation, while accelerating endothelium regeneration by inducing endothelial cell (EC) migration and endothelial progenitor cell (EPC) aggregation. An in vivo animal test further demonstrated that the nanocoating may prevent thrombus and neointimal hyperplasia after implantation for 3 months [14]. Presently, these metallic implants are mainly designed for permanent tissue replacement, but they do not account for cases where only temporary support is required. There have been numerous difficulties involved, some of which are the stress-shielding effect over a while, which then leads to bone-weakening and distortion of diagnostic images. Finally, additional surgery is required for removal of the implant when these permanent metallic implants are used [15,16]. Extensive research efforts are ongoing to develop materials that can be used as temporary implants in the human body to perform multiple functions, depending on the nature of the ailment. Among the metals, Zinc (Zn)- and Magnesium (Mg)-based materials have garnered significant attention in recent years to serve as temporary implants [17]. Magnesium and its alloys exhibit superior strength, ductility, and degradability compared to its biodegradable counterparts, Zinc and Iron [17]. The elastic modulus (∼45 GPa) of magnesium is close to that of human bone (∼10–20 GPa), and this helps to avoid stress shielding effects compared to other metallic biomaterials [18,19]. The magnesium in plasma exists as Mg²⁺ in a concentration of 1.2–1.4 mM and is excreted in urine [20,21]. Increased local degradation of Mg (and its alloys) causes increased hydrogen gas release from the implanted site, leading to an increase in pH [20]. This increase in hydrogen gas evolution and pH causes localized cell death and cytotoxic reactions, which further leads to failure [22,23,24]. For implanted devices, such as in orthopedics, where it is poorly vascularized, the accumulation of hydrogen gas is a safety concern as a gas embolism may potentially block the bloodstream, causing fatalities [25]. For coronary stents, the gas generated is of not a large concern, as it can be removed by convective transport phenomena due to a high amount of blood flow at the implantation site [26]. Nevertheless, it has been reported that the addition of Zn to Mg can significantly decrease the amount of hydrogen generated.

The favorable degradation behavior of Zn has led to the emergence and development of Zn-based alloys in recent years [22,27]. Being an essential element, Zn is involved in various biological processes, including being a metabolite for nucleic acid, signal transduction, and gene expression, and it takes part in interactions with a range of organic ligands and apoptosis coordination [28]. Zn does not corrode as quickly as Mg and the corrosion of pure Zn is known to be moderate [29]. However, Zn on its own has inferior mechanical properties with a tensile strength as low as 20 MPa, Vicker hardness of 3.7 and elongation of 0.2%. and these translate to relatively lower material strength as compared to its metallic counterparts [22]. Thus, alloying of Zn is crucial for its application in load-bearing implants. There have been several new opportunities and difficulties in the development of Mg and Zn alloys for cardiovascular applications and to overcome the drawbacks of hydrogen gas evolution [29].

2. Driving Force for Temporary Implants