Metallurgical and Materials Transactions A

, Volume 41, Issue 7, pp 1726–1734

Biocompatibility Study of Zirconium-Based Bulk Metallic Glasses for Orthopedic Applications


    • Department of Materials Science and EngineeringUniversity of Tennessee
    • Department of Mechanical, Aerospace, and Biomedical EngineeringUniversity of Tennessee
  • Andrew Chuang
    • Department of Materials Science and EngineeringUniversity of Tennessee
  • Zheng Cao
    • Department of Materials Science and EngineeringUniversity of Tennessee
  • Peter K. Liaw
    • Department of Materials Science and EngineeringUniversity of Tennessee
Symposium: Bulk Metallic Glasses VI

DOI: 10.1007/s11661-009-0150-5

Cite this article as:
He, W., Chuang, A., Cao, Z. et al. Metall and Mat Trans A (2010) 41: 1726. doi:10.1007/s11661-009-0150-5


Bulk metallic glasses (BMGs) represent an emerging class of materials that offer an attractive combination of properties, such as high strength, low modulus, good fatigue limit, and near-net-shape formability. The BMGs have been explored in mechanical, chemical, and magnetic applications. However, little research has been attracted in the biomedical field. In this work, we study the potential of BMGs for the orthopedic repair and replacement. We report the biocompatibility study of zirconium (Zr)–based solid BMGs using mouse osteoblast cells. Cell attachment, proliferation, and differentiation are compared to Ti-6Al-4V, a well-studied alloy biomaterial. Our in-vitro study has demonstrated that cells cultured on the Zr-based BMG substrate showed higher attachment, alkaline phosphatase activity, and bone matrix deposition compared to those grown on the control Ti alloy substrate. Cytotoxicity staining also revealed the remarkable viability of cells growing on the BMG substrates.

1 Introduction

Metallic materials play an important role in the biomedical field. The significant impact is evident from the extensive list of clinical uses of metallic materials, examples of which include artificial joints, dental implants, maxillofacial devices, stents, artificial hearts, neural prostheses, and more. Particularly for orthopedic applications, metallic materials have been favored due to their excellent mechanical properties, such as strength and fracture toughness. The most commonly used metallic biomaterials for orthopedics include stainless steels, titanium and its alloys, and cobalt-chromium alloys.[1] Despite great progress over the years, these materials suffer from several drawbacks, a major one being causing stress shielding. Stress shielding is a phenomenon related to the mismatch of Young’s moduli of the metallic biomaterial and the surrounding bone.[2] As shown in Table I, the Young’s modulus of a natural bone is typically 3 to 50 GPa, at least half or less compared to the currently used bulk metallic biomaterials.[3] As a result, the majority of the mechanical load is assumed by the high strength metallic implant in the repair region, shielding the surrounding bone from experiencing the stress. This trend eventually causes the bone resorption and complications at the implant/tissue interface.[4] Therefore, it is critical to maintain the Young’s modulus of the implant material close to that of bone for the desirable long-term clinical performance.
Table I

Comparison of Properties of BMG with Those of Bone and Titanium Alloy


Cortical Bone



Yield strength (MPa)

130 to 150

760 to 1050


Elastic strain limit (pct)




Young’s modulus (GPa)

3 to 50

101 to 125


Density (g/cm3)

0.7 to 1.85



Recent advancement in amorphous alloys offers a potential alternate to overcome the stress shielding effect. The amorphous alloys, commonly known as bulk metallic glasses (BMGs), have garnered intense research interest in the materials community due to their remarkable properties. These properties include extremely high elastic strains (~2 pct) in comparison with the common crystalline metallic alloys (<1 pct),[5] high fracture strength,[5] and high fracture toughness.[6] To date, the formation of metallic glasses has been reported for a variety of alloy systems, e.g., Pd-, Pt-, Au-, Mg-, Ca-, Zr-, Ti-, La-, Cu-, Fe-, Co-, and Ni-based systems. The BMGs have been studied and used for magnetic sensing, chemical, and structural applications.[7] As shown in Table I, BMGs offer a unique collection of mechanical properties compared to conventional metallic materials, i.e., lower modulus closer to that of bone and high elastic strain as well as superior strength that would be advantageous for orthopedic applications, particularly for load-bearing applications. For any potential new biomaterials, it is important for the material to have good cytocompatibility, and in the case of orthopedic applications, the material should support the bone forming osteoblast interaction leading to osseointegration. To our knowledge, very little has been reported on the interaction between osteoblast cells and BMGs. Research has been reported on the general cytotoxicity of Zr-based BMGs using fibroblast cells.[8,9] The objective of this study was to investigate the application specific osteoblast cell growth and activity on a zirconium (Zr)–based BMG substrate, to compare the results to the most commonly used titanium alloy of Ti-6Al-4V, and to determine whether this BMG material deserves further detailed biological investigations and mechanistic studies. Specifically, the osteoblast cell attachment, proliferation, differentiation, and morphology were investigated. The model BMG studied in this work is Zr62Cu15.5Ni12.5Al10. Although it contains Ni, previous study has shown good cytotoxicity,[8] and therefore, it is used as the starting material for osteoblast cell compatibility evaluation as the fabrication process is well developed and consistent.

2 Experimental Procedures

2.1 Fabrication and Characterization of Materials

The nominal compositions of the BMGs used in this study were Zr62Cu15.5Ni12.5Al10 (in atomic percent). The alloys were prepared by arc melting the mixture of high-purity Zr (99.5 wt pct), Cu (99.999 wt pct), Ni (99.9995 wt pct), and Al (99.9999 wt pct) elements in a purified-argon atmosphere. Pure titanium was melted prior to the preparation of the master ingot to absorb the oxygen atoms in the furnace, which is detrimental to the glass formability. The master ingots were remelted at least 5 times to improve the microscopic homogeneity of the constituent elements. The cylindrical rods with a diameter of 3 and 50 mm in length were prepared by subsequently suction casting the well-mixed liquid into water-cooled copper molds under an argon environment. Only the center to bottom part of the rod was used in this study to assure that all the specimens had similar cooling rates. The amorphous microstructure of the as-cast specimen was characterized by thermal analysis and high-energy X-ray diffraction. The characteristic temperatures of a metallic glass, glass transition temperature (Tg), and crystallization temperature (Tx), were measured by a PerkinElmer (Waltham, MA) diamond differential scanning calorimetry (DSC) instrument with a constant heating rate of 20 K (20 °C) per minute. High-energy X-ray diffraction was conducted with transmission geometry in the 6-ID beam line at the Advanced Photon Source (APS) at the Argonne National Lab (Argonne, IL). The incident X-ray had an energy of 100 keV and a beam size of 0.5 × 0.5 mm. Discs cut from the top, center, and bottom parts of the BMG rod were examined. For comparative purposes, the commercially available Ti-6Al-4V alloy, which is widely used as an orthopedic implant material, is selected as a positive control in this study. Grade 5 Ti-6Al-4V alloy with a diameter of 4.76 mm was purchased from the McMaster-CARR Company (Robbinsville, NJ). The difference in the diameter between the BMG and Ti alloy was taken into consideration as the results from cell culture study were normalized to the area available for cell growth. Discs of the BMG and Ti alloy with thickness of 1 mm were prepared using a precision diamond-saw cutter, and polished by 600-grit tungsten carbide sand papers to assure the comparable surface roughness between substrates. All the discs were cleaned by sonication in ethanol and acetone, respectively, 3 times for 20 minutes.

Surface wettability of the BMG and titanium alloy discs was determined by static contact angle measurement using a Ramé–Hart contact angle goniometer. Contact angles were measured by the sessile method at room temperature using distilled and deionized water (Milli-Q-Plus millipore, Billerica, MA). Contact angle values are the average result of at least five different symmetrical drops.

For cell culture studies, clean discs of the Zr-based BMG and Ti-6Al-4V alloy were sterilized by soaking in 70 pct ethanol for 30 minutes and exposed to ultraviolet light overnight before being used in cell culture experiments.

2.2 Osteoblast Culture

Newborn mouse calvaria-derived MC3TC-E1 subclone 14 preosteoblastic cells (ATCC) were maintained in an alpha minimal essential medium (α-MEM, Invitrogen, Carlsbad, CA) supplemented with 10 pct fetal bovine serum (FBS, Gibco) and 1 pct antibiotics (penicillin-streptomycin, Invitrogen, Carlsbad, CA) at 309.15 K (37 °C) under 5 pct CO2 and 95 pct air in a humidified incubator. Medium was changed every 2 to 3 days, and confluent cells were subcultured through trypsinization using 0.25 pct trypsin/0.038 pct EDTA (Invitrogen). Cells were seeded onto the BMG and Ti alloy discs in 48-well plates at a given density, i.e., 125,000 cells/cm2 for the attachment study, 50,000 cells/cm2 for the proliferation study, and 10,000 cells/cm2 for the differentiation study. After seeding, the cultures were incubated for 4 hours to permit cell attachment. The cell-seeded substrates were then transferred to new 48-well plates containing fresh media and cultured for various time periods for cell attachment, proliferation, and differentiation study, to eliminate the interference of cells initially attached to the bottom of the well plates. On day 4, the culture medium was replaced with a fresh medium. On day 7, the medium was changed to an osteogenic medium to induce differentiation. The osteogenic medium contains α-MEM supplemented with 10 pct FBS, 1 pct pen-strep, 10−8 M dexamethasone (Sigma, St Louis, MO), 50 μg/mL ascorbic acid, and 8 mM β-glycerol phosphate. The medium was changed every 2 days up to 4 weeks post cell seeding.

2.3 Osteoblast Attachment, Proliferation, and Viability

Osteoblast attachment was studied after 24 hours, and proliferation was examined after 7 days of culturing the cells.[10] The number of cells was quantified by detaching the cells from the surfaces through trypsinization using 0.25 pct trypsin/0.038 pct EDTA (Invitrogen) for 5 minutes and counting them using a hemocytometer. The difference in sample size is factored in by the normalization of the results to the surface area. The viability of the cells cultured on the substrate after 7 days was evaluated via staining using a two-color fluorescence live/dead Viability/Cytotoxicity assay kit (Molecular Probes, Inc., Carlsbad, CA). Briefly, samples were incubated concurrently in 2 μM calcein AM and 4 μM EthD-1 solution for 20 minutes at 309.15 K (37 °C). Images were then captured using a Nikon upright fluorescent microscope.

2.4 Osteoblast Differentiation

2.4.1 Total intracellular protein content

The amount of protein produced by cells was measured 3 weeks after supplementing the culture with an osteogenic medium. The intracellular protein was collected by lysing the adhered cells in deionized water using a standard four-cycle freeze-thaw method.[10] The resulting lysate solution was then collected after centrifugation. The total protein content was quantified using a BCA assay kit (Pierce, Rockford, IL). The absorbance of the solution was measured using a spectrophotometer at a wavelength of 570 nm. The amount of intracellular protein content was determined by converting the absorbance to protein content using an albumin standard curve.

2.4.2 Alkaline phosphatase activity

The lysate solution was used to measure the alkaline phosphatase activity (ALP) using a colorimetric assay (Teco, Anaheim, CA). Briefly, ALP acts upon the 2-amino-2-methyl-1-propanol buffered sodium thymolphthalein monophosphate. The enzyme activity can be stopped by the addition of an alkaline reagent, and a blue chromogen is simultaneously developed. The absorbance at 590 nm was measured and converted to concentration using the ALP standard thymolphthalein. Data were normalized with the total protein content to account for variations in the number of cells attached on different surfaces.

2.4.3 Calcification assay

The calcium deposition of osteoblast cells was studied using a colorimetric assay (Teco). Briefly, after all the lysate was removed, the sample surfaces were soaked in 6 N HCl solution overnight to dissolve the deposited calcium. The collected calcium solution was then reacted with the assay reagents where calcium reacts with cresolphthalein complexone in 8-nydroxyquinoline to form a purple color complex, and the solution absorbance at 570 nm was measured with a spectrophotometer. The amount of calcium was determined by converting the absorbance to concentration using the calcium standards, and data were further normalized against the total protein content to account for variations in the number of cells attached on different surfaces.

2.5 Cell Morphology

Cell morphology on BMG and Ti alloy surfaces was examined using scanning electron microscopy (SEM) and immunocytochemical staining. Initial cell spreading at 4 hours was studied by immunofluorescent double labeling of cytoskeleton F-actin filaments and the focal adhesion component vinculin. Briefly, cells were fixed with 4 pct paraformaldehyde (Sigma) in 1X PBS for 30 minutes at 276.15 K (4 °C). After washing with PBS, the samples were permeabilized with 0.1 pct Triton X-100 (Fisher Scientific, Rockford, IL) in 1X PBS for 5 minutes, blocked with 1 pct bovine serum albumin (Sigma) for 30 minutes, and incubated in the primary antibody mouse antivinculin (Chemicon, Billerica, MA) at 1:500 dilution for 1 hour. After rinsing, the samples were incubated with goat antimouse IgG Alexa Fluor 488 (1:200) and Alexa Fluor 594-conjugated phalloidin (Molecular Probes, Inc., 1:50) that labels the cytoskeleton F-actin filaments. Cells were visualized using a Nikon (Nikon Instruments Inc., Melville, NY) upright fluorescent microscope. For SEM study, cells cultured on the samples 3 weeks after supplementing with the osteogenic media were washed twice in PBS and then fixed with 3 pct glutaraldehyde in 0.1 M cacodylate for 1 hour. The surfaces were subjected to three 10-minute buffer washes. A secondary fixation was performed by incubating the samples in 2 pct osmium tetroxide in 0.1 M cacodylate for 1 hour. The cells were then dehydrated by replacing the buffer with increasing concentrations of ethanol (25, 50, 70, 95, and 100 pct) for 10 minutes each. Samples were then critical point dried and sputter coated in gold. The SEM imaging was conducted on the Leo 1550 (Zeiss, Germany) scanning electron microscope at an accelerating voltage ranging from 1 to 20 keV. Cells cultured on specimens for 7 days were also fixed for the SEM study.

2.6 Statistical Analysis

All data were presented as the mean value ± standard deviation. The variation between test groups was studied using the Student’s t-test. A p-value, <0.05, was considered to be statistically significant.

3 Results and Discussion

The composition of the BMG samples investigated in this study was selected based on the improved plastic strain to failure,[11] which is one of the major concerns for load bearing applications. The results of thermal analysis and X-ray diffraction of the as-cast samples were shown in Figure 1. Both measurements confirmed that the specimens were in a fully amorphous state. The DSC curve showed distinct glass transition temperature and crystallization temperature at 652.15 K (380 °C) and 754.15 K (482 °C), respectively, which were comparable to previously reported results.[11] X-ray diffraction of the discs from the top, center, and bottom parts of the BMG rod exhibited identical features. No sharp diffraction spot was found, further suggesting that the prepared BMG sample was in a fully amorphous state. In order to minimize the effect of surface roughness on cell interaction with the substrates, both Zr-based BMG and Ti-6Al-4V samples were polished with a 600-grit tungsten carbide sand paper, which corresponds to an average roughness of 16 μm. As shown in the SEM images (Figure 2), both substrates presented similar morphology of parallel grooves, typical of the polishing process using 600-grit sand paper. The samples were rinsed with water continuously throughout the entire polishing process. Therefore, the temperature is well below its glass transition temperature. Both DSC measurement and XRD showed that wet polishing did not crystallize the surface of the BMG specimen. The water contact angle on the Zr-based BMG was 72.1 ± 2.80 deg and that on the Ti alloy was 65.2 ± 1.98 deg, suggesting that the BMG surface is relatively more hydrophobic than the surface of Ti alloy. This could be attributed to the oxide layer present on the Zr-based BMG surface. Formation of an oxide layer on the Zr-based BMG surface is common and has been reported earlier by Hiromoto and co-workers.[12,13] Hydrophobicity is one of the most important surface properties that directly influence cellular interactions with a biomaterial. Its effect is reflected from the relationship between protein adsorption to a surface and the surface hydrophobicity. It has been well established that cells interact with a biomaterial through such layer of proteins adsorbed on biomaterial surface.[14] Therefore, hydrophobicity of a biomaterial will determine the profile of the adsorbed protein layer and subsequently affect cellular behavior on the surface.
Fig. 1

(a) Thermal analysis of the as-cast Zr52.5Cu15.5Ni12.5Al10 by DSC. (b) High-energy X-ray diffraction of the as-cast Zr52.5Cu15.5Ni12.5Al10
Fig. 2

SEM images showing the morphology of the polished surfaces of (a) Zr52.5Cu15.5Ni12.5Al10 BMG and (b) Ti-6Al-4V alloy. Scale bar = 50 μm

Cytotoxicity of Zr-Cu-Ni-Al-Nb (Vit106a) has been studied with mouse fibroblast, and good cytotoxic properties have been reported.[8] For a specific biomedical implant, it is of equal importance to study how the cells specific to the implantation site will interact with the biomaterial. It has been shown that osseointegration is an important issue for the proper anchorage of orthopedic implants with the surrounding bone tissue.[15] Osteoblastic cells, being one of the main cell types found in the bone tissue, play a key role during the early stages of osseointegration. Therefore, it is imperative to examine the osteoblastic response when studying the potential of any new orthopedic implant material. In the present study, various aspects of the osteoblastic cell response to the Zr-based BMG substrate were assessed using a well-studied MC3T3-E1 cell line in-vitro culture model. For anchorage-dependent cells, such as osteoblast, the initial attachment to the substrate is essential for any further cell functions, including proliferation and differentiation. As shown in Figure 3(a), the Zr-based BMG not only supported the osteoblast attachment, the number of cells attached after 24 hours was significantly higher on the Zr-based BMG substrate than on the Ti alloy. It has been reported previously that Zr implants showed a higher degree of bone-implant contact in vivo compared to Ti implants.[16,17] As Zr makes up the majority element of the tested BMG substrate, it explains the favorable attachment of osteoblasts on the Zr-based BMG surface. It could also be attributed to the difference in surface hydrophobicity between Zr-based BMG and the Ti alloy. As mentioned previously, surface hydrophobicity affects cellular interaction with biomaterials through the layer of adsorbed protein. The slightly more hydrophobic Zr-BMG could have resulted in a higher amount of adsorption of serum proteins from culture medium, which would be favorable for cell attachment. Such speculation is consistent with the study of Williams et al.[18] where surface protein concentrations on a 70 deg contact angle surface vs a 60 deg contact angle surface were 2.7 and 2.5 mg/m2, respectively.
Fig. 3

(a) Osteoblast attachment on the surfaces of Zr-based BMG and Ti alloy after 24 hours of culture; n = 3 (*p < 0.05). (b) Osteoblast proliferation on the surfaces of Zr-based BMG and Ti alloy after 7 days of culture; n = 3. (c) Cell viability evaluated using live/dead assay after 7 days of culture. Both substrates showed healthy cell growth (green) and no dead cells were observed (red). Scale bar = 100 μm

Cell proliferation was investigated to determine whether the Zr-based BMG substrate supported osteoblast growth, which would be essential for subsequent bone matrix deposition. The population of adhered cells is shown to increase when prolonging the culture time to 7 days (Figure 3(b)), suggesting that the Zr-based BMG is permissive for osteoblast proliferation. No statistical difference was observed when compared with the control Ti alloy. Such good cell growth is also reflected from the live/dead viability staining shown in Figure 3(c). Cells are very healthy on the BMG surface after 7 days of culture, as suggested by the intense and uniform green fluorescent signal (indication of live cells), and virtually no red fluorescent signal was observed (indication of dead cells). This result is comparable to cells growing on the control Ti alloy and indicates good viability of growing cells on the Zr-BMG despite the presence of Ni composition.

After 7 days of culture on Zr-based BMG and Ti alloy, the preosteoblastic MC3T3-E1 cells were provided with the osteogenic media to induce cell differentiatition into the bone forming osteoblast cells and bone matrix deposition. The extent of cell differentiation on different substrates was assessed by measuring the ALP activity and the amount of calcium deposited on the surface. In order to account for the variations in the number of cells attached on different surfaces, the results were normalized to the total protein content determined via the colorimetric BCA protein assay. Differentiation is an important aspect of cell function. It refers to cells changing from being less specialized to more specialized. A key characteristic of the differentiated osteoblastic cells is their ability to synthesize the bone tissue, which is rich in calcium and phosphorous. Alkaline phosphatase is an enzyme responsible for removing phosphatase groups in molecules and has a role in the mineralization of bone.[19] The ALP level is generally elevated during the period of active bone growth.[20] Using a colorimetric assay, ALP levels for cells growing on various surfaces were measured after 3 weeks of culture in the osteogenic media. Cells on the Zr-based BMG surfaces show higher ALP levels compared to those on the control Ti alloy surfaces (Figure 4(a)). There is an approximately fourfold increase in ALP levels on the BMG surface after 3 weeks of culture (p < 0.05). Accompanying osteoblast differentiation, bone matrix is deposited on the surface, which predominantly consists of calcium phosphate. Using a strong acid, the deposited calcium was dissolved into a solution, allowing concentration quantification with a colorimetric assay. Analysis of the calcium deposited on the substrate surfaces indicated that there is over a twofold increase in calcium content on the BMG surface compared to the Ti samples (Figure 4(b), p < 0.05). Collectively, these results suggest that the Zr-based BMG provides a favorable substrate for osteoblast differentiation and matrix production, and could potentially contribute to a successful osseointegration in vivo.
Fig. 4

(a) ALP activity measured after 3 weeks of culture of osteoblast cells on the surfaces of Zr-based BMG and Ti alloy after supplementing the osteogenic media. Data was normalized with total protein content to account for the variations in the number of cells attached on each surface; n = 3 (*p < 0.05). (b) Calcium concentration measured after 3 weeks of culture of osteoblast cells on the surfaces of Zr-based BMG and Ti alloy after supplementing the osteogenic media. Data was normalized with total protein content; n = 3 (*p < 0.05)

The morphology of osteoblastic cells growing on the Zr-based BMG surface was investigated using both immunocytochemical staining and SEM. After 4 hours of culture, the attached osteoblast cells were seen spreading on the BMG surface (Figure 5(b)) with highly organized F-actin stress fibers (red). The presence of filopodia, which is important in cell migration, can be identified by the focal adhesion formed between the cell and the substrate through a protein called vinculin (green). This observation suggests that Zr-BMG support initial osteoblast attachment and spreading, which precedes further cellular functions such as proliferation and deposition of bone matrix. Figure 6 shows SEM images of MC3T3-E1 cells on BMG surfaces at various time points of culture. One week after seeding, the osteoblast cells attached and proliferated on both the BMG surface and the control Ti alloy surface (Figures 6(a) and (b)). The cells are well spread on both surfaces, and numerous cytoplasmic extensions are observed. Higher magnification (Figure 6(e)) further revealed the filopodia extended by osteoblasts on Zr-based BMG surfaces. It confirms the immunocytochemical staining results that cells use these thin filopodia to form focal adhesion with the substrate as well as connection with neighboring cells. After 3 weeks of culture in the osteogenic media, the SEM images show that the entire surface is completely covered with a flattened confluent layer of cells along with the secreted mineralized matrix (Figures 6(c) and (d)). Cells on the BMG substrate appear to be more elongated, compared to their morphology on the Ti alloy. Qualitatively, these SEM results demonstrated the healthy osteoblast cell growth on the Zr-based BMG substrate.
Fig. 5

Double immunofluorescent staining of actin filament (red) and vinculin focal adhesion (green) of osteoblasts cultured on Ti alloy (a) and Zr-based BMG (b) for 4 hours. Scale bar = 100 μm
Fig. 6

SEM images showing osteoblast cell morphology after 7 days of culture (a), (b), and (e) and 3 weeks of differentiation (c) and (d) on Ti alloy (a) and (c) and Zr-based BMG (b), (d), and (e). A magnified view of osteoblast cells on Zr-based BMG at day 7 (e) was shown to highlight the observed filopodia extension from the cells. Scale bar = 20 μm (a); 10 μm (b) through (d); 2 μm (e)

One concern with this particular Zr-based BMG material is the presence of elements Ni and Cu. In terms of biocompatibility, Ni and Cu are classified to be a sterile abscess (toxic) group.[21] Nickel is linked to the occurrence of allergy.[22] However, results from the presented in-vitro cell study indicate that, despite the presence of Ni and Cu in the Zr-based BMG, the overall biocompatibility of BMG is comparable to the control Ti alloy, and cells displayed even more active differentiation on the BMG substrate than on the control, which encourages further development of Ni-free Zr-based BMG for orthopedic implants. The observed biocompatibility can possibly be explained by the formation of highly corrosion-resistant and biocompatible oxides such as Zr oxide in a physiological environment.[9] The resulting stable oxide layer acts as a barrier in blocking the diffusion of toxic metal ions and therefore improving the overall cytotoxicity.[23] It could also be due to the good corrosion resistant property exhibited by BMGs. For further development of BMGs into orthopedic implants, elements other than Ni and Cu should be explored to minimize the potential of toxicity and complications. Elements, such as calcium (Ca), niobium (Nb), and tantalum (Ta), are a few examples of excellent biocompatibility.[1] Another issue that should be noted is the density of the BMG material. As seen from Table I, the Zr-based BMG is slightly denser than the currently used Ti alloy (6.6 g/cm3vs 4.4 g/cm3). A possible way to address this issue is to take the inspiration from the structure of a natural bone and introduce porosity into the BMG substrate, i.e., the foamed BMG.[2426] Porosity will also improve osseointegration and implant stabilization with the adjacent tissue. It has been suggested that porosity throughout the entire implant structure could reduce the stiffness of the material, which further addresses the stress shielding issue.[27] Finite element analyses indicated that the amount of stress shared with the surrounding bone was significantly increased for porous titanium than for solid titanium.[27] Wada and Inoue[28] have reported that the Young’s modulus of a Pd-based BMG with 50 pct porosity is approximately 20 GPa, which is comparable to that of the cortical bone.

4 Conclusions

In this study, the bone cell osteoblast response to Zr-based bulk metallic glasses (BMGs), an emerging class of materials that offers an attractive combination of properties, such as high strength and low moduli, has been investigated. The Zr-based BMG supported osteoblast attachment and proliferation. After induction in the osteogenic media, cellular differentiation and bone matrix deposition of osteoblasts proceeded to a higher extent on the Zr-based BMG surface than on the traditional orthopedic implant Ti alloy. The present study demonstrates the good biocompatibility of Zr-based BMGs, and further investigation of BMGs as a potential orthopedic implant material is warranted. It would be of great interest to study the benefit of the low modulus of BMGs in reducing problems with stress shielding in appropriate animal models.


We thank Dr. John Dunlap for his help with the cell fixation for SEM; Mr. Jack McPherson for helpful discussion; and Professor Jun Shen, Harbin Institute of Technology of China, for initial supply of BMG material. One of the authors (PKL) is very grateful for the support of the National Science Foundation Integrative Graduate Education and Research Traineeship (IGERT) Program (Grant No. DGE-9987548) and the International Materials Institute (IMI) Program (Grant No. DMR-0231320).

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© The Minerals, Metals & Materials Society and ASM International 2010