Biodegradable metals (BMs), which progressively degrade in the physiological environment of the human body in fulfillment of the goal of providing temporary fixation and assisting tissue regeneration, are recognized as modern-generation metallic biomaterials for orthopedic and cardiovascular stent applications [1, 2]. BMs may exhibit promising mechanical strength and ductility, suitable degradability, and satisfying biocompatibility [3, 4]. Within the three recognized classes of BMs based on magnesium (Mg), iron (Fe), and Zn, Mg-based BMs have been broadly investigated for biodegradable implant applications. This is primarily because of their analogous mechanical properties to those of cortical bone and their acceptable biocompatibility [5, 6]. Also, Mg is an essential nutrient and plays many biological roles in the body [7]. However, Mg-based BM devices corrode very rapidly and produce undesired hydrogen (H2) gas upon corrosion which increases the localized pH and compromises the mechanical integrity of implants at an early stage (within 1–4 months) [8, 9]. Therefore, several alloying elements including aluminum (Al) and rare earth elements (REE) have been used in commercial Mg alloys, such as AZ and WE series Mg alloys, to tailor their degradation rate (DR). These alloying elements have shown promising effects in improving the corrosion and mechanical properties of Mg-based BMs, but their corrosion products (CPs) are not necessarily biocompatible [10, 11].

On the contrary, Fe-based BMs exhibit a yield strength (σYS) of 250–950 MPa, ultimate tensile strength (σUTS) of 300–1550 MPa, elongation (ε) of 2.0–19.5%, as well as reasonable biocompatibility and hemocompatibility [12, 13]. Also, similar to Mg it is an indispensable nutrient element in the body and plays a significant role in various biochemical processes, e.g., electron transfer, sensing oxygen and transport, and catalysis [14]. In addition to this, Fe-based BMs do not generate H2 gas during their degradation in the physiological medium [15]. However, Fe-based BMs degrade at the slowest DR when compared to Mg- and Zn-based BMs. In general, this class of metals has a DR of less than 0.1 mm/y, which is notably lower than the clinical requirements for BMs, which should be between 0.2 and 0.5 mm/y [16, 17]. Due to the slow DR, Fe-based alloy implants may take a relatively long time (∼ 2 years) to degrade in the body, causing complications similar to those seen with permanent metallic implants (such as stress shielding and adverse tissue reaction) [17]. Additionally, their degradation products (DPs) are not excreted satisfactorily from the body and, hence, DPs are accumulated in neighboring tissues and sensitive organs for a long time, causing cells to integrate around and within the actual footprint of the degrading implant [18].

Zn and its alloys have emerged as alternative BMs because of their moderate DRs (between those of Mg- and Fe-based BMs, DR = 0.1–0.3 mm/y), adequate biocompatibility, and effective antibacterial properties [19, 20]. Also, their DPs are completely bioresorbable and do not form excessive H2 gas [21]. Moreover, ionic Zn is also considered an essential nutrient element in the body [22] and performs several fundamental physiological and biological functions including in stimulation of new bone formation, nucleic acid metabolism, safeguarding bone mass, signal transduction, gene expression, and apoptosis regulation [23, 24]. Importantly, more than 500 enzymes depend on Zn2+ for their proper orientation and function [25]. Based on these advantages, Zn-based BMs are considered substitutes for Mg- and Fe-based BMs for bone fixation and cardiovascular stent applications [26]. However, previous studies have indicated that as-cast pure Zn exhibits poor mechanical properties (σYS = 10–29 MPa, σUTS = 18–34 MPa, ε = 0.2–1.2%, and hardness = 30–37 HV) [27, 28], which impedes the utilization of pure Zn for load-bearing implant devices. Therefore, it is vital to enhance the mechanical properties of Zn-based metallic materials via metallurgical approaches to make them suitable materials for biodegradable implant applications.

Recently, it has been proposed that metal matrix composites (MMCs) containing suitable alloying elements and reinforcing particulates can show the required mechanical properties for metallic biomaterials [13, 29]. A number of alloying elements including silver (Ag) [30], Al [31], calcium (Ca) [32], copper (Cu) [33], Fe [34], germanium (Ge) [35], lithium (Li) [36], Mg [37], manganese (Mn) [38], strontium (Sr) [39], titanium (Ti) [40], and zirconium (Zr) [41] have been investigated for alloying in Zn-based BMs. Similarly, Zn matrix composites (ZMCs) reinforced with ceramic particulates including alumina (Al2O3) [42], hydroxyapatite (HA) [43], β-tricalcium phosphate (β-TCP) [44], tungsten carbide (WC) [45], and silicon carbide (SiC) [46] have also been evaluated for biodegradable implant applications. Karimzadeh et al. [42] reported an increase in hardness and wear resistance through the addition of 32 wt.% Al2O3 nanoparticles to pure Zn. However, Al can cause health issues during long-term implantation such as senile dementia [47]. Yang et al. [43] reported that the addition of HA particles can adjust the corrosion rates of ZMCs and improve their biocompatibility. In another study, the addition of 1 vol.% β-TCP to a Zn-1Mg alloy significantly enhanced the σYS, σUTS, and ε, while maintaining a stable DR in vitro [48]. Gao et al. [46] reported that the addition of 2 wt.% SiC to a Zn matrix significantly increased the σCYS (441%) and microhardness (MH) (78%) of the ZMCs. Similarly, Guan et al. [49] found that the addition of 8 vol.% WC nanoparticles to Zn-2Fe matrices concurrently improved the tensile strength and ductility while maintaining a reasonable corrosion rate (CR), but in both cases the obtained σYS was found to be still lower than the benchmark value (σYS ≥ 200 MPa) for load-bearing implant applications.

Graphene-based materials (GBMs), such as graphene nano-sheets (GNS), GNPs, graphene oxide (GO), reduced graphene oxide (rGO) etc., are also considered promising reinforcement materials to improve the mechanical properties, corrosion performance, and wear resistance of MMCs [50, 51]. Among these carbon-based reinforcements, GNPs possess extraordinarily high surface areas (250–750 m2/g), ultrahigh mechanical strength, and chemical stability at elevated temperatures [52]. For example, single-layer graphene possesses strong covalent bonding between adjacent sp2 carbons which gives it extraordinary mechanical strength of 130 GPa [53]. However, the main concern is their biocompatibility and biodegradability. Majority of studies revealed that at high concentrations, the GBMs adversely affect cell viability (CV) but concentrations below 5–10 μg/mL do not induce a significant loss in CV [54, 55]. Akhavan et al. [54] assessed the size-dependent cytotoxicity and genotoxicity of rGO nanoplatelets and GO sheets using human mesenchymal stem cells (hMSCs) and concluded that the threshold concentration of GO nanoplatelets with average lateral dimension of ∼11 nm is 1.0 μg/mL to have a strong potential to cause cell destruction, while it is only at a high concentration of 100 μg/mL for the GO sheets with average lateral dimensions of 3.8 ± 0.4 μm to exhibit a considerable cytotoxic effect after 1 h exposure time. Wojtoniszak et al. [55] also investigated the cytocompatibility of GO and rGO using mice fibroblast cells (L929) and reported that both materials possess satisfying cytocompatibility when the concentration is between 3.125 μg/mL and 12.5 μg/mL and the cytotoxicity increases with increasing concentration. Talukdar et al. [56] investigated the cytotoxicity of graphene nano-onions, GO nanoribbons, and GO nanoplatelets using hMSCs and concluded that the cell viability of these graphene nanostructures is dose dependent (not time dependent), and a threshold concentration of 50 μg/mL is considered safe. The addition of up to 1.0 wt.% graphene nanosheets in hydroxyapatite as a composite coating showed enhanced biocompatibility toward human osteoblast cells because their filopodia inclined to move towards and get anchored by the graphene nanosheets [57]. Alumina reinforced with 10–15 vol.% GNP showed no cytotoxicity and more robust cell proliferation with higher GNP contents toward human osteoblasts [58]. In another study, addition of 0.5 and 1 wt.% GNPs in Mg composite did not show cytotoxicity toward human osteoblast-like cells (MG63), although addition of 2 wt.% GNPs in Mg showed some cytotoxic effect [59].

However, the degradation of the GNPs in ZMCs starts with detachment from the Zn matrix and continues with the dispersion in the body fluid [60]. The dispersed carbon particles can be completely metabolized in blood by human myeloperoxidase, eosinophil peroxidase, lactoperoxidase, and xanthine oxidase, because the particles will be pushed to the catalytic site following interaction of its carboxyl groups with amino acids in enzyme [61]. Also, graphene particles with a high concentration of functional groups can easily adhere to the cell membrane and be taken up by cells [62].

GNPs have been used as reinforcing materials for various metal matrices including Al [63], Cu [64], Mg [65], Ni [66], and Ti [67]. These studies have highlighted the effects of processing techniques on the resultant mechanical, corrosion, and wear properties of these GNP-reinforced MMCs. The processing techniques that have been used in these studies for the fabrication of these MMCs include stir casting, pressure infiltration, powder metallurgy (PM), laser melting deposition, electro-co-deposition, spark plasma sintering, and friction stir processing [50, 66, 68, 69]. Among these processing techniques, PM manufacturing has been widely employed to achieve uniform dispersion of reinforcement particulates in metal matrices [29]. Additionally, the MMCs fabricated by PM displayed higher strength owing to the hardening effect of the metal matrices during the ball-milling process [69]. Despite the progress in PM processing of various graphene (Gr) reinforced Al-, Mg-, and Ti-based MMCs, Gr-reinforced ZMCs are rarely reported in the literature. Owhal et al. [70] fabricated Gr-reinforced ZMCs via a two-step process, i.e., electrodeposition followed by PM, and results showed that the addition of Gr in Zn matrix remarkably improved the σYS (49 to 74 MPa) and σUTS (66 to 81 MPa), while significantly reduced the ε (12.2 to 6.0%). Yang et al. [60] prepared rGO-reinforced ZMC scaffolds using a laser additive manufacturing technique. They found a homogenous dispersion of rGO in the Zn matrix and grain boundaries free of any intermetallic compounds. Improved mechanical strength and ductility of Zn-rGO scaffolds were reported. In terms of the strength, ductility, DR, and biocompatibility, the optimal rGO content was found to be 0.2 wt.%. It is noticeable that in both composites the mechanical properties are far below than the desired properties for implant applications.

In view of such a background, biodegradable ZMCs were reinforced with 0.1–0.4 wt.% GNP via PM technique. After fabrication of GNP-reinforced ZMCs, the microstructures, mechanical and corrosion properties, and cytocompatibility of these composites were carefully and systematically investigated and compared with pure Zn sample. The optimal processing parameters for fabricating GNP-reinforced ZMCs via PM were also recommended.