Review_bmg_biomaterials.pdf

Acta Biomaterialia 36 (2016) 1–20

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Review article

Recent advances in bulk metallic glasses for biomedical applications H.F. Li, Y.F. Zheng ⇑ Department of Materials Science and Engineering, College of Engineering, Peking University, Beijing 100871, China

a r t i c l e

i n f o

Article history: Received 23 October 2015 Received in revised form 17 March 2016 Accepted 31 March 2016 Available online 1 April 2016 Keywords: Bulk metallic glasses Liquid metal Advanced materials Biomaterials Clinical applications

a b s t r a c t With a continuously increasing aging population and the improvement of living standards, large demands of biomaterials are expected for a long time to come. Further development of novel biomaterials, that are much safer and of much higher quality, in terms of both biomedical and mechanical properties, are therefore of great interest for both the research scientists and clinical surgeons. Compared with the conventional crystalline metallic counterparts, bulk metallic glasses have unique amorphous structures, and thus exhibit higher strength, lower Young’s modulus, improved wear resistance, good fatigue endurance, and excellent corrosion resistance. For this purpose, bulk metallic glasses (BMGs) have recently attracted much attention for biomedical applications. This review discusses and summarizes the recent developments and advances of bulk metallic glasses, including Ti-based, Zr-based, Fe-based, Mg-based, Zn-based, Ca-based and Sr-based alloying systems for biomedical applications. Future research directions will move towards overcoming the brittleness, increasing the glass forming ability (GFA) thus obtaining corresponding bulk metallic glasses with larger sizes, removing/reducing toxic elements, and surface modifications. Statement of Significance Bulk metallic glasses (BMGs), also known as amorphous alloys or liquid metals, are relative newcomers in the field of biomaterials. They have gained increasing attention during the past decades, as they exhibit an excellent combination of properties and processing capabilities desired for versatile biomedical implant applications. The present work reviewed the recent developments and advances of biomedical BMGs, including Ti-based, Zr-based, Fe-based, Mg-based, Zn-based, Ca-based and Sr-based BMG alloying systems. Besides, the critical analysis and in-depth discussion on the current status, challenge and future development of biomedical BMGs are included. The possible solution to the BMG size limitation, the brittleness of BMGs has been proposed. Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Contents 1. 2.

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Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Ti-based bulk metallic glasses for biomedical implants and devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 2.1. Mechanical properties of biomedical Ti-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 2.2. Corrosion behavior of biomedical Ti-based BMGs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 2.3. Biocompatibility of biomedical Ti-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Zr-based bulk metallic glasses for biomedical implants and devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 3.1. Mechanical properties of biomedical Zr-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 3.2. Corrosion behavior of biomedical Zr-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 3.3. Biocompatibility of biomedical Zr-based BMGs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Fe-based bulk metallic glasses for biomedical implants and devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 4.1. Mechanical properties of biomedical Fe-based BMGs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 4.2. Corrosion behavior of biomedical Fe-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10 4.3. Biocompatibility of biomedical Fe-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10

⇑ Corresponding author. E-mail address: [email protected] (Y.F. Zheng). http://dx.doi.org/10.1016/j.actbio.2016.03.047 1742-7061/Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

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5.

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7.

Biodegradable bulk metallic glasses designed for temporary implants and devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1. Biodegradable Mg-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2. Biodegradable Ca-based and Sr-based BMGs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2.1. Biodegradable Ca-based BMGs. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2.2. Biodegradable Sr-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3. Biodegradable Zn-based BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Future research directions and challenges for BMGs as potential biomaterials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1. BMG composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2. BMG foams . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.3. Removing toxic and noble alloying elements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.4. Improving BMG critical sizes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5. Surface modification of BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5.1. Surface modification of non-biodegradable BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5.2. Surface modification of biodegradable BMGs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.6. Metallic glass coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Concluding remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction With the improvement of the living standards and the development of science and technology, biomaterials have developed rapidly and helped in improving the quality of life and longevity of human beings greatly. Metallic biomaterials have the longest history among the various kinds of biomaterials. It has been reported that people began to use metal dental implants thousand years ago, which can be tracked back to 200 CE [1]. Stainless steel, vitallium (Co-Cr alloys), pure Ti and Ti alloys, pure Zr and Zr alloys are widely used as artificial hip joints, cardiovascular stents, artificial knee joints, bone plates, dental implants, etc. Besides, pure Mg and Mg alloys, pure Zn and Zn alloys have been studied and developed as biodegradable materials aiming to be useful in the clinical cases that need temporary supporting or fixation (such as plates and screws for bone fracture fixation, stents for cardiovascular repair), without second operation to be removed after finishing their functions [2,3]. However, these conventional crystalline metallic alloys have disadvantages such as low strength, high elastic modulus, low wear resistance, prone to crevice corrosion, pitting corrosion as well as stress corrosion cracking (SCC) and high cycle fatigue failure, incompatibility with X-ray or magnetic

Fig. 1. Illustration of the relationship of the three materials, i.e. bioglasses, biometallic alloys and biomedical BMGs.

10 10 12 12 14 15 16 16 16 17 17 17 17 17 17 18 18 18

resonance imaging, which cause various problems in clinical application [4]. For instance, poor corrosion resistance can cause relatively high toxic ion release, such as Ni +, Cr3+ and Co2+ into human and it is well known that these ions usually lead to adverse reactions when their concentrations rise above certain thresholds. Most of the conventional metallic biomaterials possess much higher elastic modulus compared to that of human bone. And the mismatching modulus can cause stress shielding of bone, resulting in bone resorption and loosening of the implant after a period of implantation [5]. Further development of novel biomaterials, that are much safer and of much higher quality, in terms of both biomedical and mechanical properties, are therefore of great interest for both the research scientists and clinical surgeons. Efforts to discover alternatives to conventional biometals led to the discovery of bioglass in the early 1970s [6], since which time various kinds of bioglasses have been developed. Key bioglass systems include the typical S45P7 (SiO2-CaO-Na2O-P2O5-B2O3), 45S5 (SiO2-CaO-Na2O-P2O5) and S52P3 (SiO2-CaO-Na2O-P2O5-B2O3Al2O3) systems, all of which have been widely used in biomedical applications, in contexts ranging from otology to cancer treatment [7]. Unfortunately, most of the bioglasses developed did not represent a real alternative to biometals, since they had unsatisfactory mechanical properties and thus remained unsuitable for clinical applications that need place them in load-bearing sites. Metallic glasses (MGs), also known as amorphous alloys and liquid metals, emerged as a newcomer to the clubs of metallic materials in 1960, with the formation of the first metallic glass of Au75Si25 being reported in Nature journal by Duwez P’s research team at Caltech, USA [8]. MGs have excellent physical and chemical functions such as high toughness and corrosion resistance, but the limited specimen sizes originally achievable (normally in the range of microns) severely restricted the study and application of MGs. These limitations were overcome by the development of bulk metallic glasses (BMGs) with much lower critical cooling rates (<100 K/s). BMGs had received a lot of attention during the last two decades for their high strength and elasticity which are the consequences of their amorphous structure and concomitant lack of dislocations and associated slip planes. From the early 1990s, a series of new kinds of BMGs with the multicomponent chemistry and excellent glass forming ability (GFA) have been discovered in Zr-, Mg-, La-, Pd-, Ti-, and Fe-based systems by various solidification methods. These BMGs exhibited strong resistance to crystallization in the super-cooled liquid state and demonstrated excellent mechanical and physical properties.

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H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20 Table 1 Comparison of bioglasses, biometals and biomedical bulk metallic glasses. Systems

Chemical composition

Classification

Bioglasses

Na2O-CaO-SiO2P2O5 oxides

Biodegradable: 45S5, phosphate-based glasses, Amorphous borate-based glasses Non-biodegradable: 60S3.8, 65S

Biometals

Ti, Zr, Fe, Mg, Ca, Biodegradable: Mg and Mg alloys, Fe and Fe Zn, Sr metals alloys, W Non-biodegradable: Ti and Ti alloys, 316L SS, 304SS, Zr and Zr alloys

Biomedical Ti, Zr, Fe, Mg, Ca, Biodegradable: Mg-based BMGs, Ca – based BMGs Zn, Sr metals BMGs, Zn-based BMGs, Sr-based BMGs Non-biodegradable: Ti-based BMGs, Zr-based BMGs, Fe-based BMGs

Microstructure Preparation methods Sol-gel, melting

Mechanical properties

Current/potential clinical applications

Low strength, low Young’s Modulus

Dental repairs, fillers, middle ear ossicular replacements

Crystalline

Conventional High strength, high casting Young’s Modulus

Dental implants, stents, orthopedic surgeries, sutures

Amorphous

Modified casting

Dental implants, fillers, stents, orthopedic surgeries, sutures

Fig. 2. Mechanical comparisons of conventional bioglasses, biometallic alloys and the novel developed biomedical BMGs.

As shown in Fig. 1, BMGs combine the microstructure of the conventional amorphous bioglasses and the elemental metal compositions of the conventional crystalline biometallic alloys. As a result, BMGs represent a class of materials that may combine the physical and chemical properties of both bioglasses and conventional crystalline metallic materials. A quick comparative review of characteristics and properties for these three classes of materials underscores the promise of BMGs.

High strength, low Young’s Modulus

Table 1 compares conventional bioglasses, biometals and biomedical BMGs according to chemical composition, microstructure, preparation methods and mechanical properties. Furthermore, Fig. 2 illustrates the compression strengths and Young’s modulus for these three classes of materials. Table 1 and Fig. 2 demonstrate that bio-glasses have low strength and low Young’s moduli while conventional metallic biomaterials have high strength and high Young’s moduli. In contrast, biomedical BMGs, with combined characteristics of bioglasses and biometals, show high strength and low elastic moduli making them, in theory, highly suitable for biomedical applications. In particular, the biomedical BMGs’ extremely high elastic limit of 2% compares favorably with bone’s elastic limit of 1%, suggesting that biomedical BMGs would be unique in their ability to flex elastically with the natural bending of the bones and would consequently distribute stresses more uniformly than current materials, minimizing stress concentrations, reducing stress shielding effects, and thus achieving faster healing rates. Thanks to the unique properties of biomedical BMGs, the BMG bone screws could have a thinner shank and deeper threads thus yielding greater holding power to the fracture bones. Compared to conventional 316L SS stents, biomedical BMG stents would require only one third of the cross-section of the strut and would have more than five times the deflection [9]. In the last decades, some of the BMGs were developed specially for biomedical purpose, and both in vitro and in vivo testing had been done to evaluate the feasibility as biomaterials. Several biomedical implants and devices made out of BMGs have been designed and developed in the past decades. As illustrated in

Fig. 3. Illustration of biomedical implants and devices made out of BMGs. (a) Commercial martensitic steel surgical blade coated with ZrCuAlAgSi BMG film (left) and ZrCuAlAgSi BMG surgical blade (right). (b) The ezlase Diode Dental Laser System, from Biolase Technology, uses Liquidmetal in its housing. (c) BMG medical stapling anvils. (d) Liquidmetal Alloys in Minimally Invasive Medical Devices.

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Table 2 Summary of biomedical Ti-based BMG systems and their mechanical properties. Chemical composition (at.%)

Preparation method

Critical diameter/thickness (mm)

Compressive fracture strength (rf) (MPa)

Young’s modulus

Specific strength (kN m/kg)

Vickers Reference hardness (Hv) (kg mm 2)

Ti40Zr10Cu36Pd14

Arc-melting/copper mold casting

6

1950

82

270

—

[19]

Ti75Zr10Si15 and Ti60Nb15Zr10Si15

Arc-melting/ejecting onto a 30 lm thick ribbons single copper

>2000

—

—

790 and 660

[21]

Ti40Zr10Cu38Pd12

Arc-melting/copper mold casting

—

—

Around 100

—

765 (7.5 GPa)

[22]

Ti45Zr10Cu31Pd10Sn4

Spark plasma sintering (SPS)

15

2060 (sintering at 643 K)

—

310 (calculated from the data of strength and density)

—

[20]

Ti45Zr10Cu31Pd10Sn4

Arc-melting/copper mold casting

4

1970

95

—

650

[23]

Ti45Ni15Cu25Sn3Be7Zr5

Arc-melting/copper mold casting

5

2480

—

—

715

[17]

(Ti0.45Zr0.1Pd0.1Cu0.31Sn0.4) 100 xMx (M: Ta and Nb) (x = 1–5)

Arc-melting/copper mold casting

3–5

1670–2150

97–119

—

—

[24]

(Ti40Zr10Cu36Pd14)100–xNbx (x = 1, 3, 5)

Arc-melting/copper mold casting

2

2050

80

—

—

[25]

(Ti40Zr10Cu38Pd12)100–xNbx (x = 0, 2, 3, 4)

Arc-melting/copper mold casting

Fully amorphous for x = 0 and 1200–2000 partially amorphous for x = 2, 3 and 4 with 2 mm in diameter

100–106

—

612–816 (6–8 GPa)

[26]

Ti40Zr10Cu34Pd14Sn2

Arc-melting/copper mold casting

10

2000–2150

—

—

—

[27]

Ti40Zr10Cu32Pd14Sn4

Arc-melting/copper mold casting

10

2000–2150

—

—

—

[27]

Ti43.15Zr9.59Cu36.24Ni9.06Sn1.96 Arc-melting/copper mold casting

3

2640

—

400

—

[18]

Ti41.5Zr2.5Hf5Cu37.5Ni7.5Si1Sn5

Arc-melting/copper mold casting

6

2000 ± 78

80 ± 12

—

600

[28]

Ti41.5Zr2.5Hf5Cu42.5Ni7.5Si1

Arc-melting/copper mold casting

5

2080

103

—

—

[29]

Ti43.3Zr21.7Ni7.5Be27.5 (LM-010) Arc-melting/copper mold casting

—

1790 (tensile 95 yield strength)

351

—

[16]

Ti41.3Cu43.7Hf13.9Si1.1

Arc-melting/copper mold casting

3

1685

95

—

—

[30]

Ti40Zr25Ni12Cu3Be20

Arc-melting/copper mold casting

—

1932

95

—

530

[31]

Ti47Cu38Zr7.5Fe2.5Sn2Si1Ag2

Arc-melting/copper mold casting

7

2080

100.4 ± 0.1

588 ± 6

[32]

Fig. 3, these biomedical devices include surgical blades with improved sharpness (Fig. 3a) [10], the housing of the iLase (Fig. 3b), which is a handheld pen laser for dental procedures, attracted by the unique strength, ability to be molded to a thin wall thickness, and ornamental finish of BMGs [11], medical stapling anvils (Fig. 3c), considering the top 4 advantages BMGs providing: 1) Superior As-Cast Surface Finishes; 2) Pocket-to-pocket dimensional accuracy; 3) Molded-in Camber for proper alignment to staple cartridge when clamped to tissue; 4) Lot-to-lot variability limited to mold cavity-to-cavity variability [12], and Minimally Invasive Medical Devices (Fig. 3d), with enhanced precision, durability, repeatability and more convenient manufacturing processing [13]. Besides the above mentioned biomedical devices, BMGs are promising candidates as biomedical implants. It has been reported that BMGs could increase the strength, corrosion resistance, biocompatibility and longevity of biomedical implants, which are aimed to be used as cardiovascular stents [14] and

orthopedic implants (such as bone plates, bone screws, articulating surfaces, artificial prostheses, absorbable sutures, dental implants and fillers) [15]. In the design and development of biomedical BMGs, the main chemical compositions of the conventional crystalline biomedical alloys, i.e. Ti alloys, Zr alloys, stainless steels, Mg alloys, Zn alloys were taken as references. Based on this consideration, there are the corresponding BMGs with the same main chemical compositions explored as potential biomaterials, including Ti- based BMGs, Zr-based BMGs, Fe-based BMGs, which are also known as amorphous stainless steels, Mg-based BMGs and Znbased BMGs. In addition, Ca-based BMGs and Sr-based BMGs are developed as potential biodegradable materials as well, considering the nutritional functions of the main alloying elements of Ca and Sr. In the following paragraphs, the recent developments and advances of these BMGs respectively as novel biomedical materials will be summarized and comprehensively reviewed.

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Ti-based BMG and it demonstrates that the alloy exhibits an elastic elongation of about 2.3%, followed by a small serrated plastic elongation of about 0.5%, indicating that the cast alloy possesses certain ductility. The fracture took place along the maximum shear stress plane, which was declined by about 45° to the direction of applied compressive load. The SEM image of the fracture surface showed that the dominant fractographic feature is a smooth and well-developed vein pattern, which is caused by viscous flow of materials in shear band [19]. 2.2. Corrosion behavior of biomedical Ti-based BMGs The study of the corrosion behavior of BMGs is of great importance to understand their chemical and environmental stability. Corrosion processes in humid environment are mostly based on electrochemical reactions. First of all, the corrosion rate of multicomponent glass-forming alloys is determined by the electrochemical reactivity of the main alloying elements. Nevertheless, in more detail, aspects of thermodynamic metastability, unusual atomic structure and chemical homogeneity due to the ideally singlephase nature are decisive as well. Researchers have studied the corrosion behavior of Ti-based BMGs in different kinds of simulated body fluids, including phosphate buffered solution (PBS) [16], Ringer’s solution [21], Hank’s balanced salt solution (HBSS) [22,26], 1 mass% Lactic acid [23,24] and specific simulated body fluid (SBF) [33]. Studies have shown that Ti-based BMGS exhibited passive behavior at the open-circuit potential with a low corrosion rate; a susceptibility to localized corrosion in the form of pitting corrosion; the localized corrosion resistance was statistically equivalent to, or better than, the conventional crystalline biomedical alloys, including 316L stainless steel [16], pure Ti and Ti-based biomedical alloys (such as Ti-6Al-4V) [23,24]. After 15 days immersed in SBF, hydroxyapatite (HA), one of the main chemical compositions of human natural bones and teeth, was reported to have deposited on the Ti-based BMG alloy surfaces [33], which would guarantee the biocompatibility of Ti-based BMGs. 2.3. Biocompatibility of biomedical Ti-based BMGs

Fig. 4. (A) Compressive stress–strain curve of the Ti40Zr10Cu36Pd14 glassy alloy rod with a diameter of 2.5 mm. (B) Compressive stress–strain curves of three Ni-free Zrbased BMGs. (C) Uniaxial compressive true stress–strain curves for four Fe–Cr–Mo– P–C–B BMGs: (a) Fe63Cr3Mo12P10C7B5, (b) Fe64Cr3Mo10P10C10B3, (c) Fe63Cr3Mo10P12C10B2 and (d) Fe71Mo5P12C10B2.

2. Ti-based bulk metallic glasses for biomedical implants and devices As summarized in Table 2, series of Ti-based BMGs have been investigated for biomedical application, ranging from Ni and Be containing Ti-Zr-Ni-Be [16], Ti-Ni-Cu-Sn-Be-Zr [17], Ti-Zr-Cu-NiSn [18] systems to the Ni and Be free Ti-Zr-Cu-Pd [19], Ti-Zr-CuPd-Sn [20] and Ti-Zr-Si [21] systems.

2.1. Mechanical properties of biomedical Ti-based BMGs Mechanical properties of materials have great importance due to their role for applications. As shown in Table 2, Ti-based BMGs have relatively low Young’s modulus (80–120 GPa), high fracture strength (1700–2500 MPa), and excellent specific strength. Fig. 4 (A) shows the typical compressive stress–strain curve of the

Biocompatibility of the selected materials has a great importance as well. It is described as the ability of the material to exist in contact with tissues of the human body without causing a non-acceptable degree of harm. That depends on numerous factors. Firstly, it concerns the human body response to it or the cellbiological activity of the implant. The higher the bioactivity, the higher is the biocompatible with surrounding tissues. And the second main factor is the material degradation in the body environment as a consequence of low wear and corrosion resistance. This could result in a release of the constituent metal ions or particles of the implant alloy system into the body causing several reactions including allergic and toxic ones. Hence, choosing materials with high wear and corrosion resistance as well as including non-toxic and non-allergenic elements is of high significance. Researchers have studied the biocompatibility of Ti-based BMGs via both in vitro cell response (via MTT/CCK8 assay and cell morphology observations) and in vivo animal implants. And results have demonstrated that Ti-based BMGs showed better biocompatibility than the conventional crystalline Ti-6Al-4V and Ti-45Ni alloys for both the in vitro human osteoblast SaOS2 cells [34], murine fibroblast cells (L929 cell and NIH3T3 cell) culture and in vivo beagle dogs implantation [28]. Fig. 5 shows the SEM images of L929 and NIH3T3 cells on crystalline pure Ti and Ti41.5Zr2.5Hf5C u37.5Ni7.5Si1Sn5 (TZHCNSS) BMG surfaces after 4 days of incubation [28]. It can be observed from Fig. 5 that cells could attach very

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well on all sample surfaces with numerous cytoplasmic extensions and filopodia. There is not much difference for the NIH3T3 cells grown on pure Ti and TZHCNSS samples, since cells on both samples have connected to other cells and tiled on the surfaces. Fig. 6 demonstrates the implantation of TZHCNSS BMG and pure Ti samples after 1 month implantation into the mandibles of beagle dogs [28]. It can be observed from Fig. 6 that after one month implantation there is no inflammation observed around

the operation sites. And no bone resorption could be observed for both pure the Ti sample and TZHCNSS samples. There is no significant difference between two samples. Both the pure Ti and TZHCNSS samples are well integrated with the bone tissue. New bone was formed around the implants. The EDS analysis also proved that the bones integrated well with both the pure Ti and TZHCNSS samples, the gap between the bone tissue and metallic samples are no more than 5 lm.

Fig. 5. SEM images of L929 and NIH3T3 cells on different sample surfaces after 4 days of incubation. (a) L929 cell on pure Ti; (b) NIH3T3 cell on pure Ti; (c) L929 cell on TZHCNSS; (d) NIH3T3 cell on TZHCNSS.

Fig. 6. Implantation of TZHCNSS BMG and pure Ti samples (a) BMG sample; (b) representative X-ray images for the implants and (c, d) representative histological images stained by methylene blue after 1 month implantation. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

7

H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20 Table 3 Summary of biomedical Zr-based BMG systems and their mechanical properties. Chemical composition (at.%)

Preparation method

Critical diameter/ thickness (mm)

Compressive fracture strength (rf) (MPa)

Young’s Vickers Reference modulus hardness (Hv) (kg mm 2)

Zr52.5Al10 Ti5Cu17.9 Ni14.6 (BAM-11)

Arc-melting/copper mold casting

7

1700 (tensile yield strength)

90

590

Zr61Cu17.5Ni10Al7.5Si4

Melt spinning

—

1800

—

510 (5 GPa)

[37]

Zr60Cu22.5Pd5Al7.5Nb5

Arc-melting/copper mold casting

—

1720

82

—

[38]

Zr60Ti6Cu19Fe5Al10

Arc-melting/copper mold casting

—

1652

70

—

[35]

Zr60Nb5Cu20Fe5Al10

Arc-melting/copper mold casting

—

1795

72

—

[35]

Zr60Nb5Cu22.5Pd5Al7.5

Arc-melting/copper mold casting

—

1724

70–85

—

[39]

Zr61Ti2Cu25Al12 (ZrxCu100 x)80(Fe40Al60)20 (x = 68–77)

Arc-melting/copper mold casting Arc-melting/copper mold casting

6 — 83 13 when x = 72.5 1560–1690 — (yield strength)

— 521–563

[40] [41]

Zr100 x y(Cu5/6Ag1/6)xAly (x = 38, 40, 42, 44, 46, 48, Arc-melting/copper mold casting 50 and 52 at.% and y = 6, 7, 8 and 9 at.%) alloys

20

1900–1916

—

—

[42]

Zr56Al16Co28

Arc melting/arc-tilt-casting, arc melting/single roller quenching

18

1830 (tensile strength)

83

—

[43,44]

Zr65Pd17.5Fe10Al7.5

Arc-melting/copper mold casting

6

Around 1500

—

422

[45]

Zr65Pd12.5Ag5Fe10Al7.5

Arc-melting/copper mold casting

6

Around 1500

—

411

[45]

xAgx (x = 0, 1, 3, 5, 7)

Arc-melting/copper mold casting

3–10 (10 when x = 3)

1640–1720 (1720 when x = 3)

68–75

—

[46],

Zr60.14Cu22.31Fe4.85Al9.7Ag3

Arc-melting/copper mold casting

10

1720 ± 28

82 ± 1.9

Zr60 + xTi2.5Al10Fe12.5

xCu10Ag5 (at.%, x = 0, 2.5, 5)

Arc-melting/copper mold casting

—

1450–1580

70–78

443–460

[48]

Zr55Cu30Ni5Al10

—

—

1830

—

416 ± 10

[49]

Zr46Cu37.6Ag8.4Al8

Arc-melting/copper mold casting

—

2158

92

554

[31]

Zr51.9Cu23.3Ni10.5Al4.3

Arc-melting/copper mold casting

—

1997

102

550

[31]

Zr51Ti5Ni10Cu25Al9

Arc-melting/copper mold casting

—

1962

100

542

[31]

Zr62.5Al10Fe5Cu22.5

Arc-melting/copper mold casting

6

1700

80

459 (4.5 GPa) [14]

(Zr0.62Cu0.23Fe0.05Al0.10)100

3. Zr-based bulk metallic glasses for biomedical implants and devices 3.1. Mechanical properties of biomedical Zr-based BMGs During biomedical application services under different physiological conditions, biomaterials would experience loadings from surrounding tissues, blood vessel walls and bones, which pose requirements on their mechanical properties. As shown in Table 3, the biomedical Zr-based BMGs have been featured with a high hardness that is about twice to three times of that for conventional crystalline biomedical 316L SS, Ti alloys and Zr alloys, and high yield strength that is considerably higher than that of those above mentioned conventional crystalline metallic biomaterials. Fig. 4(B) shows the compressive stress–strain curves of three Ni-free Zrbased BMGs. It can be seen from Fig. 4(B) that all BMGs exhibit a high yield strength over 1300 MPa, and fracture strength over 1600 MPa [35]. It is interesting to note that the three BMGs also exhibit a considerably large plastic strain. In addition, the elastic strains around 2% with low modulus of 70–80 GPa were obtained. The modulus is lower than that of 316 L stainless steel (about 200 GPa) and Ti– 6Al–4V alloy (110–125 GPa). The high strength of the Zr-based BMG greatly enables the production of biomaterials with thinner struts for cardiovascular stents, which benefits its deliverability and reduces the rate of restenosis. Compared to conventional 316L SS stents, biomedical Zr-based BMG stents would require only one third of the cross-section of the strut and would have more than five times the deflection [9]. And the BMG bone screws could have a thinner shank and deeper threads thus yielding greater holding power to the fracture bones. It is also worth mentioning that high elastic limit

[36]

[47]

of about 2% and low Young’s modulus indicated that the elastic deformation with external loading will be facilitated and would consequently distribute stresses more uniformly than current materials, minimizing stress concentrations, reducing stress shielding effects, and thus achieving faster healing rates. 3.2. Corrosion behavior of biomedical Zr-based BMGs The corrosion behavior of the biomedical Zr-based BMGs has been investigated in various physiological solutions, including PBS [50], Hank’s solution [37], 0.9% NaCl saline solution [51], Ringer’s solution [9], artificial saliva solution [52], and SBF solutions [47]. The results demonstrated that in comparison with conventional crystalline 316L stainless steel, Zr and Zr-based alloys and Ti and Ti-based alloys, the BMGs show evidently a lower passive current density, much higher pitting potential, suggesting that the passive films formed on the Zr-based BMGs are more protective than on the above mentioned control groups, indicating their enhanced corrosion resistance behavior [38,39]. The high corrosion resistance of the biomedical Zr-based BMGs can be attributed to the formation of the passive films, mainly composed of ZrO2, on the surface of the alloy. Adding Nb [53,54] and Ag [55] are benefit for the corrosion resistance, especially enhanced the corrosion resistance against pitting corrosion. 3.3. Biocompatibility of biomedical Zr-based BMGs For a candidate biomaterial to be used in clinical applications, excellent biocompatibility is an essential property in order to avoid any adverse effect in human body. Many studies have investigated

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H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20

Fig. 7. The morphologies of MG63 cells cultured on (a) Zr60.14Cu22.31Fe4.85Al9.7Ag3 BMG stained with Hochest-33258; (b) Zr60.14Cu22.31Fe4.85Al9.7Ag3 BMG stained with FITC–phalloidin; (c) Zr60.14Cu22.31Fe4.85Al9.7Ag3 BMG stained with both Hochest-33258 and FITC–phalloidin; (d) pure Zr stained with both Hochest-33258 and FITC– phalloidin; (e) Ti6Al4V alloy stained with both Hochest-33258 and FITC–phalloidin; and (f) positive control for 3 days.

the biocompatibility of Zr-based BMGs via both in vitro cell response and in vivo animal implants. And results have shown that Zr-based BMGs showed better biocompatibility than the conventional crystalline 316L stainless steel Zr and Zr-based alloys and Ti and Ti-based alloys, for both the in vitro cell culture (MTT/ CCK8 assay and cell morphology observations), including the mouse pre-osteoblast cells MC 3T3 E1 [55,56], murine fibroblast cells (L929 cell and NIH3T3 cell) [35,38,39,52], human umbilical vein endothelial cells [40] and MG63 human osteoblast-like cells [40,57], macrophage cells [9], human aortic endothelial cells (HAECs) and human aortic smooth muscle cells (HASMCs), and in vivo rabbits implantations [35]. Fig. 7 demonstrates the morphologies of MG63 cells cultured on Zr-based BMG, crystalline pure Zr, and crystalline Ti6Al4V alloy for 3 days. It can be seen from Fig. 7 that a large number of labeled cells are attached to and spread out on these surfaces. The free-growing control cells have a thin, thread-like shape (please refer to Fig. 7(f)). The cells cultured with the three different samples (Fig. 7(c–e)) exhibited an obviously larger footprint area in contrast to the free-growing ones: actin filaments are clearly seen aligned in a parallel direction, and focal adhesion plaques (green) are distributed at the cell periphery, identified as the extremities of the actin filaments. No other obvious differences in adhesion morphology were found between the three metals [47]. Fig. 8 shows the animal study results after implantations with ZrCuAlAg BMG and Ti–6Al–4V

implants. From Fig. 8, we can see that during the whole experiment period, no multinucleated giant cell and lymphocytes can be found at the bone area around the implants, meaning that no inflammation or osteonecrosis occurred. New bone was formed around the implants. After implantation for 4 weeks, gaps between bone tissue and metallic sample are very small (6 lm and 8 lm for ZrCuAlAg BMG and Ti–6Al–4V alloy, respectively) (Fig. 8j and m). After implantation for 8 and 12 weeks, no gap can be observed between the bone tissue and metallic implant samples. Both implant samples are well integrated with bone tissue [58]. 4. Fe-based bulk metallic glasses for biomedical implants and devices 4.1. Mechanical properties of biomedical Fe-based BMGs Fe-based BMGs have very low costs compared to Ti-based or Zr-based BMGs, making them quite attractive for any large-scale biomedical applications. Moreover, they also have reasonably good glass forming ability (GFA) and can be easily prepared by traditional Cu-mold casting/cooling methods. Since the first synthesis of Fe-based BMG in Fe-Al-Ga-P-C-B system in 1995 [59], a variety of Fe-based BMG have been developed, including Fe-(Zr,Hf,Nb)-B, Fe-Co-Ln (lanthanide metal, such as Tm, Er)-B, Fe-Ga-(P,C,B,Si) systems [60]. However, most of these developed Fe-based BMGs are

H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20

9

Fig. 8. The animal study results after implantation for 4 weeks (left column), 8 weeks (middle column), and 12 weeks (right column): (a), (b), (c) radiographs of thighbones with ZrCuAlAg BMG and Ti–6Al–4V implants after surgery, H&E stained results after implantation: (d), (e), (f) ZrCuAlAg BMG implant and (g), (h), (i) Ti–6Al–4V implant, and SEM images of implants and bone interface, and the corresponding EDS analysis with the elemental distribution of the white line: (j), (k), (l) ZrCuAlAg BMG implant; and (m), (n), (o) Ti–6Al–4V implant.

soft magnetic materials as functional and structural materials, for example, power inductor and soft magnetic cores and not suitable for biomedical applications considering their magnetic properties. In clinical, MRI diagnosis is inhibited with the presence of magnetic implants in body because they become magnetized in the intense magnetic field of the MRI instrument, which may produce image artifacts and therefore prevent exact diagnosis [61,62]. To decrease the artifacts, medical devices with low magnetic susceptibility are required. Nonmagnetic Fe-based BMGs were firstly developed by Ponnambalam V et al. in 2003 and are called as ‘‘amorphous steel” [63] because their composition are similar to that of stainless steel, i.e. both of the stainless steel and amorphous steel contain Fe, C, Cr, Mo elements. The chemical composition, mechanical properties of

reported Fe-based BMGs for biomedical application purpose are listed in Table 4. Table 4 demonstrated that amorphous steels have glass-forming ability high enough to form single-phase glassy rods with diameters reaching 16 mm. It is not surprising that a variety of Fe-based BMGs were being investigated as future biomaterials considering their excellent corrosion, wear resistance and relatively low material cost. Previous studies have shown that the vickers hardness of Fe-based BMGs in the range of 1200–1800 MPa and tensile fracture strengths of at least 3000 MPa, values that far exceed those reported for state-of-the-art steel alloys, and Febased BMGs have better MRI compatibility compared to 316L SS [64]. Fig. 4(C) shows the uniaxial compressive true stress–strain curves for four Fe–Cr–Mo–P–C–B BMGs. It can be seen from Fig. 4(C) that the current materials exhibit high yield stresses

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Table 4 Summary of biomedical Fe-based BMG systems and their mechanical properties. Chemical composition (at.%)

Preparation method

Critical diameter/ thickness (mm)

Compressive fracture strength (rf) (MPa)

Young’s modulus

Vickers hardness (Hv) Reference (kg mm 2)

Fe41Co7Cr15Mo14C15B6Y2

Arc-melting/copper mold casting

16

3500

—

1253

[66,67]

Fe63Mo14C15B6Er2

Arc-melting/copper mold casting

3

4000

204

1122 (11 GPa)

[68]

Fe55Cr8Mo14C15B6Er2

Arc-melting/copper mold casting

>4

—

209

1122 (11 GPa)

[68]

Fe49Cr15Mo14C(13 + x)B(8-x)Er1 (x = 2, 4, 5 and 6)

Arc-melting/copper mold casting

3–6

4040–4140

210–220

—

[69]

(Fe0.9Co0.1)58.5Cr6Mo14C(15 + x)B(6-x) Er0.5 (x = 3 and 4)

Arc-melting/copper mold casting

2–4

4070–4100

200

—

[69]

Fe48Cr15Mo14C15B6Er2

Arc-melting/copper mold casting

12

4200

213

1122 (11 GPa)

[68]

Fe70B20Si10

Melt spinning

—

2500

—

714 (7 GPa)

[37]

(Fe44Cr10Mo12.5Mn11C15B6Y1.5) 100 xMnx (x = 0, 2, 4, 8)

Arc-melting/copper mold casting

2–5

—

—

—

[64]

2.5–3

3150–3550

176–183

845–974

[65]

Fex (x = 63–71) Cry (y = 0–3) Moz (z = 5– Arc-melting/copper 12) P12C10B2 mold casting

and fracture strengths reaching 2.9 and 3.5 GPa, respectively, and high plastic strains up to 3.6%. The ductility of amorphous steels, or Fe-based BMGs in general, can be significantly enhanced by properly tuning the alloy compositions [65].

biocompatibility and MRI compatibility before its application in the biomaterials markets. 5. Biodegradable bulk metallic glasses designed for temporary implants and devices

4.2. Corrosion behavior of biomedical Fe-based BMGs It is well known that conventional biomedical stainless steels (3l6L SS and 304 SS) are usually prone to localized attack and Ni ion release in long-term applications due to their poor corrosion resistance because of the aggressive biological effects [70]. The developed Fe-based BMGs have higher pitting potential values and lower corrosion current density values both in Hank’s solution and in artificial saliva solution and have quite lower ion releasing than that of 316L SS [67], together with good biocompatibility in vitro [71]. Shahverdi et al. studied the corrosion behavior of Fe55-xCr18Mo7B16C4Nbx (x = 0, 3, and 4 at.%) amorphous ribbon s in Ringer’s solution [72], and the linear polarization and EIS measurements indicated that the Fe51Cr18Mo7B16C4Nb4 possesses larger polarization resistance value than that of 316L SS and Ti6Al-4V. 4.3. Biocompatibility of biomedical Fe-based BMGs As mentioned above, the enhanced mechanical properties and corrosion behavior guarantee the biocompatibility of Fe-based BMGs as future biomaterials. It has been reported that Fe-based BMGs exhibited better biocompatibility for the in vitro murine fibroblast cells (L929 cell and NIH3T3 cell) culture (MTT/CCK8 assay and cell morphology observations) [71]. Cell number grows quickly during the 4-day culture period for all groups, and there is no obvious difference between negative group and experimental groups. Most cells appear elongated in a spindle shape and become longitudinally aligned at high cell densities. High cell viability value of Fe based BMGs approaching to the negative control group could be explained by the protective effect of the compact oxide film, avoiding metal ion releasing in the biofluid [71]. In addition, previous study has demonstrated Fe-based BMGs have better MRI compatibility compared to 316L SS [64], and this would make it promising candidate for biomedical materials and therapeutic devices used under MRI bio-imaging and diagnostics environment. Further in vivo and pre-clinical case studies are needed to verify the

Besides the joint replacements that need permanent prosthesis implantation in the human body, there are many other clinical cases, such as bone fracture, cardiovascular diseases, in which the temporary implant materials are needed. The fixation or mechanical support are temporarily needed during the healing process of the injured or pathological tissue, and after that, the implants accomplish their mission and will no longer function in human body. In this case, biodegradable materials are the optimal choice as these materials do their job while healing and new tissue forming occur and degrade in the human body thereafter. As summarized in Table 5, the biodegradable BMG family includes members such as Ca-based BMGs, Mg-based BMGs, Srbased BMGs and Zn-based BMGs. Their fabrication methods, physical and mechanical properties are summarized in Table 5 as well. In the following paragraphs, we will separately discuss and thoroughly review the currently developed biodegradable BMG systems. 5.1. Biodegradable Mg-based BMGs Crystalline magnesium alloys have attracted considerable attention as potential implant materials in recent years [92,93], series crystalline magnesium alloys, including Mg-Zn [94–96], Mg-Sr [97,98], Mg-Ca [99,100], Mg-RE [101,102] systems have been designed and developed for biomedical applications. However, the development and deployment of biodegradable magnesium alloys faces some practical challenges. First, a much higher inherent strength would be required of such an alloy since an implant’s strength would naturally deteriorate gradually during the corrosion/degradation process. Next, pitting corrosion, resulting in surface defects, would likewise lead to the quick loss of the magnesium alloy’s strength. Third, most current magnesium alloys have fast degradation rates, exceeding rates for bone healing. Finally, the release of hydrogen and localized alkalization caused by the fast corrosion may also do harm to the surrounding tissues.

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H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20 Table 5 Summary of biodegradable BMG systems and their mechanical properties. BMG system

Chemical composition (at.%)

Preparation method

Ca based

Ca65Mg15Zn20

Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting Induction-melting/copper casting

Ca57.5Mg15Zn27.5 Ca55Mg17.5Zn27.5 Ca52.5Mg20Zn27.5 Ca52.5Mg17.5Zn30 Ca52.5Mg22.5Zn25 Ca50Mg20Zn30 Ca65Li9.96Mg8.54Zn16.5 Ca48Zn30Mg14Yb8 Ca20Mg20Zn20Sr20Yb20 Mg based

Zn based

Sr40Mg20Zn15Yb20Cu5 Sr60Mg18Zn22 Sr60Li5Mg15Zn20

a b c

Reference

364a

20a

1.42

[73–75]

mold 4

—

36.5

0.9

[76]

mold 4.5

—

36

0.9

[76]

mold 2.5

—

39

1.4

[76]

mold 0.9

—

44

1.4

[76]

mold 1.0

—

43

0.8

[76]

mold 1.2

—

46

0.7

[76]

mold 5

530

23.4

1.35

[77]

mold 2

600

31.9

—

[89]

mold 4

370

19.4

—

[79]

800 800 —

— — 66b

2.5 2.5 4b

[37] [37] [80]

47.6– 48.2 930 when x = 30 and — 830 when x = 25 817 —

2.16

[81]

—

[82]

2.16

[83]

550

—

—

[84]

787–848

2.45–2.51

[85]

—

48.5– 49.4 —

—

[86]

780

—

2.35

[87]

500 (tensile)

35

—

[88]

640c

36.6c

—

[89]

663

—

—

[78]

408

20.6

—

[90]

—

19.7

—

[91]

—

18.4

—

[91]

Melt Spinning — Melt spinning — Induction-melting/copper mold 2.6 casting Mg80 xCa5Zn15 + x (x = 5–20) Induction-melting/copper mold 1–4 casting Mg96 xZnxCa4 (x = 30, 25) Induction-melting/copper mold 2–5 casting Mg67Zn28Ca5 Melt-extraction 100 lm thin wires Mg69Zn27Ca4 Induction-melting/copper mold 1.5 casting Mg66Zn30Ca4-xSrx (x = 0, 0.5, 1, Induction-melting/copper mold 4–6 and 1.5 at.%) casting Mg66-xZn30Ca4Ndx (x = 0, 0.5, 1, Induction-melting/single roller 25–35 lm 1.5at.%) spinning Mg66Zn30 xCa4Agx (x = 0, 1, and Induction-melting/copper mold 1–4 3 at.%) casting or melting spinning Mg66Zn30Ca2Yb2, Mg66Zn30Yb4, Induction-melting/copper roller 40–100 lm Mg64Zn30Yb6, Mg60Zn30Yb10 spinning Zn38Ca32Mg12Yb18

Young’s Vickers modulus hardness (Hv) (GPa)

mold 6

Mg65Cu25Gd10 Mg67Cu25Y8 Mg60Cu29Y10Si1

Zn40Mg11Ca31Yb18 Sr based

Critical diameter/ Compressive thickness (mm) strength (MPa)

Induction-melting/copper mold 2 casting Induction-melting/copper mold 2 casting Induction-melting/copper mold 3 casting Induction-melting/copper mold 3 casting Induction-melting/copper mold 3 casting

700

Average data at various strain rates. Estimated date from Fig. 10 in [80]. Calculated from Fig. 3 in reference [89].

Recent research into Mg-based BMGs, however, has found that they have higher strength and lower elastic moduli than pure Mg and conventional Mg alloys. Fig. 9(A) demonstrates the compression properties of Mg66Zn30Ca4 and Mg70Zn25Ca5. It can be observed form Fig. 9(A) that the fracture strength rf for Mg70Zn25Ca5 and Mg66Zn30Ca4 samples is (565.8 ± 23.2) and (531.2 ± 22.8) MPa, respectively, which is much higher than that of pure Mg sample (198.1 ± 4.5) MPa [103]. Studies of the corrosion behavior, cellular response and tissue response of Mg-Zn-Ca BMGs by Zberg B et al. [104] and Gu et al. [103] demonstrate that Mg-ZnCa BMGs present more uniform corrosion morphology than conventional crystalline Mg alloys, have much lower corrosion rates,

and show higher cell viability than conventional crystalline pure Mg. Fig. 10 shows the morphology of (a–c) L929 and (d–f) MG63 cells cultured on (a, d) as-rolled pure Mg, (b, e) Mg66Zn30Ca4 and (c, f) Mg70Zn25Ca5 samples for 5 days. From Fig. 10, it can be seen that for the as-rolled pure Mg sample, a few cells are observed on the surface and unhealthy cells with round shape can be seen after 5 days culture for the two experimental cell lines. For Mg66Zn30Ca4 and Mg70Zn25Ca5 BMGs, the number of cells adhered on the surface are higher than crystalline pure Mg, and healthy cells with elongated or spindle morphologies are observed. In addition, Fig. 11 demonstrated the animal studies of Mg-based glass (a, c) in comparison with a crystalline Mg alloy reference

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H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20

Fig. 9. (A) Compression properties of Mg66Zn30Ca4 and Mg70Zn25Ca5. (B) Stress-strain curves of Ca47Mg19Zn7Cu27 BMG. (C) Stress-strain curve of Sr40Mg20Zn15Yb20Cu5 BMG. (D) Stress–strain curves of Zn38Ca32Mg12Yb18 BMG after immersion in Hank’s solution for various times.

sample (b, d). As shown in Fig. 11, results from animal studies indicate that all samples show a typical fibrous capsule foreign-body reaction (indicated by white arrows), while only the crystalline samples (implanted discs indicated by dashed lines) show pronounced hydrogen evolution (area between discs and fibrous capsules indicated by black arrows). And no tissue-imprinted hydrogen gas cavities had formed in the histological preparations of the Mg-Zn-Ca BMG samples and no inflammatory reactions were observed [104]. J.D. Cao [105] and co-workers characterized the Mg65Zn30Ca5 BMG as a potential biodegradable metal. Their studies demonstrated that confirmed via in vitro corrosion experiment and subsequent chemical analysis of the exchanged media solution, the amorphous alloy exhibited a slower corrosion rate and released Mg ions into the media at approximately half the rate of a crystalline Mg reference sample. Although both types of materials had an influence on the L929 cell viabilities over 72 h culture period, the viabilities remained relatively high, thereby indicating that they are capable of supporting cellular activities. However, direct contact with the samples created regions of minimal cell growth around both amorphous and crystalline samples, and no cell attachment was observed. Yu H-J et al. [88] achieved significantly improved ductility for Mg-based BMGs under bending and tensile loading through minor alloying with the rare-earth element ytterbium (Yb) recording a plastic strain of about 0.5% after 2% and 4% Yb was added into MgZnCa glasses. As an additional benefit, the in vitro biocompatibility of Mg-based BMGs was also improved by the Yb-alloying, a finding confirmed by indirect cytotoxicity and direct cell adhesion, extension, and proliferation assays [88]. In order to make it easier to make medical devices, such as woven stents, from Mg-based BMGs, Zberg B et al. [83] produced wires with very good surface quality via a melt-extraction setup which show extended homogeneous plastic deformation in tensile tests.

5.2. Biodegradable Ca-based and Sr-based BMGs It is well known that pure metals calcium and strontium are so active that they can react violently with water generating hydrogen gas and hydroxides. As a result, it is very hard to gain stable crystalline calcium and strontium alloys. However, amorphous calcium and strontium BMGs have nobler properties due to their unique amorphous structure. According to this, it is easy to understand that for biomedical applications, the only way to develop calcium and strontium based alloy is to form amorphous alloys. 5.2.1. Biodegradable Ca-based BMGs Li et al. [77] reported development of a class of CaLi-based BMGs, a new kind of glassy metallic plastic, which includes multiple valuable properties, such as ultralow density (<2 g/cm3). Similar to conventional glassy polymers, CaLi-based BMGs show polymer-like thermoplastic formability along with excellent plastic-like deformability and can be elongated, compressed, bent, and imprinted near room temperature [106]. Jiao W et al. [89] fabricated a Ca-based BMG with composition of Ca48Zn30Mg14Yb8 with a low Young’s modulus of 31.9 GPa and high fracture strength of 600 MPa. Fig. 9(B) shows the stress-strain curves of Ca47Mg19Zn7Cu27 BMG. From Fig. 9(B), we can see that at room temperature, an as-cast amorphous sample was loaded to 2% elastic strain and then exploded into powder. During compression at 110 °C, 120 °C, and 130 °C, yielding occurred and the flow stress decreased during plastic deformation [107]. Series of Ca-Mg-Zn BMGs were developed by low-pressure die casting [76]. It was observed that both the critical casting size and the rate of corrosion of the alloys were dependent on composition. The corrosion morphology on the Ca-BMG was not typical of that observed in crystalline alloys where localized pitting and microstructural-related (i.e., intergranular or intermetallic driven)

H.F. Li, Y.F. Zheng / Acta Biomaterialia 36 (2016) 1–20

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Fig. 10. The morphology of (a–c) L929 and (d–f) MG63 cells cultured on (a, d) as-rolled pure Mg, (b, e) Mg66Zn30Ca4 and (c, f) Mg70Zn25Ca5 samples for 5 days.

attack is expected to be dominant. Instead, the corrosion morphology consistently revealed a damage form that can be described as incongruent dissolution, similar in form to dealloying. Wang et al. [75] evaluated the feasibility of Ca65Mg15Zn20 BMG for potential skeletal applications by animal tests and found that although this Ca-based BMG degrades fast in vitro, there was no obvious inflammation reaction at the implantation site [108]. Although these Cabased BMGs show great potential as novel biodegradable metals, the rapid rates of dissolution in biocorrosion environments due to the inherently reactive nature of Ca (even more reactive than Mg) may limit their specific applications as biomaterials. In an effort to improve the corrosion resistance of metal materials, two different strategies are usually applied, i.e. surface treatment and alloying method (adding alloying elements into the master metals or alloys). Different kinds of biodegradable thin films, including fluoroalkylsilane (FAS) coating, pure Fe film and (Fe + FAS) bilayer, were introduced on the surface of Ca-Mg-Zn based BMG and has reported to reduce its degradation rate significantly [109]. For the alloying method, considering that BMGs and high entropy alloys (HEAs) both have unique properties that conventional metals and alloys are unlikely to match and the fact that an amorphous phase can be easily formed in HEAs, the research and development of new types of alloys that combine BMG and HEA concepts together would be of great importance for future novel material

studies. Ca20Mg20Zn20Sr20Yb20 high entropy BMG alloy was developed and results showed that both the mechanical properties and corrosion behavior of the Ca-Mg-Zn amorphous alloy were obviously enhanced after adding the alloying elements Sr and Yb and further formed high entropy amorphous structures. In addition, in vitro and in vivo studies showed that the Ca20Mg20Zn20Sr20Yb20 high entropy BMG could stimulate and promote osteogenesis and new bone formation [79]. Fig. 12 shows MG 63 cell staining at 3 days in the negative control group and in cells cultured in series of HE-BMG extracts. It can be observed that healthy morphologies of MG63 cells, polygon or spindle, were well spread and extended completely in both series of CMZSY HE-BMG extracts experimental groups and negative control groups. Fig. 13 demonstrates the histology of cross-sections of the distal femora and diaphyseal region of the mice. From Fig. 13, it can be seen that no inflammation or osteonecrosis was found at the bone area around the implants. The histological observations at 4 weeks after implantation show that the bone thickness around the implanted CMZSY HE-BMG is greater than that of normal bone (yellow arrows in Fig. 13), which is consistent with the radiographs and micro-CT results. The new bone at the inner edge of the cortical bone can be clearly observed (black arrows in Fig. 13). There is no new bone formation around the wound area where no implants were embedded in the control groups [79].

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Fig. 11. Animal studies of Mg-based glass in comparison with a crystalline Mg alloy reference sample. Glassy Mg60Zn35Ca5 (a, c) and crystalline Mg alloy reference (WZ21) (b, d) in two types of porcine abdominal tissue (muscle after 27 days (a, b) and subcutis after 91 days (c, d) of implantation).

Fig. 12. Cell staining at 3 days in the negative control group and in cells cultured in series of HE-BMG extracts. (a) Negative control group; (b) 5% HE-BMG extract; (c) 10% HEBMG extract; (d) 25% HE-BMG extract; (e) 50% HE-BMG extract; (f) 100% HE-BMG extract. 200; the scale bar represents 100 lm.

5.2.2. Biodegradable Sr-based BMGs Strontium-based BMG systems have attracted great interest for their potential in biomedical applications. Studies have shown that strontium can inhibit bone resorption and stimulate bone formation [110]. Sr-based drug treatments for osteoporosis, such as Protelos (Strontium ranelate) [111] can reduce the risk of fracture in

patients after one year of treatment and inhibit osteoclast activity and stimulate osteoblast proliferation [112]. Strontium, calcium and magnesium are in the same group of periodic table and the physiological distribution of Sr is similar to that for Ca, with 99% of the element being stored in bone [113]. Under normal conditions, the bone strontium/calcium ratio varies between 1:1000

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Fig. 13. Histological observation of cortical bone by light microscopy with H&E staining: (a, c) Normal bone and the bone marrow cavity; (b, d) CMZSY HE-BMG rod in the femur of a mouse 4 weeks after implantation.

and 1:2000 with the highest concentration appearing in newly formed bone [114] suggesting its potential value for biomaterials. However, as we discussed before, due to its excessively active property, the design and development of conventional crystalline pure strontium and strontium based alloys for biomedical application is infeasible. In order to overcome this situation, researchers have investigated Sr-based BMGs for their potential use in bone implants and thus in 2009 Zhao K et al. [91] reported a family of Sr-based BMGs with good glass-forming ability and many unique properties. These Sr-based BMGs had ultralow glass transition temperature Tg, low elastic moduli, wide supercooled liquid regions, and showed excellent plastic-like deformability at low temperature, similar to that for CaLi-based BMGs. Moreover, these Sr-based BMGs demonstrated tunable degradation behavior and a corrosion rate that could be controlled simply through minor alloying. Li et al. [90] investigated the bio-corrosion and biocompatibility of Sr40Mg20Zn15Yb20Cu5 BMG. And Fig. 9(C) shows the stress-strain curve of Sr40Mg20Zn15Yb20Cu5 BMG. From Fig. 9(C), it can be observed that the fracture strength rf for Sr40Mg20Zn15Yb20Cu5 BMG is (408.2 6 20.0) MPa [90]. The in vitro cell culture studies demonstrated that most of the MG63 cells that were cultured in the Sr40Mg20Zn15Yb20Cu5 BMG extraction medium were flattened, had polygonal configuration and dorsal ruffles, and were well attached to the substrate by cellular extension, indicating their healthy status [90]. The combination of increased mechanical strength, greater corrosion resistance, and excellent biocompatibility make it as potential material for biodegradable orthopedic implant applications. 5.3. Biodegradable Zn-based BMGs Zinc is an essential element for human beings, serving as a cofactor in all six classes of enzymes [115] as well as several classes of regulatory proteins [116], and is the second most abundant transition metal element in human body [117]. Zinc has also

been widely used as an alloying element in Mg-, Ca-, and Sr- based BMGs, which have great potential for use as biodegradable implants. Recently, crystalline pure Zn and Zn-based alloys, such as Zn-X (Mg, Ca, Sr) alloys have been investigated as potential biodegradable materials and results demonstrated that pure Zn and Zn-based alloys have great potential orthopedic and cardiovascular applications [118–121]. Considering that, biodegradable Zn-based BMGs were developed by Jiao W and co-workers, who further studied their magnetic susceptibility, mechanical properties, corrosion behavior and cytocompatibility [78,89]. Aiming to develop a class of Znbased BMGs as new category of bone repair and fixation materials, Jiao et al. found that a Zn-based BMG with a composition of Zn38Ca32Mg12Yb18 showed much higher strength (above 600 MPa) than conventional 837 Mg (about 200 MPa) crystalline materials, with much smaller magnetic susceptibility (22.3  10 6) than that of commonly used biomedical alloys, and slower degradation rates than pure Mg. In addition, practically no hydrogen was generated during the material’s immersion time. Fig. 9(D) shows the stress–strain curves of Zn38Ca32Mg12Yb18 BMG after immersion in Hank’s solution for various times. As shown in Fig. 9(D), the Zn38Ca32Mg12Yb18 BMG’s compression fracture strength showed no obvious decline after 30 days of immersion in Hank’s solution [89]. This feature of the Zn38Ca32Mg12Yb18 BMG is clearly superior to Mg–6Zn alloy which exhibits a large decrease in bending strength resulting from surface defects such as holes formed during non-uniform corrosion. Cytotoxicity tests (MTT/CCK8 assay and cell morphology observations) revealed that this Zn-based BMG showed good cytocompatibility with MG63 osteoblast cells. Compared to the cell group, the viability of all cultured in extraction is better than 90%, similar to that of the negative group. There was no significant difference (p > 0.05) between MG63 cells cultured in extracts and negative group. MG63 cells presented healthy elongated spindle shape when cultured on the Zn38Ca32Mg12Yb18 BMG substrate. Whereas MG63

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cells after incubation on pure Mg substrate showed round shape in unhealthy or apoptosis state. This further confirms that the Znbased BMG has good cytocompatibility with no negative effects on cell morphology and viability. 6. Future research directions and challenges for BMGs as potential biomaterials Although BMGs have great potential as future biomaterials, there are considerable challenges associated with the development of biomedical BMGs. Firstly, the limited critical sizes of BMGs would hinder the design of biomedical devices using BMGs. Secondly, the toxic elements (such as Ni, Be) that commonly exit in many BMG systems may bring the long term biosafety concerns. Besides, structural relaxation and the associated embrittlement in some BMG systems may occur when processing the BMGs via annealing treatments. Fig. 14 illustrated the future research directions and challenges for BMGs as potential biomaterials. In order to further meet the various requirements for clinical applications, the future studies for BMGs as potential biomaterials are suggested to be focused on the following aspects: overcoming the brittleness, increasing the GFA thus obtaining BMGs with larger sizes, removing/reducing toxic elements, and surface modifications, as schematically illustrated in Fig. 14. 6.1. BMG composites Although BMGs exhibit high strength and show substantial fracture toughness, some of them lack ductility and fail in an apparently brittle manner in unconstrained loading geometries (such as in tensile manner) [122]. The potential for catastrophic failure associated with the rapid propagation of shear bands is a concern for the utilization of BMGs in biomedical applications. As mentioned above, the strain localization and propagation is particularly problematic under tensile stress states where failure may occur along a single shear plane with very limited measurable plastic deformation [123]. For instance, some BMGs exhibit significant plastic deformation in compression or bending tests, but with negligible plasticity (<0.5% strain) in uniaxial tension. To overcome brittle failure in tension, BMG–matrix composites have been introduced [124,125]. In BMG

composites, the second-phase particles can act both as initiation sites for shear bands and as barrier to shear band propagation which result in a dramatic increase in a number of shear bands, and in turn allows for significant ductility [126,127]. Various approaches have been taken to fabricate BMG composites and they can be divided into two groups. One is precipitation of a quasicrystalline phase from the amorphous matrix during annealing [128,129], the other is incorporating a ductile phase in the amorphous matrix [130–132]. For example, The Zr38Ti17Cu10.5Co12Be22.5 BMG/porous tungsten phase composite prepared by pressure infiltration shows greater plastic deformation (30% strain) and flow stress than the unreinforced metallic glass and the pure W [133]. Future studies need to focus on their performances under physiological circumstances, including bio-mechanical properties, bio-corrosion behavior and biocompatibility evaluations. 6.2. BMG foams As mentioned above, the lack of macroscopic plasticity is considered the Achilles’ heel of BMGs and has prevented widespread proliferation of BMGs as biomedical materials. Besides BMG composite, fabricating BMG foams is another solution to overcome this critical defect of BMGs. It has been reported that when BMGs are used in geometries where one dimension is below about 10 times its critical crack length (1 mm for a medium range Zr-based BMG), they exhibit significant bending plasticity. This feature and other size effects have been widely explored in foams to design overall plasticity [134,135]. On the other hand, previous studies have demonstrated that porous structures with suitable pore sizes (200–500 lm) are favorable for cell attachment, proliferation and differentiation [136,137]. Thirdly, the strength and the Young’s modulus of the porous materials can be adjusted through the adjustment of the pore size and the porosity in order to match the requirement of bone repairs as orthopedic implants. Thus, the BMG foams may offer unique properties in biomedical implants. Porous bulk metallic glass Zr57Nb5Cu15.4Ni12.6Al10 (Vit106) has been fabricated by low-pressure infiltration of carbon microspheres and by casting the alloy into a bed of leachable salt (BaF2 has been chosen for its high stability and melting point) that is subsequently dissolved after solidification [138,139]. And the

Fig. 14. Illustration of future research directions and challenges for BMGs as potential biomaterials.

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obtained foam structure is macroscopically homogeneous with a uniform pore size of 212–250 lm and the strength is 100 MPa at 50% strain. In addition, the BMG foams with honeycomb structures have been fabricated by combination of lithography and thermoplastic forming (TPF) fabrication, and the deformation made of a BMG honeycomb can be altered from brittle to ductile through varying the ratio of ligament length to ligament thickness and a 0.2% increase in density doubles strength and energy absorption [140]. Considering the above mentioned information, it seems that the physical and mechanical properties of the porous BMGs are suitable for the biomedical application requisition. Unfortunately, however, to the present authors’ knowledge, the physiological properties of the porous BMGs have not been reported till now. Future studies need to focus on their performances under physiological circumstances, including bio-mechanical properties, biocorrosion behavior and biocompatibility evaluations. 6.3. Removing toxic and noble alloying elements Many BMGs developed so far contain toxic elements for clinical applications, such as beryllium [141], nickel [142] etc. The toxicity of these ions and their release may require further attention, particularly in the light of recent problems with metal toxic effects seen with metal on metal articulations. On the other hand, some BMGs with high GFA usually contain noble metal palladium [143], which increases the production cost. These two factors may severely hinder the widespread applications of BMGs in biomedical applications, thus removing the toxic and noble alloying elements in BMGs while still keep their high GFA at the same time is another research direction and challenge for the further research and development of biomedical BMGs. 6.4. Improving BMG critical sizes Although it is now possible to fabricate BMGs with diameters of several centimeters at certain chemical compositions, most of the Ni-free and Be-free BMGs developed have relatively low GFA so far, resulting in forming small size samples [38,144], and some of the amorphous alloys could only form amorphous ribbons with the thickness of several or ten microns [33,145]. The low GFA and small sizes of amorphous alloys could not satisfy the requirements of future biomedical applications, as most of biomedical devices are in the geometric sizes of several micrometers to several centimeters. Therefore, in order to meet the requirements of biomedical applications, one of the main directions and challenges of further exploring and developing biomedical BMGs is to enhance their GFA and critical sizes significantly. 6.5. Surface modification of BMGs Multiple approaches have been developed to obtain the desired surface properties of biomaterials, and their mechanical, chemical and biological properties can be improved selectively using the appropriate surface treatment techniques while the desirable bulk attributes of the biomaterials are retained [146]. The proper surface treatment will undoubtedly expand the use of biomaterials in clinical. Considering on this point, surface modification of BMGs for biomedical applications is one of the important directions for the further studies of biomedical BMGs. The biggest challenge for surface modification of biomedical BMGs is to choose the proper surface modification method while the amorphous nature of the substrate is not destroyed during the modification process. Thus the desirable properties attributed to the unique amorphous structure of BMGs could be retained during the surface modification process.

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6.5.1. Surface modification of non-biodegradable BMGs Proper surface treatment can further expand the usage of BMG alloys in the biomedical fields. For example, many currently developed non-biodegradable BMGs contain nickel, beryllium, which would have adverse effects on human health when the content is over a certain concentration threshold. Besides developing new BMG systems without these toxic elements, surface modification can also be considered in order to prevent or reduce the release of these elements into the human body. Although there are few literatures about reducing the toxic elements releasing of BMGs by surface modification, there are many literatures about reducing the toxic elements releasing of conventional crystalline biometallic materials, which we can use as references when further studies for surface modifications of BMGs. For instance, the Tantalum [147], zirconium nitride [148], titanium nitride [149,150], TiC/Ti [151,152], diamond-like carbon (DLC) [153,154] coatings that are applied for the preventing of toxic ions releasing and improvement of biocompatibility of conventional NiTi alloys can be considered when further studying and development surface modifications of biomedical BMGs. On the other hand, just as conventional bioinert materials, such as pure Ti and Ti-6Al-4V, non-degradable BMGs are bioinert by nature and cannot form a bioactive bond with the living bone after they are implanted in bony sites. Typically, surface modifications are required in order to prepare a bioactive coating on top of the BMG surface. For the surface modification of non-degradable BMGs, we can consider the methods applied in conventional pure Ti and Ti alloys, such as sand-blasting, electrochemical treatment (anodic oxidation), sol–gel, alkali treatment etc. [155]. A few surface treatments of BMGs have been reported in order to improve their physiological properties and biocompatibility. For instance, Liu’s research group [156] modified the surface of Zr-based BMGs by using the microarc oxidation technique. A rough and porous oxide layer mainly containing tetragonal ZrO2 and amorphous SiO2 incorporating some Ca and P was formed. The Ti- and Zr-based BMGs’ surfaces have been modified by sandblasting and improved the cell attachment, proliferation, and differentiation, thus effective for improving implant osseointegration in vivo [157]. Hydrothermalelectrochemical treatment was employed on Zr-based BMGs [158], and Ti-based BMGs [159], which resulted in the formation of a bioactive surface and can accelerate nucleation and growth rate of apatite on the BMG surfaces. 6.5.2. Surface modification of biodegradable BMGs For some of the developed biodegradable BMGs, such as Ca-MgZn BMG systems, their fast corrosion rate in physiological environments is the limit for their further clinical application. Surface modification is an effective method for slowing down the corrosion rate of biodegradable alloys. It has been reported that fluoroalkylsilane (FAS) coating, pure Fe film and (Fe + FAS) bilayer were employed on the surface of Ca60Mg15Zn25 BMG, and the results revealed that these coatings are quite effective to control the corrosion behavior of Ca60Mg15Zn25 BMG [109]. Other surface modification methods, such as alkaline heat treatment [160], MgF2 coating [161], microarc oxidation [162,163], which are already used for biomedical Mg and Mg alloys are recommended for further studies in biodegradable BMGs’ surface modification. 6.6. Metallic glass coatings Considering their attractive mechanical, physical and chemical properties of metallic glass alloys, there has been substantial and long-standing interest in producing amorphous alloy layers on conventional crystalline metallic substrates since the 1970s [164]. The metallic glass coatings can enhance the properties of

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the substrates, such as the mechanical behavior, wear resistance and corrosion resistance [165]. However, unfortunately, most of the reported amorphous alloy coatings are mainly for structural and engineering applications, and studies on amorphous coating for biomedical application have not yet been explored. Therefore, one of the directions of further studies is to utilize the amorphous alloy coatings onto the existing biomedical implants and devices in order to enhance the properties of the substrates, such as the mechanical behavior, wear resistance, corrosion resistance and biocompatibility.

7. Concluding remarks BMGs have great potential in biomedical applications ranging from orthopedic, cardiovascular to dental implants and fillers. To be specific, as non-degradable BMGs, Ti-based, Zr-based and Febased BMGs with combined excellent mechanical properties and corrosion resistance can be used as biomedical devices, such as surgical blades, pacemakers, medical stapling anvils, and minimally invasive surgical devices; and biomedical implants, such as articulating surfaces, artificial prostheses and dental implants, which are needed to serve a long time in the severe human body environments. On the other hand, biodegradable BMGs, including Mg-based, Ca-based, Zn-based and Sr-based BMGs have great potential as fracture repair materials (such as intramedullary needles, bone plates and bone screws) as well as cardiovascular stent materials, absorbable sutures, fillers around dental implants, and fillings of bone after cyst/tumor removal in arthroplasty; as 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. Acknowledgements This work was supported by the National Basic Research Program of China (973 Program) (Grant No. 2012CB619102), National Science Fund for Distinguished Young Scholars (Grant No. 51225101), National Natural Science Foundation of China (Grant No. 51431002 and 31170909), the NSFC/RGC Joint Research Scheme (Grant No. 51361165101), Beijing Municipal Science and Technology Project (Z141100002814008). References [1] JS Temenoff, AG Mikos, in: Biomaterials: The Intersection of Biology and Materials Science, Pearson/Prentice Hall, 2008. [2] H.F. Li, Y.F. Zheng, L. Qin, Progress of biodegradable metals, Prog. Mater. Sci. 24 (2014) 414–422. [3] Y.F. Zheng, X.N. Gu, F. Witte, Biodegradable metals, Mater. Sci. Eng. R 77 (2014) 1–34. [4] M. Geetha, A. Singh, R. Asokamani, A. Gogia, Ti based biomaterials, the ultimate choice for orthopaedic implants – a review, Prog. Mater. Sci. 54 (2009) 397–425. [5] R. Huiskes, H. Weinans, B. Van Rietbergen, The relationship between stress shielding and bone resorption around total hip stems and the effects of flexible materials, Clin. Orthop. 274 (1992) 124–134. [6] T. Kokubo, Recent progress in glass-based materials for biomedical applications, Nippon Seramikkusu Kyokai Gakujutsu Ronbunshi-J. Ceram. Soc. Jpn. 99 (1991) 965–973. [7] L.L. Hench, The story of Bioglass (R), J. Mater. Sci. 17 (2006) 967–978. [8] W. Klement, R. Willens, P. Duwez, Non-crystalline structure in solidified gold–silicon alloys, Nature 187 (1960) 869–870. [9] J. Horton, D. Parsell, Biomedical potential of a zirconium-based bulk metallic glass, MRS Proceedings, Cambridge Univ Press, 2002. [10] P.H. Tsai, Y.Z. Lin, J.B. Li, S.R. Jian, J.S.C. Jang, C. Li, et al., Sharpness improvement of surgical blade by means of ZrCuAlAgSi metallic glass and metallic glass thin film coating, Intermetallics 31 (2012) 127–131. [11] http://wwwmddionlinecom/article/could-liquidmetal-be-newest-miraclematerial-medical.

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