Development of Mg-Based Bulk Metallic Glasses and Applications in Biomedical Field

Abstract

Mg and its alloys have attracted much attention recently as a biomaterial due to their excellent biocompatibility, similar mechanical properties to bone and biodegradability. However, the rapid degradation proved to be challenging to act as an implant. To improve the corrosion resistance and overcome rapid degradation of Mg-based alloys, researchers have been working on the synthesis of Mg-based bulk metallic glasses (BMGs). The first research on Mg-Cu-Y BMG was conducted by Inoue in 1991. Since then, Mg-based BMGs with different systems have been developed. Compared to the crystalline metallic Mg-based implants, the structure of Mg-based BMGs without any regular pattern offers low corrosion reactivity and increase passivity. Mg-based BMGs exhibit a good combination of biological, mechanical and corrosion properties and ease of fabrication. Thus, Mg-based BMGs can be considered an efficient candidate in the biomedical field.

1. Introduction

Bulk metallic glasses (BMGs) can be considered a new group of engineering materials in the field of biomedical materials. As crystallize, the molecules have to find the right place in the lattice and position themselves at the right angle relative to their neighbors. Once the position of the molecule is confused or missing, it will cause defects that affect the comprehensive properties of the material. For glass state, due to fast supercooling the liquid from melting temperature (T_m) to glass transition temperature (T_g) inhibits the crystalline phase to nucleate and grow. Liquid-like disorder is “frozen” to solid BMGs structure due to the fast supercooling rate. This fast-growing new class of materials is known as BMGs. BMGs present more homogeneous structures without microstructural defects like dislocations, grain boundaries, or precipitates, thus exhibiting better and more stable mechanical properties. Comparing to the conventional crystalline metallic counterparts, BMGs exhibit higher strength, lower Young’s modulus, improved wear resistance, good fatigue endurance, and excellent corrosion resistance due to the unique amorphous structures. For this purpose, BMGs have recently attracted much attention for biomedical applications. BMGs such as Ti-based [1, 2, 3], Zr-based [4], Mg-based [5, 6], and Ca-based [7] are popular in the biomedical field. Ti-based and Zr-based BMGs are almost impervious to corrosion in the human body; hence they are suitable for permanent implant. And Mg-based and Ca-based BMGs for temporary due to be easily corroded by body fluids [5, 7]. Compared to other materials, Mg is exceptionally light in weight with a density of around 1.74 g/cm³, and the elastic modulus of Mg is similar to that of natural bone [6]. Further, the ability for magnesium alloys to undergo biodegradation eliminates the requirement for a second surgery to remove the implant. Here, Mg seems to be the promising candidate for biodegradable implants. As a natural element in human body, Mg shows good biocompatibility with no systemic inflammatory reaction or effects on cellular blood composition [8]. Additionally, its degradation releases Mg²⁺ ion and is a cofactor in numerous numbers of enzymes. And the products such as MgO and Mg(OH)₂ are expected to be nontoxic to the surrounding tissue. Therefore, researchers viewed Mg-based BMGs as alternative biomaterials for implant applications. Despite the advantages of Mg alloys, the rapid degradation act as a challenge as the implant is unable to retain its structural integrity during the implantation period. To improve the corrosion resistance of Mg alloys, Mg-based bulk metallic glasses have been synthesized and characterized, which can significantly improve the corrosion resistance of Mg-based alloys. In addition to the aforementioned properties, the development of the Mg-based BMGs is of great economic interest as magnesium is abundantly available in land and seawater.

In this chapter, the development of Mg-based BMGs and their application in the biomedical field are introduced. In addition, the high strength, suitable corrosion rate, and biocompatibility of the Mg-Zn-Ca BMGs, which can be considered the most potential system in biomedical field, are described in detail of the mechanical properties, corrosion resistance, and biocompatibility.

2. Development of Mg-based BMGs

2.1 The formation of Mg-based BMGs

The most conventional approach for the development of bulk metallic glasses is through the rapid cooling technique at very low critical cooling rates (<100 K/s). Examples of methods involved in rapid cooling are liquid quenching methods such as injection casting, tilt casting, and melt spinning. The working principle of these processes is that rapid quenching causes the freezing of atoms and does not allow rearrangement and diffusion of an atom. Thus, materials can maintain the glassy structure because of no order or short-range order as rapid quenching freezes the liquid-like atomic arrangement. Cooling rate and degree of crystallization have an important role in dictating the microstructure of the Mg-based BMGs. So the critical cooling rate determines the critical casting thickness—the maximum thickness for glass formation. One can speak of a material’s glass-forming ability (GFA) as being either inversely proportional to its critical cooling rate or proportional to its critical casting thickness [8].

The formation of BMG can be concluded as three criteria [9, 10]: (1) Requiring a multicomponent alloy with three or more elements; (2) Exhibit mismatch of atomic size ratio greater than 12% among major constituents; (3) Having a negative heat of mixing between major constituents. This rule signifies the range of glass forming composition, which often tallies with low melting temperature and eutectic region. It reduced the glass transition temperature according to Eq. (1) as follows [11]:

$$T_{\mathit{rg}}=\frac{T_g}{T_m}$$E1

T_g is the glass transition temperature and T_m is the melting temperature of the alloy. This is for easy glass-forming alloys where a reduced glass transition temperature ($T_{\mathit{rg}}$) is normally equal to or higher than 0.6.

2.2 Introduction of Mg-based BMGs

In 1980s, Mg-based metallic glass has been studied. The initial research on Mg-based metallic glass mainly focused on the physical properties such as glass-forming ability, resistivity, hall-coefficient, thermopower, and electron transport properties. In 1991, Inoue et al. [12] first developed Mg-based BMGs, which proved the system of Mg-Cu-Y BMG. It exhibited excellent GFA and a wide supercooled liquid region. The diameter amorphous cylinder (D_c) of prepared Mg-Cu-Y BMG is 1.5 mm with the length of 50 mm (as shown in Figure 1).

Figure 1.

Mg-Cu-Y BMG produced by low-pressure casting into a copper mold [12].

Copper mold casting method is the most conventional of preparing the Mg-based BMG [13], however, the bulkier a sample is, the more slowly its interior cools. The sample size thus will be limited. With the development of technology, more and more Mg-based BMG systems have been developed successfully with various methods. And these Mg-based BMGs exhibit a wide supercooled liquid region before crystallization and have large GFA. The system of Mg-Ni-Nd BMG with a thickness as large as 3.5 mm is produced through chill-block melt-spinning [14]. Mg-Cu-Ni-Ag-Zn-Y-Gd BMG with a diameter of 14 mm was successfully fabricated by conventional copper mold casting method [15]. Mg-Cu-Ni-Gd BMG systems with a maximum diameter of 2.5–5.0 mm were prepared by copper mold casting (Figure 2) and exhibited high strength above 900 MPa [16]. Some Mg-based BMG systems and preparation methods are summarized in Table 1.

Figure 2.

Outer shape and surface appearance of Mg-Cu-Ni-Gd BMG rods with diameters of 2.5, 4.0, and 5.0 mm, respectively [16].

System	Production method	Diameter (mm)	Refs.
Mg-Cu-Y	Copper mold casting	4	[12]
Mg-Ni-Nd	Chill-block melt-spinning	3.5 (thickness)	[14]
Mg-Cu-Ni-Gd	Copper mold casting	2.5–5.0	[16]
Mg-Cu-Y-Zn	Induction furnace under an inert atmosphere	3	[17]
Mg-Cu-Ni-Zn-Ag-Y-Gd	Copper mold casting	14	[15]
Mg-Cu-Ni-Zn-Ag-Y	Copper mold casting	9	[14]
Mg-Cu-Ag-Er	Copper mold casting	8	[18]
Mg-Ni-La	Injection casting	2.5 ∼ 4.0	[19]
Mg-Zn-Ca-Sr	Copper mold casting	4 ∼ 6	[20]
Mg-Y-Cu-Ag-Pd	Water quenching	12	[21]
Mg-Zn-Ca	Spark plasma sintering	15	[22]

Table 1.

The summary of Mg-based BMGs systems with different diameters by various preparation processes.

2.3 Structure of Mg-based metallic glasses

Metallic glasses are of utmost importance because of their exceptional chemical, mechanical, and physical properties. They are amorphous in structure without any regular pattern, such as crystalline metallic alloys. Their disorder structure can be realized from a wide range of glassy compositions because of the absence of specific stoichiometries. This is also favorable for the microscopic tuning of properties under a certain range through the optimization of the glass transition composition. Take the Mg-Cu-Y-Ni metallic glass as an example, researchers described the structure of Mg-Cu-Y-Ni metallic glass by experimental and modeling methods. Mg-Cu-Y-Ni metallic glass was prepared in the form of ribbons with a thickness of 0.08 mm and a width of 10 mm by the melt spinning (MS) technique. And they also found that a small amount of Ni could improve the glass-forming ability of an Mg-Cu-Y alloy. As shown in Figure 3, the structure of Mg-Cu-Y-Ni metallic glass is presented by the 3D atomic configuration obtained from Monte Carlo (RMC) modeling of the random configuration of 8000 atoms. It can be seen that the distribution of atoms is not completely homogeneous. And the Cu and Y atoms segregate in some areas, indicating the formation of local ordering with more and less dense regions.

Figure 3.

Structure of Mg-Cu-Y-Ni metallic glass [23].

Metallic glass has a special microstructure, which is different from crystalline alloy. The X-ray diffraction (XRD) pattern and Differential Scanning Calorimetry (DSC) curve of Mg-Cu-Y-Ni metallic glass are displayed in Figure 4. It reveals a typical broad diffraction peak that is centered at about 43° and indicated the formation of an amorphous phase. The sample was heated from room temperature to 600 K at a heating rate of 20 K/min. It can be observed that the Mg-Cu-Y-Ni metallic glass exhibits an endothermic effect of the glass transition followed by a distinct exothermic peak. The detected effects confirm the amorphous structure of the studied sample and allowed the glass transition temperature (T_g = 420 K), the onset crystallization temperature (T_x = 467 K), and the peak crystallization temperature (T_p = 474 K) to be determined. The supercooled liquid region(ΔT_x = T_x − T_g) of Mg-Cu-Y-Ni metallic glass can be calculated at about 54 K.

Figure 4.

(a) XRD pattern and (b) DSC curve of Mg-Cu-Y-Ni metallic glass [23].

3. Applications of Mg-based BMGs in the biomedical field

The development of biodegradable BMGs is still in its infancy. Since the inception of an Mg-Zn-Ca BMG in 2005 [24], only Mg-based biodegradable BMGs have been successful as potential contenders for temporary implant applications. For crystalline Mg and its alloy, the rapid degradation rate will precipitate a large amount of hydrogen. Hydrogen evolution that cannot be removed will form pores in the tissue, affecting tissue healing. In contrast to crystalline Mg, the potentially much greater range of alloying elements in an amorphous single-phase structure allows the production of particular metallic glasses with significantly improved corrosion characteristics. This implies that glassy Mg alloys may exist for hydrogen evolution during degradation, which is significantly reduced or even prevented completely [25]. And it is very necessary for the biomedical application.

Recent research into Mg-based BMGs has found that they have higher strength and lower Young’s modulus than pure Mg and conventional Mg alloys [26]. In addition, the system of Mg-Zn-Ca BMG presents more uniform corrosion morphology than conventional crystalline Mg alloys, has much lower corrosion rate, and shows higher cell viability than conventional crystalline Mg [5, 27, 28]. All of Mg, Zn, and Ca are essential elements for the human body, the completely nontoxic Mg-Zn-Ca BMG thus has attracted huge attention in biomedical applications.

3.1 Mechanical properties

Although Mg-Zn-Ca BMG shows great potential biomedical applications, however, Mg-based BMGs are among the most brittle BMGs with almost no plasticity [29] at room temperature (Figure 5); this limits their potential application [30]. To improve their mechanical properties, it is always preferred to use the particulate-reinforced Mg-based BMGs composite [31, 32, 33]. Metallic elements, alloys, polymers, ceramics, and oxides can be selected as the second phase. However, from the previous research, it seems that metal elements are the best reinforcer. Xie et al. [34] introduced Fe particles into Mg-Zn-Ca MG to fabricate composites by ball milling process. After sintered by spark plasma sintering (SPS), the samples exhibit a microstructure of core-shell. The scanning electron microscope (SEM) of Mg-Zn-Ca BMGs and composites with Fe particles are displayed in Figure 6. It can be observed that Fe particles distribute on the surface of Mg-Zn-Ca forming a composite material with a core-shell structure where Fe as the second phase is located at boundaries of amorphous regions.

Figure 5.

Compression properties of Mg₆₆Zn₃₀Ca₄ and Mg₇₀Zn₂₅Ca₅ BMG [29].

Figure 6.

The SEM images of Mg-Zn-Ca BMG and Mg-Zn-Ca/Fe BMG composite with different Fe after SPS processing, respectively. High resolution of Mg-Zn-Ca/Fe BMG composite is inserted in (d) and (e) mapping images for the element Mg, Zn, Fe, and Ca of Mg-Zn-Ca/Fe BMG composite (with 10% of Fe) [34].

Figure 7 shows the compressive strength of Mg-Zn-Ca BMG and composites with different addition contents of Fe. The stress of Mg-Zn-Ca BMG is 574 MPa while those of the composites are 679 MPa, 738 MPa, 765 MPa, and 641 MPa for the composites with addition of 5%, 8%, 10%, and 12% (mol.%) Fe, respectively. Due to introducing Fe into Mg-based BMG, the composites exhibit an enhanced compressive strength. Therefore, “ex-situ” addition of the elements into metallic glass (MG) matrix can be regarded as an effective and attractive way to improve the mechanical strength. However, the addition of Fe did not improve the plasticity of Mg-Zn-Ca BMG.

Figure 7.

The compressive stress-strain curves of Mg-Zn-Ca BMG and Fe@Mg-Zn-Ca BMG composites at room temperature [34].

It is worth mentioning that Mg-Zn-Ca BMG exhibits good plasticity with the addition of rare earth elements [35]. Figure 8 displays the compressive stress-strain curves of Mg-Zn-Ca BMG with rare earth Y. The Mg-Zn-Ca glassy alloy also exhibits linear elastic behavior and it only fails in the absence of macroscopic yielding and plastic strain. However, the addition of Y can significantly improve fracture strength. Fracture strength increases from 550 to 1012 MPa (80% increment) with the addition of 1 at.% Y, which is four times higher than that of traditional magnesium alloys and 26% higher than Mg-Zn-Ca metallic glass. The Mg-Zn-Ca-Y Mg-Zn-Ca BMG also exhibits an excellent capacity for plastic strain, above 3.1%, which is 2.5 times higher than the capacity of Mg-Zn-Ca metallic glass. However, the biocompatibility of rare earth elements is still unclear. So rare earth elements as biomedical materials are still very controversial.

Figure 8.

Engineering strain-stress curves of the Mg_69−xZn₂₇Ca₄Y_x: (a) x = 0, (b) x = 1, (c) x = 2 at.% [35].

More mechanical properties of Mg-Zn-Ca BMG and their composites via different methods are summarized in Table 2. It is worth noting that the mechanical properties of Mg-Zn-Ca BMG are differently prepared by different methods. The mechanical properties of Mg-Zn-Ca BMG composites are highly dependent on the inherent properties of the second phase of reinforcement and its volume fraction. However, the plasticity of Mg-Zn-Ca BMG is still not ideal and needs more researches to overcome it.

System	${\sigma }_f$	${\sigma }_y$	A (%)	Hv	Ref
Mg60Zn35Ca5	571	—	—	—	[31]
Mg67Zn28Ca5	432	—	—	—	[31]
Mg68Zn28Ca4	828		1.28	—	[36]
Mg70Zn25Ca5	642	—	0.4	—	[36]
Mg80Zn15Ca5	513		3.58		[36]
Mg66Zn30Ca4	716–854	—	—	—	[37]
Mg71Zn25Ca4	672–752	—	—	—	[37]
Mg60Zn34Ca6	—	888	—	296 ± 25	[38]
Mg73Zn23Ca4	—	636	—	212 ± 19	[38]
Mg66Zn30Ca4	378–587	—	—	257–263	[22]
Mg-Zn-Ca-Cu	583–797	—	—	—	[39]
Mg60Zn35Ca5/Ti	807	—	—	—	[31]
Mg67Zn28Ca5/Ti	690	—	—	—	[31]
Mg66Zn30Ca4-xSrx	787–848	—	—	2.45–2.51	[20]
Mg-Zn-Ca-Ga	>600	—	—	—	[40]

Table 2.

The mechanical properties of Mg-Zn-Ca BMG and their composites.

${\sigma }_f$: Compressive fracture strength, ${\sigma }_y$: compressive yield strength, A: elongation, H_v: microhardness.

3.2 Corrosion properties

Crystalline Mg and its alloys show poor corrosion resistance in body liquid, which leads to a fast degradation rate and an uncontrollable hydrogen release. Because of the active chemical reaction between the alloys and the body fluid, the biomedical application of Mg-based alloys is extremely limited. Due to the unique atomic configuration in the absence of translational and rotational symmetry down to sub-nanoscale, Mg-based BMGs exhibit better corrosion resistance compared with crystalline alloys. However, they still cannot meet the clinical needs. Cardiovascular and orthopedic applications are the most important applications of Mg-based alloys, however, for the bioabsorbable cardiovascular and orthopedic applications, it should degrade slowly in the early stage of implantation. Vascular and orthopedic remodeling is generally completed in 90 and 120 days, respectively. After that, it should gradually degrade at an appropriate rate so that the degradation products do not cause adverse reactions in the surrounding tissues. Generally speaking, a full absorption of the cardiovascular and orthopedic applications is expected within two years. Obviously, it cannot be achieved for the current Mg-based alloys and BMGs. Therefore, improving the corrosion resistance of Mg-based BMGs is still the biggest challenge. The effective ways to improve the corrosion resistance of Mg-based BMGs are introducing other alloying elements [40] and preparing surface coating [41].

The researchers introduced Ga into Mg-Zn-Ca and investigated the corrosion behavior in Figure 9. According to the results of electrochemical tests and immersion tests, it can be seen that with the increase of Ga addition, Mg-Zn-Ca-Ga metallic glasses have significantly higher corrosion potential and lower corrosion current density corresponding to higher corrosion resistance. The addition of element Ga can help to form a dense oxide or hydroxide film on the metallic glasses. Ga thus improves the anticorrosion performance of Mg-Zn-Ca metallic glass.

Figure 9.

(a) Open circuit potential (OCP) curves, (b) Nyquist plots, and (c) potential dynamic polarization (PDP) curves of different Mg-Zn-Ca-Ga metallic glasses (Ga0, Ga0.25, Ga0.50, Ga0.75, Ga1.00, and Ga1.25) tested in hanks’ solution at 37°C, (d) electrical equivalent circuit (EEC) used in the fitting of EIS result [40].

The methods for preparing the coating mainly include plasma spraying, pulsed laser deposition, sol-gel method, electrochemical deposition, and so on. Among these methods, electrochemical deposition is popular because of the relatively low deposition temperature, controllability for a coating composition, and fast deposition rate. In this part, take the electrochemical deposition method as an example. The researchers synthesized a series of calcium phosphate (CaP) coatings on Mg-Zn-Ca metallic glass via the electrodeposition method. Among all bioactive coatings, calcium phosphate (CaP) compounds exhibit outstanding biocompatibility and low toxicity. However, the coating thickness, chemical composition, and morphology will strongly influence the biological activity and degradation rate. Different voltages will influence the coating thickness and morphology. In this work, different voltages including −3.3 V, −3.5 V, and − 3.7 V are carried out to prepare coatings with different coating thicknesses. Figure 10 displays the SEM morphologies of surface and cross-section of electrodeposited coating under different voltages. The flake-like morphology CaP coatings are formed on the substrate under different electrodeposited voltages. The thicknesses of CaP coating are 3.6 ± 1.6 μm, 8.3 ± 2.1 μm, and 31.1 ± 6.8 μm under different electrodeposited voltages of −3.3 V, −3.5 V, and −3.7 V, respectively. The result indicates the CaP coating is increased with the increase of electrodeposited voltage. Electrodeposited CaP coating can effectively inhibit the reaction between sample surface and liquid environment.

Figure 10.

The SEM morphologies of samples: (a1) (a2) CaP-3.3 V, (b1) (b2) CaP-3.5 V, and (c1) (c2) CaP-3.7 V and the cross-section morphologies of (a3) CaP-3.3 V, (b3) CaP-3.5 V, and (c3) CaP-3.7 V. the corresponding EDS results are provided in (a4), (b4), and (c4) [42].

Figure 11 shows the electrochemical results of different electrodeposited CaP coating on Mg-Zn-Ca metallic glass in different samples. From Figure 11(a), the open-circuit potential (OCP) of CaP coating samples is higher than that of Mg-Zn-Ca metallic glass, which indicates higher corrosion potential. In results of fitted electrochemical impedance spectroscopy (EIS) data (Figure 11(b) and (d)), comparing with the electrodeposited coating samples, Mg-Zn-Ca metallic glass exhibits a significantly smaller capacitive loop and polarization resistance value (3516 Ω cm²). The polarization resistance of the sample CaP-3.3 V increased to 5701 Ω cm², which is slightly higher than that of the Mg-Zn-Ca metallic glass substrate. When the electrodeposition voltage was 3.5 V, the polarization resistance of sample CaP-3.5 V reaches the highest value of 27,110 Ω cm². As the electrodeposition voltage further increased to 3.7 V, the polarization resistance of sample CaP-3.7 V dropped to 14,454 Ω cm². According to the EIS fitting parameter (R_p), the corrosion resistance of samples CaP-3.3 V, CaP-3.5 V, and CaP-3.7 V are 1.6 times, 7.7 times, and 4.1 times higher than that of the Mg-Zn-Ca metallic glass. Figure 11(c) shows that the value of Mg-Zn-Ca metallic glass is significantly lower than that of electrodeposited coating samples, this result indicates that the electrodeposited CaP coating can effectively inhibit the electrochemical reactions on the sample surface, thus reducing the corrosion rate.

Figure 11.

The electrochemical results of different samples. (a) OCP curves, (b) Nyquist plots, (c) PDP curves, and (d) bode Nyquist plots tested in Hanks’ solution [42].

Although a great deal of research has been carried out on the corrosion of Mg-based BMGs, it is worth mentioning: the research on the corrosion mechanism of Mg-based BMGs is still in the early stage, so it is a great challenge to further study the corrosion mechanism of degradable medical Mg alloys.

3.3 Biocompatibility

For implant materials, biocompatibility is the most important evaluation index. It is necessary to ensure that the material is completely nontoxic before further research is necessary. Numerous studies have been conducted on the biocompatibility of Mg-Zn-Ca BMGs both in vivo and in vitro.

The system of Mg-Zn-Ca BMGs has been proven with excellent biocompatibility. Figure 12 shows the animal studies in the abdominal walls and cavities (two tissue types apiece) of domestic pigs to evaluate the tissue reactions of the Mg glass during degradation and the hydrogen evolution in vivo. Mg-Zn-Ca BMGs discs, together with a crystalline reference Mg alloy (WZ21), were implanted and analyzed after 27 and 91 days. Around both materials, a fibrous capsule, typical of wound-healing processes after implant surgery, has developed in all four different tissue types (see the white arrows). Within the capsules, however, gas cavities formed by hydrogen evolution (as indicated by the black arrows in Figure 12(b) and (d)) are only observed around the crystalline Mg discs. In contrast, no tissue-imprinted hydrogen gas cavities have formed in the histological preparations of the glassy Mg-Zn-Ca samples (see Figure 12(a) and (c)). As no inflammatory reaction was observed for any of the implants, the animal tests show that the biocompatibility of Mg-Zn-Ca metallic glass as good as the crystalline Mg alloy [25].

Figure 12.

Animal studies of Mg-based BMG in comparison with a crystalline Mg alloy reference sample. Mg-Zn-Ca BMG (a) and (c) and crystalline Mg alloy reference (WZ21) (b) and (d) [25].

The biocompatibility of Mg-Zn-Ca BMG in vitro has also been studied. According to ISO 10993-5: 2009, Mg-Zn-Ca BMG and Mg-Zn-Ca (with 10% content of Fe) BMG composite specimens immersed into high glucose Dulbecco’s modified eagle medium (HGDMEM) (Hyclone) supplemented with 10% fetal bovine serum (FBS and Gibco) and penicillin/streptomycin (PS, Gibco). Raw264.7 cells were seeded in a 24-well plate with a concentration of 1.6 × 104/ML and replaced with a prepared extract after 24 h and cultured for 1, 3, and 5 days. Cells were washed with PBS three times, 500 μl mixed dye was added to each well, and cells were cultured in a cell incubator for 15 min. Live/dead BacLight (Invitrogen) was used to detect the live or dead Raw264.7 cells. Then observe the cell fluorescence by fluorescence microscopy. Green respects living cells and red means dead cells.

Figure 13(a) shows the numbers of RAW264.7 cells culturing in extract of Mg-Zn-Ca BMG and Mg-Zn-Ca BMG composites with addition of Fe for one day, three days, and five days. The number of cells increases significantly from day one to day five, which proves that the two materials are nontoxic and with good biocompatibility for RAW264.7 cells. Figure 13(b) presents the quantitative results in Figure 13(a). With the addition of Fe, the absorbance value becomes higher proves the addition of Fe is beneficial to cell proliferation.

Figure 13.

(a) Live cell staining assay and (b) cell proliferation after culturing RAW 264.7 measured for one day, three days, and 5 days [34].

Figure 14(A) shows the representative gross sagittal image of a harvested rabbit femur exposing the surgical implant site. All the implanted rods were found to be well-fixed inside the bone tunnel with no evidence of loosening. Before they were sent for histological sectioning, all rods were removed manually. The results show remarkable new bone formation surrounding the Mg-Zn-Ca BMGC and Ti6Al4V alloy at 12 and 24 weeks after implantation (Figure 14(B)). In contrast, many osteoblasts are found surrounding the Mg-Zn-Ca BMG and Ti6Al4V alloy rods compared to PLA. No inflammatory cells such as leukocytes or macrophages are localized at the implant sites, signifying good biocompatibility for Mg-Zn-Ca BMG. All rabbits demonstrated good health before sacrificing with no evidence of local (surgical site) or systemic inflammation.

Figure 14.

(A). Gross image of harvested rabbit femur indicating the surgical implanted site of Mg-Zn-Ca BMG, Ti6Al4V, and PLA. The fixation rod was first removed before sent for histologic section. (B) Histological images of the implanted site at 12 weeks and 24 weeks. Black arrows indicate new bone formation [43].

A conclusion can be drawn from the above results, no matter whether in vivo or in vitro, Mg-Zn-Ca BMG shows superior biocompatibility. With an improved biodegradation rate, excellent biocompatibility, and most importantly, osteogenic ability, Mg-based BMG has great potential for future surgical implant development and application.

4. Conclusions

A large number of studies have shown that Mg-based bulk metallic glasses had a huge application prospect in the field of biodegradable biomaterials. For biodegradable applications, they will degrade gradually in human body after completing their temporary mission (would dissolve completely upon fulfilling the mission of fixing or supporting) during which arterial/bone remodeling and healing would occur. They can be used as a scaffold and bioresorbable implants because they offer an excellent combination of mechanical properties, low degradation rate, and biocompatibility. Through different methods to improve the mechanical properties and corrosion resistance of Mg-based bulk metallic glasses, great progress has been made. However, brittle fracture and fast corrosion rate of metallic glass have not been completely solved, which is the biggest obstacle limiting its application. Despite better resistance to degradation, more suitable mechanical properties, and biocompatibility, the development of magnesium-based bulk metallic glasses for biomedical applications is still in the stage of infancy. From the processing point of view, challenges remain in obtaining a bulk sample that is a sample that is sufficiently large enough to be of use. What is certain, however, is that the biocompatibility of Mg-Zn-Ca metallic glass system has been fully recognized by implanting in animals experiments. Although initial degradation studies showed promising results compared to their crystalline counterparts, the mechanism behind their in vivo degradation is not yet well understood. Mg-based bulk metallic glasses certainly are promising candidates for future biomedical applications. Breakthrough in processing methods to obtain a sufficiently large bulk metallic glass without compromising the amorphous structure should pave the way for accelerated research as a substitute for current biomaterials and in targeted engineering applications.