Magnesium Alloys Processed by Severe Plastic Deformation (SPD) for Biomedical Applications: An Overview

Abstract

Ultra-fine grained and even nanostructured magnesium alloys obtained by processing with methods of severe plastic deformation (SPD) are promising biomaterials for absorbable orthopaedic implants due to their enhanced mechanical properties, adequate corrosion resistance and biocompatibility. This paper presents an overview of the impact of the most important SPD methods – equal-channel angular pressing (ECAP) and high-pressure torsion (HPT) – on microstructure refinement and improvement of the mechanical properties of magnesium alloys intended for medical implants. Several selected groups of magnesium alloys which have the potential for use as bioabsorbable implants are discussed. The presented results of many years of research indicate the beneficial effect of SPD methods on obtaining ultra-fine and even nanostructures of magnesium alloys with improved mechanical and better functional properties, which are necessary for bioabsorbable implants.

1. Introduction

Musculoskeletal problems affect millions of people worldwide, including bone and joint pathologies.¹⁾ Moreover, bone fractures are a frequent results of various types of accidents every year. Some of them are complex and require the use of implants in the form of plates, screws, nails, wires, pins, rods, micro-clips etc. to hold the bone in place. Currently, in such cases, permanent implants made of stainless steels, Co–Cr alloys or titanium alloys are used because of their biocompatibility and corrosion resistance, high strength and toughness. However, they must be removed after a certain period of treatment.

Permanent implants can be substituted by new-generation biodegradable implants intended to work as support for fractured bone until it heals, which then will be absorbed by the body at a controlled rate.^2,3) Using biodegradable implants in osteosynthesis will bring several benefits to the healthcare system, first and foremost, increasing the comfort of patients and reducing the cost of treatment. The treated patient will not be exposed to the risks of a second surgery for removal or replacement of the implant and further morbidity. Moreover, it is particularly advantageous for paediatric cases when the bone is still in the process of growth.

Magnesium alloys are the most important biodegradable materials for orthopaedic applications and cardiovascular implants.⁴⁾ They have great potential to replace permanent metallic implants and revolutionize the implant market. Research on them has intensified in the last two decades,⁵⁾ however, the history of magnesium’s use in surgery dates back to the late 19th century, as described in more detail by Witte.⁶⁾

The main feature that distinguishes magnesium alloys from other implant materials is the similarity to the mechanical properties of human bones. The closeness of Young’s modulus (41–45 GPa) and density (1.74–2.0 g/cm³) of magnesium alloys to bone (17.4–20 GPa and 1.7–2.1 g/cm³) makes it possible to minimize or even avoid the so-called stress shielding phenomenon and related problems such as loss of bone density, poor healing, implant loosening, etc.^7,8) In this respective, Mg alloys show advantages compared to permanent implant materials but also compared to other biodegradable materials such as iron–manganese or zinc-based alloys.⁷⁾ Moreover, Mg ions are common metabolites in the body with a daily intake of 240–420 mg/day and are naturally stored in the bones,⁹⁾ while for example the recommended daily consumption of Zn for an adult is about 50 times lower (8–11 mg/day).¹⁰⁾

Bioceramics, which are also biocompatible, and have other desirable properties such as good thermal stability, corrosion resistance, chemical inertness and good wear resistance are commercially used for example for coating of implants, drug delivery devices and dental implants.^11,12) However, bioceramics such as hydroxyapatite (HAP), alumina and zirconia are brittle. Polymers, both natural (e.g., collagen and protein-based gels, hyaluronic-based derivatives) and synthetic (e.g., poly-L-lactic acid - PLLA, poly lactic-co-glycolic acid - PLGA) also exhibit high biocompatibility and biodegradability, and have found use in many tissue engineering applications due to their properties similar to materials in the body.^13–15) In most cases, metal implants are preferred due to their high mechanical strength and durability, which makes them superior to polymers for load-bearing applications.⁷⁾ Therefore, magnesium alloys could largely replace permanent implant materials for many orthopaedic applications in the future. Moreover, studies show that magnesium-based alloys used as bone fixation stimulate bone formation in a physiological environment^16–18) and promote bone healing by participating in immunomodulatory, angiogenesis, osteogenesis and regulation of osteoclast function.¹⁹⁾ Other studies (in vivo and in vitro) prove that magnesium-based implants have anti-tumour properties and retard the growth of cancer cells through their degradation products Mg²⁺ and H₂.^20–23)

Despite the many advantages of magnesium alloys listed above, they also have their limitations. One of their major drawbacks is rapid corrosion occurring in water and body fluids, accompanied by the release of excessive hydrogen gas at the implantation site,^24–26) especially in enriched chloride environments and in the first 2–3 weeks after surgery.²⁶⁾ Slow and homogenous degradation of Mg implants is required, however, many factors affect the rate of corrosion in physiological environment, including concentration and types of ions, pH value and biochemical activities of surrounding tissues.^27–29) Over the last two decades, many strategies have been adopted to control the rate of degradation of magnesium alloys in such conditions. They involve for example the modification of the chemical composition and microstructure with appropriate heat treatments and surface modification with different corrosion-resistant coatings.^2,30,31)

Despite these improvements in the corrosion resistance of magnesium alloys, there are still many challenges in this matter, primarily matching the degradation rate of magnesium alloys with the individual recovery rate of different populations without adverse reactions.³⁰⁾

So far, various forms of magnesium alloy implants have been preclinically tested around the world and in vivo research confirmed the role of Mg-based implants in promoting fracture healing. However, the research methodology is diverse and there is no reference to a standardized model.³²⁾ A breakthrough in the use of biodegradable metallic implants was the clinical application of the MAGNEZIX^® compression screw in 2013.³³⁾ These studies demonstrated that the degradable Mg-based screw was radiographically and clinically equivalent to the conventional titanium screw. After clinical trials,^34–37) MAGNEZIX^® screws are now available on the market (with the CE-label and HSA approval for medical applications) for hallux valgus correction or fixation of small joints, for example, fragments of cartilage separated from the bone.³⁸⁾ Furthermore, as MAGNEZIX^® has similar X-ray density than cortical bone, it does not generate artefacts in X-rays, CT or MRT. Examples of magnesium-based implants, which are already used in medical applications, are shown in Fig. 1. These are resorbable magnesium scaffolds (RMS), Biotronik (Fig. 1(a)), as well as a screw, Syntellix AG (Fig. 1(b)) and pin, Syntellix AG (Fig. 1(c)).

Fig. 1

Magnesium-based implants: (a) resorbable magnesium scaffolds (RMS), Biotronik, (b) screw and (c) pin, Syntellix AG.

For magnesium-based implants, there is concern that the materials may degrade too quickly and lose their mechanical stability before the fracture has healed. Another concern is that the lower strength compared to titanium alloys requires the usage of bigger implants, e.g., screws with larger diameters. Therefore, the mechanical properties and mechanical integrity of Mg alloy implants in a physiological environment are important for successful fracture fixation surgery.

One of the strategies to improve the mechanical properties of magnesium alloys used for medical application, besides solid-solution hardening and second phase strengthening, is reduction of the grain size by severe plastic deformation (SPD) methods. It is well known that these methods offer the possibility to obtain submicron and or even nanometric structures, leading to advanced mechanical and functional properties not achievable by traditional methods. The history of using SPD methods to obtain improved mechanical properties of materials (mainly metals and their alloys) dates back to ancient times, however, the last two decades have brought a significant intensification of research.^39–42) This also includes magnesium, its alloys and other metallic materials and composites for implants, e.g. titanium-protein nanocomposites⁴³⁾ or high-entropy alloys.⁴⁴⁾ Apart from, there are also new tendencies to apply SPD processing to various kinds of intermetallics and non-metallic materials such as glasses, carbon polymorphs, semiconductors, ceramics and polymers.⁴⁵⁾ The most popular and effective SPD methods for obtaining ultra-fine grained materials or nanomaterials are equal-channel angular pressing (ECAP),⁴⁶⁾ high-pressure torsion (HPT),⁴⁷⁾ accumulative-roll bonding (ARB),⁴⁸⁾ twist-extrusion (TE)⁴⁹⁾ and multi-directional forging (MDF).⁵⁰⁾ In addition to the most well-known ones mentioned above, there are other methods: friction stir processing (FSP),⁵¹⁾ hydrostatic extrusion,⁵²⁾ cyclic extrusion compression (CEC),⁵³⁾ KoBo extrusion,⁵⁴⁾ high-pressure sliding⁵⁵⁾ and even more.^39–41)

In conclusion, for biodegradable magnesium alloys, in addition to biocompatibility and bioactivity within the human body, appropriate mechanical properties with a controlled degradation rate in body fluids at 37°C without unacceptable accumulation of degradation products around the implantation site play a crucial role. All the SPD methods mentioned above are able to improve the mechanical properties without changing the chemical composition, thus, with no or only minor influence on biocompatibility and bioactivity. Therefore, this review focuses mainly on the research on the improvement of mechanical properties by grain refinement of magnesium alloys for medical applications using SPD methods.

2. Magnesium Alloys Processed by ECAP

ECAP is one of the most popular SPD methods and can deform a wide range of materials. It generally leads to a fine-grained structure of polycrystalline metals to the sub-micron level with accompanying significant improvement of mechanical properties. A solid material in the form of rods or bars is – usually repeatedly – pressed through a die with a bent channel (with an internal angle, usually 90° or 120°) without changing its cross-section, resulting in a large plastic deformation by simple shear.⁴⁶⁾ The principles of this method have been described in detail for example by Valiev and Langdon.⁵⁶⁾ ECAP is among the most developed SPD processes used worldwide and the most popular method applied to refine the microstructure of magnesium alloys for various applications because the material volumes obtained are large compared to other SPD methods. So far, also variations of the ECAP process have been developed, e.g. I-ECAP (Incremental-ECAP),⁵⁷⁾ TCAP (Twist channel angular pressing),⁵⁸⁾ D-ECAP (Double-ECAP).^59,60) Several parameters, such as the operating temperature (typically 100°C–350°C for Mg alloys), the pressing speed, the channel angles (an internal and a curvature angle), the number of passes and the rotation of the sample between subsequent passes (routes A, B_A, B_C and C), lubrication and the application of back pressure control the deformation process in this method.

2.1 Microstructure

Below the ductile-brittle transition temperature, magnesium and its alloys exhibit only two independent slip systems (basal slip {0001}$\langle 11\bar{2}0\rangle $) due to the hexagonal close-packed crystal structure.⁶¹⁾ Therefore, the material does not meet the Taylor criterion of five independent slip systems required for homogenous plastic flow⁶²⁾ and is rather brittle at room temperature. Thus, magnesium alloys are usually processed at elevated temperatures. Under these conditions, prismatic ($\{ 10\bar{1}0\} \langle 11\bar{2}0\rangle $) and pyramidal slip systems ($\{ 112\bar{2}\} \langle 11\bar{2}3\rangle $) are also activated, causing the material to become ductile. In addition, plastic deformation of magnesium alloys is supported by the formation of deformation twins, particularly at process temperatures below 250°C or in cases where the initial grain size is large. The effect of grain size and temperature on the deformation mechanism was investigated for a AZ31 alloy during compression and it could be shown that a transition from twinning to slip-dominant flow occurs with increasing temperature and decreasing grain size.⁶³⁾ Similar observations were made for ZK60 alloy during ECAP at 200°C.⁶⁴⁾ Twining occurred during the first two ECAP passes, but the twin density decreased with subsequent passes and decreasing grain size. Moreover, the twins are preferred recrystallization sites for new grains. In a further example, a large volume fraction of twins was observed after the first ECAP pass of a Mg–4Ag alloy at 370°C which reduced the grain size from initially about 350 µm down to 38 µm.⁶⁵⁾

During the ECAP process, the microstructure is severely refined by mechanical shearing but especially in case of Mg alloys also dynamic recrystallization (DRX) plays a significant role due to the high process temperature (higher than half the melting point), the energy stored in deformation-induced defects and the low deformation rates. The degree of refinement of the microstructure depends significantly on the initial structure of the processed alloy. In most cases, ECAP-treated magnesium alloys are pre-prepared by heat treatment,^65–67) rolling^68–70) or extrusion.^60,66,71) The smaller the initial grains, the fewer ECAP passes are needed to obtain a homogeneous and fine-grained structure. For example, the ZK60 alloy was processed by 6 ECAP passes with various initial structures, i.e. after annealing with an average grain size of about 180 µm and after extrusion with an average grain size of approximately 2.9 µm.⁶⁶⁾ As a consequence, quite different grain structures were observed after the same ECAP process. In the first case, the obtained microstructure had a bimodal character, it consisted of large grains in the range of 20–50 µm and fine grains with a size of about 1 µm surrounding the large ones. In the second case, an ultrafine-grained structure was obtained with an average size of about 0.8 µm. However, in another study of the AX41 (Mg–4Al–Ca, mass%) alloy processed by ECAP at 220°C up to 8 passes via route B_C,⁷²⁾ it was shown that it is possible to obtain finer grains despite the coarser initial grain size in the as-cast state (200 µm) compared to the state after extrusion (10 µm). In the first case a grain size of 1.4 µm is obtained after ECAP, while in the second case it was 2.4 µm. This difference was attributed to different dislocation densities and its fractions in different slip systems. A general grain refinement model of Mg alloys processed by ECAP, depending on the initial grain size and based on the principles of DRX, was proposed by Figueiredo and Langdon.⁷³⁾ Based on this model, it can be concluded that bimodal or multimodal grain size distributions, usually reported after ECAP, are transient and can be changed if pressing is continued for a sufficient number of passes.

During the ECAP process, a texture is induced which also affects the mechanical properties of the processed material. There have been many studies on ECAP texture development focusing on magnesium alloys.^74–77) Some of the results showed a post-ECAP softening effect, which was ascribes to the alignment of the main basal plane (0002) with the shear plane of the ECAP tool. This caused a decrease in the yield strength in extrusion direction but also a significant improvement of the ductility as compared to the as-extruded state.^74,77) For example, ECAP processing of AZ31 magnesium alloy enforce a crystallographic reorientation resulting in texture-induced softening and increased ductility.⁷⁸⁾ Similarly, it was shown that via modifying the microstructure of AZ31B alloy by I-ECAP with different processing routes (B_C and A) its mechanical properties - strength and ductility - can be varied and controlled and thereby the texture plays an important role.⁵⁷⁾ As an example, EBSD (Electron Backscatter Diffraction) maps and corresponding micro-textures of AZ31B alloy after I-ECAP are shown in Fig. 2. After four passes at 250°C and two passes at 200°C using route A, two overlapping peaks, indicating hcp cells tilted at 75° and 62° to the extrusion direction (ED), can be distinguished on the (0001) pole figure, while texture obtained after the same ECAP procedure using route B_C revealed that hcp cells are tilted at 57° to ED. It was proved that the texture produced by route A was more effective in improving the strength than route B_C, hence, the role of the texture cannot be neglected.

Fig. 2

EBSD maps and corresponding micro-textures of AZ31B alloy: (a), (d) as-extruded (initial); (b), (e) after 4 I-ECAP passes at 250°C and 2 I-ECAP passes at 200°C using route A; (c), (f) after 4 I-ECAP passes at 250°C and 2 I-ECAP passes at 200°C using route B_C.⁵⁷⁾

Another effect of the ECAP process on the microstructure of magnesium alloys is the fragmentation and redistribution of hard particles in the magnesium matrix. In the AZ91 alloy, fragmentation as well as precipitation of Mg₁₇Al₁₂ phase particles during ECAP were observed, which affected strength and plasticity, apart from grain refinement.⁷⁹⁾ A Mg–10Al–0.5Sb alloy obtained excellent mechanical properties at room and high temperatures due to the grain refinement and fragmentation of β-Mg₁₇Al₁₂ and Mg₃Sb₂ particles.⁸⁰⁾ An additional example is the fragmentation and redistribution of hard and brittle (Mg,Zn)₁₂RE and T-phase (Mg₇Zn₃RE) particles in the ECAP-processed EZ33A and ZE41A alloys, respectively, that decorated grain boundaries forming a semi-continuous network.^81,82) Fragmentation of these networks in both alloys and redistribution of particles have contributed to the strengthening and increase of ductility. Moreover, during the ECAP, nanometric (Mg,Zn)₁₂RE particles (below 100 nm) precipitated, which contribute to the strengthening of the EZ33A alloy by inhibiting the movement of dislocations.

2.2 Mechanical properties

Magnesium alloys have significantly higher strength and moderate ductility due to the ECAP process, so they have great potential for medical implant applications. Table 1 shows the mechanical properties obtained by the ECAP process for selected magnesium alloys discussed below.

Table 1 Mechanical properties of selected magnesium alloys processed by ECAP.

2.2.1 Mg–Al based alloys

The Mg–Al system is the most studied one also in the context of implant materials because the resulting alloys are characterized by good mechanical properties, excellent castability and sufficient corrosion resistance. The addition of aluminium leads to solid solution strengthening, as its solubility in magnesium is high (about 12.7% by weight), as well as precipitation strengthening through Mg₁₇Al₁₂ particles which typically form along the grain boundaries. Alloys with an Al content of up to 6 mass% are intended for plastic working and mainly squeeze die casting. To improve their mechanical properties, mainly Zn, Si and Mn are also added to Mg–Al alloys.

The AZ31 is the most popular industrial magnesium alloy processed by ECAP. Strong refinement of grains can be obtained, depending on the initial grain size, process temperature and the number of passes (Table 1). For example, after 4 passes at 300°C with reducing the grain size from 27 µm to 8 µm, the yield strength (YS) and tensile strength (TS) were improved to 160 MPa and 220 MPa, respectively.⁸³⁾ However, in another study with a lower process temperature of 150°C, the grain size was reduced from 10 µm to 0.9 µm, giving an even more significant increase in the yield strength (305 MPa) and tensile strength (380 MPa) at similar elongation to fracture.⁵⁷⁾ The grain refinement in AZ31 and AZ61 was investigated in various ECAP experiments with different processing temperatures, number of passes and initial states of the microstructure in particular different initial grain sizes.^34,84–88) A significant improvement of the mechanical properties of the AZ91 alloy, namely an increase in yield strength and tensile strength to 290 MPa and 417 MPa, respectively, was achieved by two-stage ECAP at two different temperatures (with a total of 6 passes). The most promising magnesium alloys for biodegradable implants with aluminium with regards to their mechanical properties and corrosion resistance contain rare earth elements and even lithium: AE21, AE42, and particularly LAE442.^89–92) Significant grain refinement was obtained in these alloys by ECAP processing.^93,94) After 4 passes, the average grain size was reduced to about 2.5 µm. This was accompanied by an increase in yield strength which was largest for the LEA442 alloy. However, this is not only due to the grain size reduction but also due to the addition of lithium influencing the c/a ratio and the texture created during the ECAP process.

Although aluminium-containing magnesium alloys have good tensile strength and good corrosion resistance,⁸⁶⁾ the potential use of such alloys as AZ61, and AZ91 as biodegradable implants is limited because the neurotoxic effect of aluminium when exceeding the daily dose can cause Alzheimer’s disease or dementia^95,96) and increases the risk of muscle fibre damage.⁹⁷⁾ Thus, more attention has been paid recently to magnesium alloys with additives that are components of nutrition like Zn, Ca, Sr, etc.

2.2.2 Mg–Zn, Mg–Ca and Mg–Zn–Ca based alloys

Both Zn and Ca are important metals in the human body that support many processes including bone healing.^98,99) Due to the relatively high solubility of Zn in magnesium (6.2 mass%), it may increase strength by combining solid-solution strengthening and precipitation hardening by intermetallic compounds.¹⁰⁰⁾ Mg–Zn alloys are characterized by good biocompatibility.^101,102) To improve the mechanical properties, they have also been ECAP processed. For example, for the ZK60 alloy (Mg–6%Zn–0.5%Zr), after 4 ECAP passes, an increase in tensile strength from 285 MPa in the initial state to 326 MPa was obtained.¹⁰³⁾ Moreover, after 8 ECAP passes of this alloy, a notable increase in the yield strength (260 MPa) and ultimate tensile strength (371 MPa) with a good ductility (18.5% elongation at fracture) was obtained. This increase in mechanical properties was associated with a significant grain refinement (from 200 µm to 3–0.5 µm) and precipitation of metastable particles (MgZn₂) during the ECAP process.¹⁰⁴⁾

The addition of Ca has a positive effect on the corrosion behaviour and cytocompatibility of magnesium.^105,106) Moreover, Ca affects the grain size reduction and increases the strength due to Mg₂Ca precipitates along the grain boundaries in alloys with up to 1% of Ca. Mg alloys with higher Ca concentration possess a lower strength, but also lower ductility and corrosion resistance.^107,108) Due to the rather poor mechanical properties in the as-cast condition, the Mg–1Ca alloy was processed by ECAP. Significant grain refinement was obtained from 150 µm to 1 µm after 4 passes at 400°C and 2 subsequent passes at 300°C. In addition, the Mg₂Ca particles after the ECAP process were more evenly distributed in the microstructure. This is accompanied by a noticeable increase in yield strength to 110 MPa, tensile strength to over 200 MPa and elongation to 7%.¹⁰⁹⁾

The Mg–Zn–Ca system is one of the most promising orthopaedic biodegradable magnesium alloy systems, having a better combination of strength properties and degradation rate than the two systems mentioned above. For example, studies have shown that Mg–Zn–Ca (ZX)-lean alloys: ZX10 (Mg–1Zn–0.3Ca, mass%) and ZX20 (Mg–1.5Zn–0.25) have good strength properties and slow and uniform in vitro and in vivo degradation.^110,111) Furthermore, the biocompatibility and strength retention of Mg–Zn–Ca alloy screws implanted into New Zealand rabbits was in detail studied and compared with a commercially available biodegradable polymer implant made of PLLA.¹¹²⁾ It was found that the strength retention of Mg–Zn–Ca alloy bone screws was similar to those of biodegradable polymers with Mg alloy having a generally higher strength level than polymer. Therefore, this alloy is an excellent biodegradable biomaterial for osteosynthesis applications.

To improve the mechanical properties of the low-alloy magnesium, it was processed by ECAP as well. For example, a series of D-ECAP experiments at various combinations of process temperatures and passes were performed for the ZX00 alloy (Mg–0.6Zn–0.5Ca, in mass%).⁶⁰⁾ The results of these tests showed the strong influence of the processing parameters, e.g. the tensile strength was greatly increased from 225 MPa (as-extruded) to over 370 MPa after 4 passes at 280°C, with an elongation at fracture of 7%. After 2 passes at 300°C, the value of the tensile strength was 298 MPa and the elongation at fracture was as much as 26%. Such values of mechanical properties were a consequence of grain refinement, but also different proportions of coarse-grained and very fine-grained fractions, resulting from dynamic recrystallization during processing at elevated temperatures. Based on these results, it can be concluded that the mechanical properties of Mg alloys can be controlled in an extensive range by changing the parameters of the ECAP process - the chemical composition of the material is not changed. A comparison of these results for the ZX00 alloy with other Mg–Zn–Ca alloys with higher alloying content is presented in Fig. 3. The outstanding combination of mechanical properties of low-alloyed material can be attributed to using D-ECAP die exhibiting a higher degree of deformation per pass as well as an increased hydrostatic pressure.⁶⁰⁾ For instance, biodegradable Mg–4.71Zn–0.6Ca (mass%) alloy was processed by 4 conventional ECAP passes at 250°C. This led to a reduction of the grain size from 54.5 µm to 1.6 µm and an increase in tensile strength from 206 MPa to 290 MPa.¹¹³⁾ However, despite the higher among of alloying elements, the strength is significantly lower than that of the lean ZX00 alloy after D-ECAP.

Fig. 3

Comparison of tensile strength and elongation at fracture for Mg–Zn–Ca alloys processed by ECAP with different content of alloy additions.^{60,195–200)}

2.2.3 Mg–Zn-RE, Mg-RE based alloys

Inserting rare earth elements (RE) into magnesium alloys is an effective way to improve their mechanical properties, both by solid solution and precipitation strengthening.^114–116) Most Mg-RE alloys have enhanced mechanical properties compared to other Mg alloy systems and improved corrosion resistance both in in vitro and in vivo.¹¹⁷⁾ Some of the RE elements exhibit biocompatibility within specific amounts of addition.^116,117) However, it has been reported that some RE elements like Ce, La, Pr, Eu have toxic effects.^118–121) Therefore, it is necessary to conduct further research on the toxicity of rare earth metals and carefully select the chemical composition of magnesium alloys.

An example of a successfully used magnesium alloy with additions of rare earth metals in the technology of bioresorbable scaffolds providing temporary vascular support with the possibility of drug delivery is the WE43 (Mg–Y–RE–Zr) alloy.^122,123) The technology of absorbable metal scaffolds using this magnesium alloy with modifications is constantly developed to improve their performance and safety. Moreover, screws made of Mg–Y–RE–Zr alloy were implanted in the femur of rabbits and their chronic local effects on bone tissue and systemic reactions were studied. This study showed that Mg–Y–RE–Zr screws have good biocompatibility and osteoconductivity without chronic toxicity.¹²⁴⁾ Other pilot studies demonstrated that such screws are radiographically and clinically equivalent to the conventional titanium screw for fixation during chevron osteotomy in patients with a mild hallux valgus.¹²⁵⁾ Similar results were obtained with the MAGNEZIX^® CS compression screw (also made of Mg–Y–RE–Zr alloy), which was applied to the fixation of displacement 1st metatarsal osteotomies in the surgical management of hallux valgus.¹²⁶⁾

The effect of ECAP on the mechanical properties of the WE43 alloy was also investigated, proposing two different approaches with varying process temperatures.¹²⁷⁾ In the first approach, ECAP was performed in two stages: 6 passes at 400°C and the subsequent 6 passes at 350°C. In the next approach, the process temperature was gradually lowered by 25°C, starting from 425°C to 300°C. Two ECAP passes were performed at each temperature. ECAP changed this alloy’s grain size (from initial 70 µm) radically, leading to the formation of an ultrafine-grained structure with an average grain size of 1 µm for the first approach and 0.7 µm for the second approach. This resulted in a significant increase in strength properties. The tensile strength increased from 220 MPa in the initial state to 250 MPa by the first approach and 300 MPa by the second approach. An even higher tensile strength for the WE43 alloy was obtained in another study after 12 passes at 375°C, i.e. 322 MPa, thanks to grain refinement (from 50 µm to 1.5 µm) and homogeneity of the alloy microstructure after ECAP.¹²⁸⁾

Commercial EZ33 and ZE41 alloys, mainly used in the aerospace and military industries, are also considered potential materials for medical implants.¹²⁹⁾ ECAP processing significantly improved the mechanical properties of these alloys. For example, after four passes through the D-ECAP die with stepwise decreasing temperature from 350°C to 240°C, a significant increase in yield strength and tensile strength for the EZ33A (Mg–2.5Zn–0.4Zr–3RE, mass%) alloy was obtained, i.e. 327 MPa and 346 MPa, respectively.⁸¹⁾ Such strength characteristics were obtained both by refining the grains and by fragmentation and redistribution of the hard intermetallic (Mg,Zn)₁₂RE phase by ECAP. Identical parameters of the ECAP process were also adopted for the ZE41A (Mg–4Zn–0.52Zr–1.3RE, mass%) alloy. After ECAP processing, the yield strength was improved by nearly 150% from the initial condition (from 165 MPa to 321 MPa) and the tensile strength increased from 186 MPa to 336 MPa. These changes in mechanical properties of the ZE41A alloy were caused by grain refinement (with a bimodal grain size distribution) and the fragmentation and the redistribution of the ternary phase particles through the ECAP process.⁸²⁾ Another study¹³⁰⁾ of this alloy demonstrated a homogeneous distribution of essentially equiaxed grains with a size of approximately 1.5 µm after 32 ECAP passes at 330°C. The result was an increase in yield strength to about 275 MPa, tensile strength to about 310 MPa and elongation at fracture to 14%.

2.2.4 Mg–Ag based alloys

The addition of antibacterial silver to magnesium is an interesting proposal for biodegradable medical implants, which might reduce the symptoms of local inflammation or infection that may arise in the first weeks after implantation. Studies carried out on Mg–Ag alloys with different silver content (2%, 4% and 6%, mass%) showed cytocompatible and antibacterial properties.¹³¹⁾ Other studies on as-extruded Mg–Ag alloys confirmed their antibacterial properties with almost equivalent cytocompatibility to human primary osteoblasts as pure Mg.¹³²⁾ However, as the mechanical properties were insufficient for implant applications, these alloys were processed using the ECAP method. For example, the Mg–4Ag alloy processed by two ECAP passes, the first at 370°C and the second at 330°C, demonstrated an increase in compressive strength from 290 MPa after T4 heat treatment (with an average grain size of 350 µm) to 325 MPa, and an increase in the compressive yield strength from 31 MPa to 62 MPa (with an average grain size of 15 µm).⁶⁵⁾ Performing 12 ECAP passes at stepwise decreasing temperature (from 375°C to 250°C with a reduction after every second pass), the grains of Mg–2Ag and Mg–4Ag alloys were refined from about 40 µm to 2–3 µm.¹³³⁾ Despite this, a decrease in the strength characteristics compared to the initial state was observed, although nearly twice the increase in ductility (up to 30%) due to the basal and prismatic slip promoted by ECAP. Almost identical results were obtained for the Mg–6Ag alloy after a comparable ECAP process.¹³⁴⁾

3. Magnesium Alloys Processed by HPT

The HPT process introduces especially high strains to materials and results in grain refinement to a size typically finer than that produced by other SPD methods, e.g. by the ECAP method. The fundamentals of HPT and its history have been described in detail in the literature.^39,135,136) In this technique, the material, in the form of disc or ring, is torsionally strained between two anvils under high pressure. The process is usually performed under quasi-constrained conditions,³⁹⁾ allowing limited material flow between the upper and lower anvils and a high hydrostatic pressure. There are also some variants of the HPT method, e.g. high-pressure tube twisting (HPTT),¹³⁷⁾ and shear diamond anvil cell (SDAC).¹³⁸⁾ The imposed strain can be higher than any other SPD technique because of the rotational feature of the process. However, it is not uniform over the sample but depends on the radius. The combination of high hydrostatic stress and continuous shear stress in HPT makes it possible to process brittle and hard deformable materials such as magnesium alloys at ambient temperature. However, the main limitation of the process with regard to industrial applications is the small size of the discs, usually 10 mm in diameter and 1 mm in thickness. Despite the small sample sizes, magnesium alloys processed by the HPT method with additional heat treatment can be used for thin barrier membranes of several tens of micrometres for guided bone regeneration (GBR). Such degradable membranes,¹³⁹⁾ produced by rolling and Ca–P coated, have been used in the clinical study to separate bone graft material from the surrounding fibrous tissue and provide for bone regeneration, avoiding a second removal operation.

The HPT process is controlled by the following parameters: temperature, pressure, strain by the number of revolutions and strain rate by rotation speed. Magnesium alloys are usually processed at ambient temperature, at the pressure of 6 GPa, with several revolutions from 0.25 to 10 and at a rotation rate of 1 rpm. A significant achievement was using an up-scaled HPT facility with a maximum capacity of 5 MN (500 tons) to process disks with 30 mm in diameter of AZ31 and AZ61 magnesium alloys.¹⁴⁰⁾ This opens new opportunities, both in terms of research and future commercialization of the HPT method.

3.1 Microstructure

An intrinsic characteristic of HPT is that the applied strain depends on the radius and in theory it is zero at the centre of the processed disk. Therefore, it can be assumed that the microstructure produced by HPT will be exceptionally inhomogeneous. However, available experimental studies show that due to saturation tendencies a gradual evolution to relatively uniform microstructures is possible for many materials, including magnesium alloys.¹³⁵⁾ Information on the homogeneity of the microstructure can be taken or by using scanning or transmission electron microscopy or also indirectly by hardness measurements along the radius of the processed disk.

In the early stages of HPT processing of magnesium at ambient temperatures, multimodal grain size distributions evolve. The continuation of plastic deformation causes an increase in the volume fraction of ultra-fine or even nanosized grains until the total proportion of initial coarse grains is consumed. The research shows that grain refinement starts with sub-grain formation, shear banding and twins.^141–143) The typical microstructure of magnesium and magnesium alloys processed by the HPT method consists in the early stage, namely after 1/8, 1/4 and 1/2 turn, of coarse-grained areas, bands of elongated sub-grains and fine grains, especially along original grain boundaries.^143–146) Moreover, as shown in Fig. 4, the grain sizes vary through the radius and thickness of the HPT sample of AZ31 alloy,¹⁴⁴⁾ which indicates the heterogeneity of the deformation. The grain refinement and the formation of elongated bands of sub-grains occurred already during the early stages of HPT straining after 1/4 turn (Fig. 4(d)), whilst a similar effect in the centre of the disk was observed only after 1 turn (Fig. 4(a)). Moreover, nanograins with sizes in the range of 150 nm to 250 nm are observed after 1 turn in the periphery of disks (Fig. 4(e)), whereas a similar effect was obtained after 15 turns in the centre of the disk (Fig. 4(c)). Additional HPT straining results in slight grain fragmentation without substantial microstructure changes.

Fig. 4

TEM micrographs showing the evolution of microstructure of HPT-processed AZ31 alloy in the near centre (a), (b), (c) and edge (d), (e), (f) after different numbers of turns.¹⁴⁴⁾

The area fraction of fine grains increased with increasing imposed strain and refined grains dominated most of the microstructure after 1 turn. However, the number of turns where almost complete homogenous structure was achieved depends on the pressure applied during the HPT process. An example is deformation of an AZ31 alloy by HPT at different hydrostatic pressures, i.e. 2.5 GPa¹⁴⁴⁾ and 6 GPa,¹⁴⁷⁾ where a homogeneous microstructure was obtained after 15 and 5 rotations, respectively.

In addition to pressure, process temperature also is important. The variation of average grain size with the number of turns and with the temperature at different positions of disks after the HPT of AZ31 alloy was reported.^148,149) Processing at different temperatures resulted in different grain structures also affecting the microhardness values.¹⁴⁹⁾ Sub-micron grain structure was observed after HPT at room temperature, as well as at 100°C, whilst the process at 200°C led to grain growth during processing, so that the grain size after 5 turns was larger than after 1 turn, and the microhardness reduced. Similarly, for the Mg–Y–Gd–Zr (Mg–4.7Y–4.6Gd–0.3Zr, mass%) alloy processed by HPT,¹⁵⁰⁾ a significant difference in grain size was observed, i.e. grains of 20–30 nm were obtained for the process at room temperature, and of 60–90 nm at 200°C.

Magnesium alloys show a strong texture after plastic deformation, hence a significant texture is also revealed in HPT-processed disks and can be expressed as a function of strain. The texture during HPT evolves from as-received state to basal (0001) fibre texture, which is attributed to the activation of twinning. A case in point, at early stages of deformation of the disks of ZK60¹⁵¹⁾ and AM60¹⁴⁵⁾ alloys, i.e. after 1/2 turn, the $(00\bar{1}0)$ texture remained from the as-extruded state (Fig. 5(a)). When the number of turns increases to one, the $(00\bar{1}0)$ texture disappeared and evolved to the majority of basal (0001) texture with the c-axis parallel to normal direction of the disk (Fig. 5(b)). However, with the increasing HPT straining, i.e. after 3 and 10 turns, the basal texture weakened (Fig. 5(c) and 5(d)), which also led to a decrease in the hardness value.¹⁴⁵⁾ Similar texture evolution was reported for AZ31 alloy,¹⁴⁸⁾ for Mg–0.41Dy¹⁵²⁾ alloy and for pure magnesium.¹⁵³⁾

Fig. 5

Grain orientation maps and corresponding inverse pole figures of AM60 alloy processed by HPT for the centre of the disks after: (a) 1/2 turn, (b) 1 turn, (c) 3 turns and (d) 10 turns.¹⁴⁵⁾

As in the ECAP process, particles in the magnesium alloy matrix are also fragmented and redistributed during the HPT processing. In this regard, the magnesium alloy EZ33A was investigated by both, ECAP and HPT methods.^81,154) As a result of the strong deformation, the hard and brittle (Mg,Zn)₁₂RE particles decorating the grain boundaries in the as-cast condition undergo strong fragmentation and redistribution. However, smaller particles were obtained after the ECAP process despite a considerably lower strain. The high hydrostatic pressure of the HPT process likely prevents these particles from breaking and spreading further in the magnesium alloy matrix.

3.2 Mechanical properties

The HPT process significantly improves the mechanical properties of magnesium alloys. Table 2 summarizes the mechanical properties of selected magnesium alloys after HPT. Due to the small dimensions of the samples, most of the papers report hardness test results only.

Table 2 Mechanical properties of selected magnesium alloys processed by HPT at room temperature.

3.2.1 Mg–Al based alloys

As mentioned earlier, the AZ31 alloy is one of the Mg alloys most often subjected to SPD processes. The HPT method allows processing at room temperature without cracking or segmentation, as in the case of the ECAP method.^85,155,156) It can be seen that an AZ31 alloy subjected to HPT processing has several times smaller grain size (∼100–200 nm) compared to the alloy processed by ECAP at elevated temperatures (∼0.9–9 µm, see Table 1), which is accompanied by a significant increase in hardness (∼120 HV)^{140,144,147,157)} compared to ∼87 after ECAP.¹⁵⁸⁾ For the AZ91 alloy with a higher aluminium content, a spectacularly small grain size of about 35 nm was observed after HPT processing at ambient temperature,¹⁵⁹⁾ which was reflected in an increase in hardness to 135 HV. A similar hardness value (130 HV) was obtained for the AM60 alloy after grain refinement from 16 µm to 800 nm after only 1/2 HPT turn.¹⁴⁵⁾ Compared to ECAP carried out for this alloy at 220°C, the grain size after the HPT process was about eight times smaller, and the maximum dislocation density was twice as high. This difference can be explained by the retarding influence of lower temperature and higher pressure on diffusion during HPT, which hindered dislocation annihilation. As a result, the hardness of the sample processed by HPT was about 30% higher than that processed by ECAP.

3.2.2 Mg–Zn, Mg–Ca and Mg–Zn–Ca based alloys

As previously mentioned, when discussing the post-ECAP mechanical properties of this group of materials, biodegradable magnesium alloys only alloyed with essential elements found in the human body, such as Zn and Ca, can support the bone healing without side effects. Based on this approach, many alloys of this group have also been HPT processed with the aim to obtain high strength materials for orthopaedic applications. A comparison of tensile strength and elongation at fracture for Mg–Zn–Ca alloys processed by HPT with different content of alloy additions is presented in Fig. 6.

Fig. 6

Comparison of tensile strength and elongation at fracture for Mg–Zn–Ca alloys processed by HPT with different content of alloy additions.^{160–162,191)} The number in parentheses after HPT indicates the number of turns. The HT designation means an additional heat treatment after the HPT process.

The HPT-processed Mg–1Ca (mass%) alloy showed improved mechanical properties as its microstructure was modified and the grain size was reduced to 171 nm: a hardness of 91 HV, yield strength of 229 MPa and tensile strength of 316 MPa. However, the elongation at fracture only reached 1.6%. A much higher value of about 7% was obtained for this alloy processed by ECAP,¹⁰⁹⁾ but the other parameters are much lower (YS = 110 MPa, TS = 200 MPa) as the grain size is as large as 1 µm.

Application of the HPT process to the Mg–1Zn–0.13Ca (mass%) alloy caused a significant grain size reduction from 30 µm to 150 nm and the accompanying increase in hardness from 45 HV of initial coarse-grain state to 99 HV.¹⁶⁰⁾ Due to the low ductility of magnesium alloys after the HPT process, additional annealing at 200°C was used in this case. This resulted in an increase in grain size to 1.5 µm but also a higher TS (270 MPa) as compared to the initial coarse-grained material (140 MPa) with retention of a good elongation at fracture (8.5%). The improvement of TS was explained by grain refinement and dispersion hardening and retention of good ductility by activation of dislocation slip in non-basal planes.

Another study showed that the HPT processing of low-alloyed Mg–Zn–Ca alloys causes an increase in strength and hardness due to grain refinement.¹⁶¹⁾ Moreover, this alloy shows an additional substantial strength increase caused by post-HPT heat treatment. A hardness of more than 110 HV and tensile strength of more than 300 MPa were obtained in Mg–0.2Zn–0.5Ca by aging treatment at 142°C for 1 hour after 2 turns of HPT. However, the elongation at fracture was very low (1.1%). Due to the low content of alloying elements, the strength by the heat treatment of the HPT deformed alloy is not associated with precipitation hardening but with the high number of vacancies after HPT processing, which form agglomerations and dislocation loops that act as obstacles and hinder the movement of dislocations.

Similarly, Mg–Zn–Ca alloys with higher Zn content (Mg–5Zn–0.3Ca, Mg–5Zn–0.15Ca, Mg–5Zn–0.15–Ca–0.15Zr, mass%) in different initial states (homogenised and either furnace-cooled or quenched) were investigated after HPT and aging treatment, which led to the precipitation of intermetallic particles and vacancy agglomerates and thus tremendous increases of the alloy’s strength.¹⁶²⁾ For instance, a Mg–5Zn–0.15Ca alloy after HPT and aging treatment at 100°C for 24 hours reached a YS of 395 MPa and a TS of 418 MPa with the initial values being YS = 120 MPa and TS = 339 MPa. However, the material was also rather brittle and the elongation at fracture was reduced from 17% to 5%. An analogous tendency after the HPT process combined with post-HPT heat treatment was observed for the other alloys.

3.2.3 Mg–Zn-RE, Mg-RE based alloys

As earlier mentioned, in the case of Mg alloys, the addition of rare earth elements results in a significant increase in mechanical properties compared to traditional commercial Mg alloys. The relatively high solubility of RE elements in the Mg matrix causes substantial solution strengthening and additional precipitation strengthening after ageing. It was shown that HPT significantly strengthens the WE43 alloy.¹⁶³⁾ HPT processing (10 rotations) resulted in a partially nanocrystalline structure with a grain size of 30–100 nm. As a result of this, hardness, YS and TS increased from 77 HV to 120 HV, from 161 MPa to 353 MPa and from 234 MPa to 371 MPa, respectively. However, as in the previous cases considered above, the plasticity of the material was reduced. An additional ageing of the material at 200°C caused a further increase in YS and TS with the elongation to break remaining at the same level of 1%.

A Mg–4.7Y–4.6Gd–0.3Zr (mass%) alloy in the annealed and quenched state was processed by HPT up to 10 turns at the temperatures of 20–200°C. That led to a partially nanocrystalline structure with an average grain size of 60–90 nm.¹⁵⁰⁾ By HPT at 200°C, a YS of 450 MPa, a TS of 475 MPa and an elongation at fracture of 2.5% were obtained. Furthermore, HPT samples were also isothermally aged at 200°C for 16 hours, and additional hardening due to a precipitation process was achieved. Thereby, the hardness increased from 140 HV after HPT to 155 HV.

A similar strategy was proposed for a Mg–4.97Sm–0.84Ca (mass%) alloy: the alloy in the solution-treated state was subjected to hot rolling and HPT, followed by ageing at 150°C for 3.5 hours.¹⁶⁴⁾ This led to an exceptionally large increase in the hardness of 175 HV which was 165% higher than for the solution-treated alloy.

Magnesium alloys with additions of lithium after HPT are characterized by a high plasticity. As an example, HPT processing of a LZ91 alloy (Mg–9Li–1Zn, mass%) with 10 turns resulted in a reduction of the grain size from 30 µm to 230 nm and good hardness and tensile strength, but especially the ductility was improved significantly and elongation at fracture reached about 400% at 200°C.¹⁶⁵⁾ It is worth mentioning that superplasticity was also observed for a Mg–8Li (mass%) alloy after ECAP, where the maximum elongation of ∼970% was reached at 200°C after combination of extrusion and ECAP.¹⁶⁶⁾ Moreover, other studies have shown superplasticity of the Mg–9Li alloy even at room temperature after 200 HPT revolutions,¹⁶⁷⁾ where the transition to grain-boundary sliding can be attributed not only to the grain refinement to 230 nm but segregation of lithium to grain boundaries that influence their mobility.

4. Corrosion and Biodegradation of Mg Alloys

Magnesium is one of the most non-noble metals with a corrosion potential of −1.65 V_SCE.¹⁶⁸⁾ Therefore, its corrosion rate in technical as well as in biological environments is high. When using Mg alloys for biodegradable implants, controlling the degradation rate is a big issue. The implant should not degrade too fast, as for example the bone which is fixed by it needs a certain time to heal, but also localized corrosion or pitting might have detrimental effects on the stability of load-bearing implants.

The basic corrosion mechanism of Mg in aqueous solutions is Mg + 2H₂O → Mg(OH)₂ + H₂. The evolving magnesium hydroxide forms a partly protective layer on the surface, which may also be transformed into MgO. The further corrosion steps depend strongly on the environment, i.e. the test solution or the surrounding of an implant in a living organism. Especially chlorine ions have a strong influence as they dissolve the magnesium hydroxide layer.¹⁶⁹⁾ Beside magnesium oxide and hydroxide surface layers, also carbonates and phosphates can precipitate on the surface of Mg. This depends on the composition of the test solution, however, calcium phosphate formation is frequently observed when testing in physiological environments.¹⁷⁰⁾ Such insoluble surface layers lead to a passivation of the surface and a decrease of the degradation rate. Under in vivo conditions, even more reactions and processes occur, e.g. adherence of proteins and cells.¹⁷¹⁾

Several papers deal with mimicking physiological conditions in the laboratory, the influence of different testing methods and solutions and the comparison of degradation rates and behaviour observed in in vitro and in vivo experiments.^172–175)

The influence of alloying elements and thermomechanical treatments of Mg alloys on corrosion and biodegradation partly depends on the testing environment, but anyhow, some general conclusions can be drawn. Alloying usually changes the corrosion potential of Mg. When intermetallic phases are formed, they normally have a higher corrosion potential than the surrounding Mg matrix (Mg₁₇Al₁₂ for example has −1.35 V_SCE¹⁶⁸⁾). This leads to galvanic coupling and micro-galvanic corrosion¹⁷⁶⁾ which accelerates the corrosion rate of the Mg matrix while the particles are not or only very slowly dissolved. One exception is Mg₂Ca, which exhibits a corrosion potential of −1.75 V_SCE that is lower than that of pure Mg.¹⁶⁸⁾ In this respect, the size and distribution of noble secondary phases and particles also plays a significant role. If there is a small amount of discontinuous secondary phases, the surrounding Mg base material shows enhanced dissolution while in case of a continuous network of the second phase, this can have a barrier effect and protect the base material underneath.¹⁷⁶⁾

Impurities can also increase the degradation rate of Mg alloys, with Fe being one of the most critical.^177,178)

4.1 Influence of SPD processing on biodegradation

As described above, the electrochemical condition of the material (in terms of overall potential as well as micro-galvanic coupling) and the testing conditions (in terms of evolving passivating surface films) generally have a strong influence on the observed degradation behaviour of Mg alloys. When examining the influence of different SPD methods, it is therefore strongly recommended to compare the SPD processed alloy in the same test set-up to its non-deformed counterpart and in case of in vitro tests to apply physiological test conditions as good as possible (body temperature, physiological pH value, salt concentrations like in the body, and so on).

When examining the literature about degradation of ECAP processed Mg alloys, reports about increasing degradation rates (e.g. ZX40 in NaCl solution,¹⁷⁹⁾ AZ31 in NaCl solution,¹⁸⁰⁾ AZ61 in NaCl solution,¹⁸¹⁾ AZ91D in NaCl solution,¹⁸²⁾ Mg–Zn in NaCl solution¹⁸³⁾), decreasing degradation rates (e.g. pure Mg in simulated body fluid (SBF),¹⁸⁴⁾ Mg–Gd–Nd–Zn–Zr in SBF¹⁸¹⁾) or unchanged degradation rates (e.g. Mg–4Ag in SBF,⁶⁵⁾ Mg–Zn–Ca in SBF⁶⁰⁾) can be found.

There are two main reasons for this large variation in results. The first is the variety of investigated Mg alloys and SPD processing temperatures, the second one is the different testing environments. As discussed previously, SPD processing generally refines the grain structure of Mg alloys and in case of materials containing precipitates, their size and distribution can be changed. From theory it is expected that a high density of grain boundaries and lattice defects like dislocations leads to faster corrosion as these sites have lower density and it is easier to remove atoms there.¹⁸⁵⁾ However, even this depends on the environment. In case of passivating environments (e.g. with low NaCl concentration) the passivation capacity is increased if the grain size is smaller, thus, the corrosion rate is decreased.¹⁸⁶⁾ However, the influences of stored defect energy or other features of the Mg matrix like orientation of the grains (i.e., texture) are generally considered smaller than the influence of second phase particles and micro-galvanic corrosion. In view of this, the amount and distribution of second phase particles in the initial as well as in the SPD processed condition have to be taken into account. For example, a study on AE21 and AE42 Mg alloys¹⁸⁷⁾ showed that for the first alloy the corrosion resistance in NaCl solution was slightly reduced by ECAP while for the second it was significantly increased. The second alloy had higher amount of alloying elements, thus, a higher number of Al₁₁RE₃ second phase particles. ECAP led to a more homogenous distribution of these particles compared to the initial extruded state and therefore an overall decreased degradation rate. Another example is a study on a lean Mg–Zn–Ca alloy processed by D-ECAP.⁶⁰⁾ This alloy exhibits a slow and homogenous degradation behaviour as it contains only Mg₂Ca precipitates which are less noble than the Mg matrix. As D-ECAP processing does not change the size and distribution of the particles, also the degradation rate measured in simulated body fluid is not changed.

The testing environment also has a significant influence on measured degradation rates. In general, the more physiological the solution (this means, it contains not only NaCl but also further salts, glucose or even organic compounds like proteins), the more complex surface layers evolve on the Mg alloys during immersion. A study¹⁸⁸⁾ comparing the degradation behaviour of a LAE442 (Mg–4Li–4Al–2RE) alloy in different media showed that ECAP improved the corrosion resistance in simple NaCl solution as well as in more physiological solutions compared to the extruded condition. The absolute values, however, vary by up to 80%. A further study compared as-cast with ECAP processed ZM21 alloy (Mg–2Zn–1Mn)¹⁸⁹⁾ in different media with and without a buffer to keep the pH value in a physiological range. ECAP samples showed in all conditions faster and more localized corrosion behaviour, however, the difference between the two conditions also depended on the test conditions. The higher degradation rate is ascribed to the crystal defects produced by the process and the change in the distribution of second phase particles, both preventing the formation of a continuous protective layer.

There are also degradation studies on HPT processed Mg alloys. For example, a Mg–2Zn–0.2Ca alloy after HPT¹⁹⁰⁾ showed improved degradation resistance in simulated body fluid (SBF) because a more uniform and compact layer of degradation products was formed. Another study on different binary and ternary HPT processed Mg–Zn, Mg–Ca and Mg–Zn–Ca alloys¹⁶²⁾ showed only rather small variations of the degradation rate in SBF but depending on the alloy as well as the immersion time, HPT processing slightly increased or decreased the degradation rate.

In vivo studies of SPD processed Mg alloys are still very rare. An in vivo study with rabbits of annealed as well as ECAP processed AZ31 alloy⁸⁶⁾ showed that the smaller grain size after ECAP increased the bioactivity of the alloy leading to a faster evolution of protective surface layers and therefore slower overall degradation. In vivo studies on ECAP as well as HPT processed pure Mg, Mg–Ca and Mg–Sr alloys done on rats, showed that both alloys exhibit slower degradation than the pure Mg while there is no significant difference between the two processing methods.^191,192)

5. Summary and Outlooks

This review shows that ultra-fine grained or even nanostructured magnesium alloys obtained by SPD methods can be seriously considered for bioabsorbable medical implants characterized by better mechanical properties than those classically fabricated.

Particularly noteworthy are magnesium alloys containing Zn and Ca because they are characterized by an excellent biocompatibility and demonstrate attractive properties desirable for medical applications without side effects, as in the case of alloys with the addition of Al. On the other hand, magnesium alloys with rare earth additives have the highest strength combined with good degradation resistance, but their content must be controlled to avoid the risk of harmful effects on the human body.

Severe plastic deformation has become a suitable tool for optimizing and controlling the appropriate combination of mechanical properties of magnesium alloys. Although many SPD methods are available, most of the magnesium alloys considered in this paper are generally processed using two of the best-known SPD processes, equal-channel angular pressing (ECAP) and high-pressure torsion (HPT).

This paper comprehensively shows the enhancement of mechanical properties such as ultimate tensile strength, yield strength and hardness of selected magnesium alloys by ECAP and HPT processes. However, the article does not discuss (and only mentions) the phenomenon of superplasticity, which has been proven for magnesium alloys processed by both SPD methods,¹⁹³⁾ which is very interesting from the point of view of product shaping.

Ultra-fine grained magnesium alloys processed by ECAP have desirable and much higher mechanical properties than those currently used for bioabsorbable implants. This applies to Mg–Zn–Ca alloys in particular, which, in addition to excellent mechanical properties, have a beneficial effect on the formation of new bone tissue during healing. Furthermore, they endure equally good biocompatibility and corrosion resistance in the environment of body fluids. Therefore, they can be an alternative to the current materials for bone fixation implants.

Nanostructured magnesium alloys obtained by the HPT technique are characterized by exceptionally small grains and excellent mechanical properties, especially a high hardness, and unattainable by other methods. Moreover, in several cases the hardness can be further increased by post-HPT ageing. On the other hand, the main disadvantage of materials obtained by the HPT method, apart from the restricted volume of material, is their low plasticity, which can be modified to some extent by heat treatment.

The effect of SPD processing on the biodegradation behaviour and rate is diverse as several things have to be taken into account, especially the influence of the processing on second phase particles and the evolution of protective surface layers during testing, which again depend on the testing environment and time. However, the number of papers reporting decreased or unchanged degradation rates after ECAP or HPT is higher than the number of papers reporting an increase.^185,194)

Due to the high requirements for materials for medical implants, magnesium alloys processed by SPD methods with good mechanical properties probably will require additional heat treatment and an appropriate coating to control the rate of degradation in the environment of body fluids. Further issues are increasing the volume of processed materials by SPD methods and long-term clinical trials that should be carried out in the future.

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