Updates on the research and development of absorbable metals for biomedical applications

The three families of absorbable metals

So far, absorbable metals are basically made of iron, magnesium, or zinc as the main ingredient (base metal). They are essential elements needed for proper metabolic function of human body, and are relatively compatible with human cells and tissues as confirmed in many in vitro and in vivo studies (Drynda et al. ²⁰¹⁵ ; Myrissa et al. ²⁰¹⁶ ; Kubásek et al. ²⁰¹⁶ ; Marco et al. ²⁰¹⁷ ; Wang et al. ^2017a ; Drelich et al. ²⁰¹⁷ ). To improve both mechanical properties and corrosion behavior, the base metals are mostly alloyed with elements which are considered “non-toxic”. The first base metal, i.e., iron, has attractive mechanical properties approaching to those of the 316L stainless steels which is considered as a benchmark for metallic biomaterials. Indeed, pure iron has superior mechanical properties compared to pure magnesium and zinc (Table  ¹ ), making it a suitable candidate for implants which require a high structural strength such as coronary stents (Francis et al. ²⁰¹⁵ ; He et al. ²⁰¹⁶ ; Lin et al. ²⁰¹⁶ , ²⁰¹⁷ ). Alloying with manganese turns the ferromagnetic iron into non-magnetic as a single austenitic phase is formed (Feng et al. ²⁰¹⁶ ; Hufenbach et al. ²⁰¹⁷ ). Iron was made into composites with calcium phosphate or calcium silicate that led to enhanced bioactivity in both in vitro and in vivo experiments (Ulum et al. ²⁰¹⁵ ; Dehestani et al. ²⁰¹⁶ ; Wang et al. ^2017b ).

The second base metal, i.e., magnesium, is a lightweight metal with density of 1.74 g/cm ³ , and it has a low elastic modulus of 40–45 GPa which is near to that of the bone (Li et al. ²⁰¹⁶ , Zhao et al. ^2017b ). Therefore, it is a very attractive candidate material for bone implants (Chaya et al. ²⁰¹⁵ ; Han et al. ²⁰¹⁵ ; Yu et al. ²⁰¹⁷ ). Advanced alloying and processing techniques, such as thermomechanical treatment and severe plastic deformation, have improved the properties of magnesium-based metals (Sunil et al. ²⁰¹⁶ ; Griebel et al. ²⁰¹⁷ ) (Table  ¹ ). So far, the most complete literature coverage has been on magnesium-based metals. This has facilitated their clinical translation (Agrawal et al. ²⁰¹⁶ ; Zhao et al. ^2017b ). The third base metal, i.e., zinc, has been lately added as a new family of absorbable metals (Liu et al. ^2016b ; Wang et al. ^2016a ; Levy et al. ²⁰¹⁷ ). The cytocompatibility of zinc-based metals was studied in view of its applications for bone and vascular implants (Murni et al. ²⁰¹⁵ ; Shearier et al. ²⁰¹⁶ ; Guillory et al. ²⁰¹⁶ ). Alloying zinc with magnesium and other trace elements, such as manganese, strontium, etc., improved the mechanical properties to those levels of some magnesium alloys (Gong et al. ²⁰¹⁵ ; Mostaed et al. ²⁰¹⁶ ; Liu et al. ^2016a , ^b ) (Table  ¹ ). Thermomechanical treatment, such as extrusion (Fig.  ² a), was employed to improve both the strength and ductility of many magnesium and zinc alloys (Wang et al. ²⁰¹⁴ ; Gong et al. ²⁰¹⁵ ). The two alloys were also subjected to severe plastic deformation processing such as the equal channel angular pressing (ECAP) (Fig.  ² b) to further enhance their mechanical properties (Mostaed et al. ²⁰¹⁴ ; Dambatta et al. ²⁰¹⁷ ).

Fig. 2

Adapted with permission from John Wiley and Sons, and Elsevier (Gong et al. ²⁰¹⁵ ; Dambatta et al. ²⁰¹⁷ )

Attempts to improve the corrosion behavior

In terms of corrosion behavior, it is generally acknowledged that iron-based metals corrode slowly, magnesium-based metals corrode rapidly, and zinc-based metals corrode moderately (Li et al. ²⁰¹⁶ ; He et al. ²⁰¹⁶ ; Zhao et al. ^2017b ; Mostaed et al. ²⁰¹⁸ ). Their corrosion rates in simulated body fluid indicate this trend (Table  ¹ ). A direct comparison done by Vojtech et al. ( ²⁰¹⁵ ) showed corrosion rates of pure iron, pure zinc, and pure magnesium after 168 h of immersion in saline solution at 37 °C were 0.2, 0.6, and 4 mm/year, respectively. Many attempts have been done to control the corrosion behavior of absorbable metals, i.e., increasing corrosion rates of iron-based metals where its low rates become more evident under in vivo condition (Kraus et al. ²⁰¹⁴ ; Drynda et al. ²⁰¹⁵ ; Lin et al. ²⁰¹⁷ ) and lowering those of magnesium-based metals as it also becomes problematic under in vivo condition (Shi et al. ²⁰¹⁷ ; Cui et al. ²⁰¹⁷ ; Liu et al. ^2017a ). Various advanced metallurgical processes have been used to achieve it. Basically, it was done by optimizing the composition and microstructure of the metals, either in the bulk, e.g. precise alloying, purification and improved manufacturing process, or on the surface, e.g. surface treatment process and protective coating (Table  ² ). Attempts to accelerate the corrosion rate of iron-based metals has been done by various methods such as alloying with manganese, palladium, silver, gallium, sulfur, and intermetallics (Čapek et al. ²⁰¹⁶ ; Wang et al. ^2017a ; Hufenbach et al. ²⁰¹⁷ ; Sotoudehbagha et al. ²⁰¹⁸ ; Sikora-Jasinska et al. ²⁰¹⁸ ), vacuum plasma nitriding process (Lin et al. ²⁰¹⁶ ), implanting silver using a vapour vacuum arc technique (Huang et al. ²⁰¹⁶ ), and making composite with polymers (Yusop et al. ²⁰¹⁵ ; Cysewska et al. ²⁰¹⁷ ; Qi et al. ²⁰¹⁸ ). Slowing down the corrosion rate of magnesium-based metals has been addressed by various means such as purification from detrimental impurities (Hofstetter et al. ²⁰¹⁵ ; Qu et al. ²⁰¹⁷ ), alloying with calcium, zinc, rare earth elements (RE) and other elements (Zander and Zumdick ²⁰¹⁵ ; Mao et al. ²⁰¹⁷ ; Li et al. ²⁰¹⁸ ), and coating with bioceramics or biopolymers (Hiromoto et al. ²⁰¹⁵ ; Witecka et al. ²⁰¹⁶ ; Su et al. ²⁰¹⁶ ; Patil et al. ²⁰¹⁷ ).

In addition, using porous structure of iron was recently studied as a way to control its corrosion rate while targeting a new potential application for bone scaffolds (Heiden et al. ²⁰¹⁶ ; Feng et al. ²⁰¹⁷ ). This actually opened a new direction of using porous absorbable metals as alternative to polymers for scaffolds by exploiting mostly their superior mechanical properties over biodegradable polymers. The emerging additive manufacturing technology helps to advance the design and process of ideal topological porous metals suited for bone scaffolds and orthopaedic implants (Wang et al. ^2016c ; Gordeladze et al. ²⁰¹⁷ ). The high strength and slow corrosion of iron give a higher degree of freedom to vary its surface area/weight ratio for controlling the corrosion rate and matching the different strength and flexibility requirement for bone scaffolds (Yusop et al. ²⁰¹⁵ ; Alavi et al. ²⁰¹⁷ ; Yang et al. ²⁰¹⁸ ). Differently, the high surface area of porous structure increases the challenge of controlling the rapid corrosion of magnesium scaffolds (Aghion and Perez ²⁰¹⁴ ; Cheng et al. ²⁰¹⁶ ). As for its solid form, alloying, composite fabrication, and surface coatings are the methods used to control the corrosion of porous magnesium (Yazdimamaghani et al. ²⁰¹⁷ ). As for zinc-based metals, Zhao et al. ( ^2016d ) suggested porous zinc for low load-bearing bone scaffolds, since they observed its good corrosion resistance in simulated body fluid and its ability to induce CaP precipitation during immersion tests. The following section describes in vivo corrosion behavior of absorbable metals during implantation in animals.

In vivo corrosion behavior

Accelerating the in vivo corrosion of iron-based implants has been the focus of many works. The interestingly improved corrosion rates during in vitro corrosion tests are often not replicated in the in vivo tests. Drynda et al. ( ²⁰¹⁵ ) observed that a series of Fe–0.5Mn, Fe–2.7Mn, and Fe–6.9Mn alloys exhibited no significant corrosion even after 9 months implanted subcutaneously in mouse. The formation of barrier (Fe–Mn phosphates) layers was determined as the cause of the high corrosion resistance, so strategies to dissolve these layers or to prevent their formation need to be developed to expedite the in vivo corrosion of Fe–Mn implants. Kraus et al. ( ²⁰¹⁴ ) had similar finding when observing Fe–10 Mn–1Pd and Fe–21Mn–0.7C–1Pd pins implanted in a growing rat skeleton over a period of 1 year. They suspected that a dense layer of corrosion products acted as a barrier against oxygen transport, whereas oxygen is prerequisite for iron corrosion and its availability is rather limited in bony tissue. Thus, the corrosion of alloys should depend on hydrogen evolution, as it has a sufficiently low electrode potential, however, at a much lower rate than in presence of oxygen. More recent studies on long-term in vivo implantation of iron stents also indicated the need of increased corrosion rate. Lin et al. ( ²⁰¹⁶ ) implanted 53 μm strut drug-eluting coronary stents made of nitride iron (+zinc barrier layer and 12 μm PLLA coating) in rabbit abdominal aorta. The stents maintained adequate scaffolding (125 kPa) after 3 months of implantation while having a shortened corrosion period to 13 months. However, a complete bioresorption of the corrosion products [Fe ₃ O ₄ , FeOOH, Fe ₂ O ₃ , and Fe ₃ (PO ₄ ) ₂ was not observed (Fig.  ³ a, d)]. The same group reported another study that showed 70-μm strut nitride iron stents to lose its mass twice as much as pure iron stent after 36 months of implantation in rabbit abdominal aorta (Fig.  ³ b, c) (Lin et al. ²⁰¹⁷ ). Although the study showed signs of good long-term biocompatibility (Fig.  ³ e), the complete corrosion of the stent and the absorbance of the insoluble corrosion products may take 4–6 years long, and thus, further increase of corrosion rate is necessary. The results obtained by Qi et al. ( ²⁰¹⁸ ) showed that some PLLA-coated stents could totally corrode in the abdominal aorta of New Zealand white rabbits, whereas all the non-coated stents left some remnants of struts several months after implantation.

Fig. 3

Adapted with permission from Elsevier (Lin et al. ²⁰¹⁶ , ²⁰¹⁷ )

Although magnesium-based implants have demonstrated high potentials, fast corrosion is often observed in the in vivo studies. A recent study by Yue et al. ( ²⁰¹⁵ ) acknowledged the need for longer period of corrosion of magnesium stents which can shed light on the incidence of late thrombosis and the extent of late lumen loss. Their 28-day-long implantation of Mg–Al–Zn alloy stent (Fig.  ⁴ a) in coronary artery of mongrel dogs (Fig.  ⁴ b) proved that the stent has good plasticity and strong resistance, and it induced mild neointimal hyperplasia 2–4 weeks after stenting (Fig.  ⁴ c). However, the short corrosion time of the stent could give some side effects such as the vascular elastic recoil and late lumen loss. Besides being targeted for endovascular stents, magnesium alloys were also utilized for esophageal stents. Zhu et al. ( ²⁰¹⁶ , ²⁰¹⁷ ) successfully determined its feasibility through the implantation of silicone-covered magnesium esophageal stent (Fig.  ⁴ d) in the esophagus of rabbits. The stents showed good flexibility, elasticity, and patency (Fig.  ⁴ e), and they corroded slower than bare magnesium stents. The stents assisted the esophageal wall remodelling with desired thin epithelial and smooth muscle layers (Fig.  ⁴ f) and provided a reliable support for at least 2 weeks with acceptable migration rates without causing severe injury or tissue reaction when compared with plastic stents or collagen deposition.

Fig. 4

Adapted with permission from John Wiley and Sage Publications (Yue et al. ²⁰¹⁵ ; Zhu et al. ²⁰¹⁶ )

A different in vivo corrosion behavior was observed for zinc implants. In a long-term implantation study, Drelich et al. ( ²⁰¹⁷ ) demonstrated that chronic inflammation in relation with corrosion activity of zinc wires implanted in the murine artery was subsided between 10 and 20 months (Fig.  ⁵ a). The wires exhibited a steady corrosion rate for up to 20-month postimplantation (Fig.  ⁵ b) without causing local toxicity despite a steady build-up of passivating corrosion products and intense fibrous encapsulation of the wire. This continued corrosion indicates that zinc stents could safely corrode within a time frame of approximately 1–2 years. Yang et al. ( ²⁰¹⁷ ) reported the implantation of pure zinc stents in rabbit abdominal aorta and found that the stents conserved its mechanical integrity for 6 months and almost half of the stent volume was corroded after 12 months of implantation (Fig.  ⁵ c). They figured out that the corrosion process involved a conversion mechanism in a microenvironment that evolved from dynamic blood flow to neointimal tissue. The stent strut was covered by neointimal layer starting from the first month of implantation indicating a rapid endothelialization, but there was no intimal hyperplasia or obvious accumulation of corrosion products even after 12 months of implantation (Fig.  ⁵ d). These studies demonstrate an optimistic in vivo corrosion behavior of pure zinc.

Fig. 5

Adapted with permission from Elsevier (Drelich et al. ²⁰¹⁷ ; Yang et al. ²⁰¹⁷ )

Correlation between in vitro and in vivo corrosion

A desirable corrosion behavior is mostly observed in the in vitro setting where a full control on the testing parameters can be established. Once the metals are exposed to the in vivo environment, they often behave differently (Sanchez et al. ²⁰¹⁵ ; Bowen et al. ²⁰¹⁵ ). In the in vivo environment, corrosion is influenced by the complex body fluid content (water, organic compounds, dissolved oxygen, anions, cations, amino acids, proteins, plasma, etc.) and the reciprocal interaction with the tissue response (Johnston et al. ²⁰¹⁸ ). As the metal implantation causes injury, the body responses to it by decreasing the pH value around the implantation site (i.e., 5.3–5.6), and this in turn may accelerate corrosion process of the implant and reduce the local oxygen concentration (Paramitha et al. ²⁰¹⁷ ). In their comprehensive review about factors influencing in vitro corrosion of magnesium and its pertinence to in vivo corrosion, Johnston et al. ( ²⁰¹⁸ ) urged the inclusion of factors such as (i) proteins, (ii) amino acids, (iii) vitamins, and (iv) tissue encapsulation. At this point, there is a lack of knowledge on the best method to characterize the in vivo environment and corrosion mechanism and to correlate the results with the processing technique and properties of the metals. In addition, the correlation between the results of in vitro and in vivo corrosion experiments is still barely established (Zheng et al. ²⁰¹⁴ ; Sanchez et al. ²⁰¹⁵ ; Johnston et al. ²⁰¹⁸ ).

After systematically reviewing the results of more than 20 in vitro and in vivo corrosion studies of magnesium and its alloys, Sanchez et al. ( ²⁰¹⁵ ) found that deriving a correlation between in vitro and in vivo test results is yet challenging. This is attributed mainly to the difficulties in mimicking the complex in vivo physiological conditions with in vitro experiments and to the wide and variable testing parameters and procedures used in the studies. Myrissa et al. ( ²⁰¹⁶ ) found that corrosion rates of pure magnesium, Mg–2Ag and Mg–10Gd alloys, are comparable between in vitro and in vivo conditions only after 4 weeks of experimentation period. In their work, the magnesium pins were immersed in DMEM (+ 10% FBS at 37 °C, 20% O ₂ , 5% CO ₂ , and 95% rH) and then measured for their mass loss, whilst similar pins were also implanted in Sprague–Dawley rats which then subjected to μCT scans for corrosion rate determination using the software Mimics ^® . Although the experiment indicated that the in vivo corrosion rate can be represented by in vitro rates to some extent, a more complex in vitro set-up (including biological component, mechanical and dynamical exposure of the metals) is needed to better mimic the in vivo condition.

The effect of dynamical exposure on corrosion behavior of absorbable metals was studied in two recent works. Wang et al. ( ^2016b ) analyzed the corrosion mechanism of Mg–Zn–Ca alloy under mimetic hydrodynamic conditions using an in situ and real-time electrochemical set-up in a vascular bioreactor. With this in vitro set-up, they demonstrated that flow-induced shear stress accelerated mass and electron transfer that led to an increase in uniform and localized corrosions. Effect of flow on corrosion of magnesium was also observed by Saad et al. ( ²⁰¹⁶ ) using a test rig that mimicked the environment surrounding a cancellous bone where mass loss and mechanical integrity of the porous magnesium deteriorated linearly with an increase in porosity and degradation time. These studies confirm the influence of the in vivo condition, i.e., fluid flow, on corrosion behavior of absorbable metals. This is of a great importance to correlate the in vitro and in vivo corrosion behavior. The large variability between in vitro corrosion test methods makes comparison between studies increasingly difficult (Johnston et al. ²⁰¹⁸ ). A standard approach is urgently needed to allow their searchers to observe the exact effect of a certain parameter of in vitro corrosion, and thus, a reliable iterative approach can be taken to improve the structure and processing of the materials.

Real-time corrosion monitoring

The body responds to the corrosion of an absorbable metal implant via a cascade process of protein absorption, coagulation, acute inflammation, chronic inflammation, and foreign body response. Acute interaction occurs shortly after implantation. It induces the formation of an interspace between the implant and the tissue which then is filled with a large volume of body fluid. This is followed by cellular inflammation and then proliferation to finally fill the interspace (Kim et al. ²⁰¹⁶ ). The body-response phenomena have been successfully evaluated using the tools adopted from those used for inert biomaterials. The common tools for assessing absorbable metal implants are radiography, ultrasonography, micro-computed tomography (μCT) and its advanced synchrotron version (SRμCT), magnetic resonance imaging (MRI), blood evaluation, and histological and implant retrieval analysis (Paramitha et al. ²⁰¹⁷ ). However, these methods provide segmental, discontinued and offline results, so they are mostly unable to reveal the fundamentals of in vivo corrosion of the metals which is a continuous and time-dependent process. The nature of corrosion process of absorbable metals that produces metal and hydroxyl ions, hydrogen gas (for magnesium), and a flow of electric current should be exploited as the basis for developing in vivo corrosion assessment tools.

Over the past 5 years, several reports on the continuous assessment of corrosion behavior of absorbable metals have been published. Wang et al. ( ^2016b ) combined a vascular bioreactor that circulates simulated body fluid with an electrochemical cell that is connected to a potentiostat (Fig.  ⁶ a). With this set-up, they analyzed the in situ and real-time electrochemical corrosion mechanism of magnesium alloy under the influence of dynamic flow as previously described. Lately, Natasha et al. ( ²⁰¹⁸ ) developed an online monitoring system which consisted of a microdialysis probe (as a tool to sample the fluid adjacent to magnesium implant) and a fabric-based electrochemical device (FED) (as a catalytic biosensor specific to Mg ²⁺ ), both of which were connected to a potentiostat (Fig.  ⁶ b). The device demonstrated a pseudo-linear response of concentration vs. time from 0.005 to 0.1 mmol/L with a slope of 67.48 μA/mmol L. It also showed high ion selectivity with detectable interfering species less than 1%, a high temporal resolution, and a reduced sampling time (as it required only 3 μL of fluid sample to complete a measurement). This system could be further developed as a potential tool for real-time assessment of the in vivo corrosion behavior of magnesium implant.

Fig. 6

Adapted with permission from Elsevier and Springer Nature (Wang et al. ^2016b ; Zhao et al. ^2016b , ^c ; Natasha et al. ²⁰¹⁸ )

Zhao et al. ( ^2016a ) developed a set-up composed of hydrogen gas sensor and capillary pH and Mg ²⁺ microsensors to measure the real-time concentration of magnesium ion, hydroxyl ion, and hydrogen gas. By using the set-up, they were able to generate a map of hydrogen concentration in the vicinity of magnesium alloy sample. The set-up was further developed to a transdermally (non-invasive) electrochemical hydrogen microsensor (Fig.  ⁶ c) and tested for measuring in vivo corrosion of magnesium implants in mouse (Fig.  ⁶ d) (Zhao et al. ^2016c ). Using the sensor, they realized that hydrogen permeated through the skin at the concentration as low as 30–400 μM with a fast response time of 30 s. Although the hydrogen levels permeated through the skin were very low, the sensor changed its color to give a 3D visualization of hydrogen permeation (Fig.  ⁶ e) (Zhao et al. ^2016b ). When the transdermal hydrogen sensor measurement results were combined with those of ICP-MS and XPS, a more comprehensive understanding of in vivo corrosion behavior of magnesium was obtained. It was realized that the impurities were among the determinant factors responsible for rapid in vivo corrosion of magnesium (Zhao et al. ^2017a ). However, since it is a non-invasive sensor which depends on the hydrogen permeation through the skin, it may not be practical for monitoring in vivo corrosion of magnesium implants placed deeper beneath the skin, i.e., bone screw and plate implanted under the muscle.

The available in vitro corrosion data of absorbable metals could not be directly compared, as the test condition varies from one report to another. This points out the importance of using some standardized methods or common protocols that allow a comparison between results obtained in different labs. Many in vivo parameters are not yet mimicked in the in vitro corrosion testing, and thus, a correlation between corrosion rates obtained in vitro with those measured in vivo is difficult to establish. Thus, studies leading to a better understanding of in vivo corrosion behavior of absorbable metals should become a focus of the future works. In situ and real-time corrosion monitoring techniques based on the electrochemical nature of corrosion, e.g., using a minimally invasive sampling probe which allows a deeper sampling penetration adjacent to the implant, are worth further attentions. Aside from the major works on iron- and magnesium-based metals, there is a rapid increase of reports on zinc-based metals where pure zinc has been rapidly becoming an object of in vivo implantation studies, and some suggested its potential application for vascular stents. However, more basic research, i.e., on in vitro corrosion and cytocompatibility, is still needed to confirm the suitability of zinc-based metals especially its alloys for absorbable metals. The low mechanical properties of pure zinc were improved by alloying, yet the in vivo corrosion of zinc alloys is still unclear (Mostaed et al. ²⁰¹⁶ ). For the iron-based metals, considering their very low corrosion rates and the phenomenon of barrier layer of corrosion product, its suitability for large temporary bone implants such as pins and screws appears questionable. Its usage for fine-structured implants such as stent or scaffold is, however, promising. As indicated in the two PLLA-coated stent studies (Lin et al. ²⁰¹⁶ ; Qi et al. ²⁰¹⁸ ) and the PLGA-infiltrated iron scaffold (Yusop et al. ²⁰¹⁵ ), polymer coating with adjustable composition could become a way to regulate the in vivo corrosion of fine-structured iron-based implants. The in vivo studies on magnesium-based metals have acknowledged the limitation related to the fast corrosion of magnesium which only allowed an effective mechanical support for shorter period of time than is desired, so further research to delay the corrosion is mandatory. The following section describes the clinical studies of magnesium-based implants and the attempts to tailor their corrosion behavior.

Magnesium coronary stents

Coronary stents made of Mg–RE alloy have been delivered by Biotronik (Berlin, Germany), and they have been scrutinized in a continuous set of clinical studies since 10 years ago. To slow down its corrosion, the stent was coated with PLGA containing paclitaxel drug. Therefore, it was named as the drug-eluting absorbable metal scaffold (DREAMS), and then, the stent was subjected to the first-in-man trial called BIOSOLVE-I at five European centres (Bartosch et al. ²⁰¹⁵ ). The DREAMS was viewed as a good alternative to polymeric absorbable scaffolds, as, in 3-year follow-up, the overall long-term outcome was excellent, i.e., no cardiac death or scaffold thrombosis occurred (Haude et al. ^2016a ). The angiographic performance measures of DREAMS were still considered inferior to those of contemporary drug-eluting stents, and thus, the second generation device (DREAMS-2G) was developed (Fig.  ⁷ a). It was then assessed in the BIOSOLVE-II study to determine its safety and performance in symptomatic patients with de-novo coronary artery lesions. In 6-month follow-up, the DREAMS-2G showed a substantially improved performance with a superior safety profile to that of its DREAMS-1G precursor, i.e., no thrombosis or cardiac death up to 6 months (Haude et al. ^2016b ). This desirable profile was continuously observed up to 12 months with stable angiographic parameters (Haude et al. ^2016c ). Additional evidence on the excellent safety profile of DREAMS-2G was provided by the 24 months of clinical outcomes after implantation where the definite or probable scaffold thrombosis was absence (Fig.  ⁷ b) (Haude et al. ²⁰¹⁷ ). With these encouraging clinical results, the company landed CE Mark approval in 2016 for their absorbable magnesium stent named Magmaris. The stents will be soon available in the wider market, and it will offer the benefits of temporary arterial scaffolding for the patients.

Fig. 7

Adapted with permission from Europa Digital and Publishing (Haude et al. ²⁰¹⁷ )

Magnesium bone screws

Magnesium bone implants, such as bone screws, achieved a relatively rapid clinical translation and commercialization (Fig.  ⁸ a). Many in vitro and in vivo studies demonstrated the promising applicability of magnesium alloys for bone implants. The clinical studies were conducted and reported by at least two groups in Germany and Korea and a few groups in China (Zhao et al. ^2017b ). In Germany, Plaass et al. ( ²⁰¹⁸ ) reported a clinical study, wherein the Mg–Y–RE–Zr screw was compared with the standard titanium screw for fixation of a modified distal metatarsal osteotomy in 26 patients with a symptomatic hallux valgus (Fig.  ⁸ b). The 3-year postoperative MRI demonstrated a significant improvement for all tested clinical scores, i.e., AOFAS, SF-36, FAAM, and Pain-NRS, from preoperative to postoperative investigation. As no statistically relevant difference was found between the groups, the clinical results of the magnesium screw are comparable to those of the standard titanium screw. Less artifacts in the MRI was shown for magnesium implant without any implant related cysts, and the implant was corroding 3 years postoperatively. In a larger prospective clinical study involving 44 patients, they showed the feasibility of using the screws in hallux valgus surgery and advised surgeons to be aware of the proper handling of these implants and to know about corrosion effects during healing and its radiographic appearance (Plaass et al. ²⁰¹⁶ ). These magnesium screws are now commercialized by Syntellix (Hanover, Germany) under the name Magnezix after receiving the CE Mark approval in 2017 (Biber et al. ²⁰¹⁷ ).

Fig. 8

Adapted with permission from National Academy of Sciences and Elsevier (Lee et al. ²⁰¹⁶ ; Plaass et al. ²⁰¹⁸ )

Lee et al. ( ²⁰¹⁶ ) reported the results of a long-term clinical study of Mg–5Ca–1Zn alloy screws in 53 distal radius fracture fixation cases performed at Ajou University Hospital (Suwon, Korea). Two conventional stainless steel pins were also used to fix a scaphoid fracture. Cortical continuity was observed after 6 months, while the diameter of the screw was significantly reduced (Fig.  ⁸ c). After 12 months of implantation, the distal radius fracture was completely healed, while the remaining screw was nearly impossible to differentiate from new surrounding bones. The patients did not feel any discomfort or pain throughout the entire study. They returned to their everyday life without any sign of pain or decrease in range of motion, and they gained back a normal range of grip power. The controlled corrosion of the alloy results in the formation of biomimicking calcification matrix at the corroding interface and initiation of the bone formation process. This facilitates the early bone healing process and leads to the complete replacement of the implant by the new bone tissue. After being approved by the Korean FDA in 2015, the Mg–5Ca–1Zn alloy screws are commercialized by U&i Corp. (Gyeonggi-do, Korea), and they are now available in the market as K-MET bioresorbable bone screws.

As commercial products, the exact materials’ composition and fabrication process of the aforementioned absorbable magnesium implants are not publicly available in scientific publications. However, one may have a look at patent databases such as that of the United States Patent and Trademark Office (USPTO). The following examples are three US patents assigned to the three companies who already commercialized the absorbable metals technology. First, the US Patent #US9700652B2 was assigned to Biotronik claiming an absorbable medical implant made of fiber-reinforced magnesium or fiber-reinforced magnesium alloys (Loeffler et al. ²⁰¹⁷ ). It is comprised of a matrix of crystalline magnesium or magnesium alloy reinforced by amorphous or nanocrystalline fibers made of Mg-(10–40 wt%) Zn-(0–20 wt%) X , wherein X is selected from the group of lanthanum, yttrium, silicon, aluminum, and calcium. This composite-like structure allows an increased strength and tunable corrosion behavior which are desirable for coronary stent applications. Second, Syntellix assigned the US Patent # US9402669B2 which claims a method for producing a medical implant, such as bone screw, nail, pin, plate, and suture anchor, from a magnesium alloy which contains 2.5–5 wt% RE, 1.5–5 wt% yttrium, 0.1–2.5 wt% zirconium, and 0.01–0.8 wt% zinc (Neubert and Schavan ²⁰¹⁶ ). The method consists of melting the alloy at 700–900 °C, atomizing the molten alloy into alloy powder under a protective-gas atmosphere, shaping the powder into a green body, extruding it at 300–400 °C to obtain a molded part, and finally producing the medical implant from it. Third, the US Patent US20170119922A1 was assigned to U&I claiming a magnesium alloy for forming a medical implant, comprising 0–23 wt % calcium and 0–10 wt % zinc (Koo et al. ²⁰¹⁷ ). The alloy possesses controlled corrosion resistance properties due to the electrical potential difference between magnesium phase and Mg ₂ Ca phase which is not greater than 200 mV as measured in a biomimetic solution specified in the patent. As can be seen, the inventions claimed in those patents based on the material composition that leads to a controlled corrosion resistance, the composite structure that leads to an improved strength, and the processing method that results into a molded semi-finished material.