Journal of Non-Crystalline Solids 411 (2015) 45–52

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Journal of Non-Crystalline Solids

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Influence of heat treatment on the mechanical properties, corrosionbehavior, and biocompatibility of Zr56Al16Co28 bulk metallic glass

Rongrong Wang a, Yingying Wang b, Jie Yang a,⁎, Jiaming Sun a, Lingyun Xiong a

a Department of Plastic Surgery, Wuhan Union Hospital, Tongji Medical College of Huazhong University of Science and Technology, Wuhan 430022, Hubei Province, People's Republic of Chinab MATEIS — CNRS UMR 5501, INSA de Lyon, 69621 Villeurbanne, France

⁎ Corresponding author.E-mail address: abrams18@163.com (J. Yang).

http://dx.doi.org/10.1016/j.jnoncrysol.2014.12.0180022-3093/© 2014 Elsevier B.V. All rights reserved.

a b s t r a c t

a r t i c l e i n f o

Article history:Received 22 September 2014Received in revised form 12 December 2014Accepted 13 December 2014Available online 29 December 2014

Keywords:Bulk metallic glass;Heat treatment;Mechanical properties;Corrosion;Biocompatibility

This paper reports the influence of heat treatment on themechanical properties, corrosion behavior and biocom-patibility of Zr56Al16Co28 bulkmetallic glass (BMG) investigated by dynamicmechanical analysis (DMA), electro-chemical analysis in vitro hemocompatibility and in vivo experiments. Zr56Al16Co28 alloyswith amorphous, semi-crystalline and fully crystallized microstructures were obtained by thermal treatment of as-cast Zr56Al16Co28BMG below or above glass transition temperature (Tg). The modulus slightly decreased with increasing temper-ature below Tgwhereas themodulus dramatically increased treated above Tg due to the appearance of crystallinephase. The passivation potential did not change for samples heat-treated below Tg but decreased significantlyabove Tg, which indicated that Zr56Al16Co28 BMG possessed better corrosion resistance than crystallizedZr56Al16Co28 alloys. Zr56Al16Co28 alloys showed good hemolytic resistance regardless of the microstructural con-dition and presented lower hemolytic rate than their crystallized counterparts. In vivo test revealed that as-castZr56Al16Co28 BMGpresented similar biocompatibility to the structurally relaxed BMG and better biocompatibilitythan the crystallized counterparts. Mice implanted with Zr56Al16Co28 BMGs showed lower degree of inflamma-tory infiltration response compared to mice implanted with the crystallized counterparts. The results showedthat the properties of as-cast Zr56Al16Co28 BMGwere generally comparable to its structurally relaxed counterpartand superior to the crystalline samples.

© 2014 Elsevier B.V. All rights reserved.

1. Introduction

Bulk metallic glasses (BMGs) have been developed in a number ofsystems and among these BMGs, the Zr-based BMGs exhibit a uniquecombination of properties due to their high glass forming ability, highstrength, low modulus, good impact toughness, good biocompatibilityand corrosion resistance, which make them attractive candidates forbiomedical application [1–10].

The Zr-based BMGs are considered as promisingmaterials for ortho-pedic applications [11]. For implant materials, a comparablemodulus tobone is critical to avoid stress shielding. The Young's modulus of Zr-based BMGs is 70–80 GPa [11,12], which is closer to that of bone(3–50 GPa [11,12]) than the values of conventional materials (i.e. mod-ulus of Ti–6Al–4V is 101–125 GPa [11,12]). Zr-based BMGs also showhigh corrosion resistance and low cytotoxicity, which are very impor-tant for the biocompatibility of materials [9,13–16]. Zr–Al–Co BMGsare biocompatible due to the absence of copper, which may result inhigh cytotoxicity if released excessively during corrosion [2,12,17]. Coions are also known to be cytotoxic, but the total mass of Co ions re-leased per unit surface area of glassy Zr56Al16Co28 alloy (0.3 ng/mm2)

is dramatically low due to the excellent corrosion resistance ofZr56Al16Co28 BMGs [18]. It has been reported that the Zr–Al–Co BMGsare spontaneously passivated with low passive current densities andwide passive potential regions in physiologically relevant environ-ments, proving that Zr–Al–Co BMGs are highly corrosion resistant [2,18]. Pang et al. [19] studied the corrosion behavior of the Zr55Al20Co25alloy in 3 mass% NaCl solution and pointed out that the surface filmenriched by Zr and Al is responsible for the high corrosion resistanceand spontaneous passivation behavior of the Zr–Al–Co alloys.

Many publications reported the influence of heat treatments on themechanical properties of BMGs [20–22]. Materials heat-treated belowthe glass transition temperature Tg tend to move to a more stable state(at low-energy state) [22–25]. The mechanical behaviors are improvedby physical aging in metallic glasses [24–26]. If the annealing time islong enough or the annealing temperature is high enough, crystalliza-tion of the metallic glasses will occur, which dramatically alters theirmechanical properties. Heat treatments alsomodify the surface compo-nents andmicrostructures inducing changes in corrosion resistance andbiocompatibility [13,15,27,28]. The corrosion resistance of BMGs afterheat treatments may change to a greater or less extent depending onthe nature of the alloys, for instance, Pd40Ni40P20 BMGs present a highercorrosion resistance [29] whereas Zr60Cu20Al10Fe5Ti5 shows a lowercorrosion resistance after high-temperature heat treatment [30].

Fig. 1. XRD patterns for as cast Zr56Al16Co28 BMG and samples after thermal treatment at660 K, 800 K and 973 K.

46 R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

As a potential method to modify the biomaterials, the influence ofthermal treatment on the mechanical properties, corrosion resistanceand biocompatibility requires comprehensive examination. Thus, thecritical issues of this work are (1) to investigate the effect of heat treat-ment on the mechanical properties of Zr56Al16Co28 BMGs; (2) to inves-tigate the corrosion behavior of Zr56Al16Co28 alloys after heat treatmentin a physiologically relevant environment; and (3) to study the in vitroand in vivo biocompatibility of Zr56Al16Co28 alloys. All of the results pro-vide the foundational information of Zr56Al16Co28 BMGs for the biomed-ical applications.

2. Experiment

2.1. Material fabrication

Master alloy with a nominal atomic composition of Zr56Al16Co28(atom ratio) was prepared in melting equipment under a Ti-getteredargon atmosphere. The details of preparation method were enclosedelsewhere [23]. All the samples were re-melted at least four times toensure the chemical homogeneity. Plates with a thickness of 2 mmand 70–80 mm in length were made from a copper-mold castingmethod.

2.2. DSC

The thermal properties of themetallic glasswere determined by dif-ferential scanning calorimetry (DSC), whichwas conducted at a heatingrate of 20 K/min using a standard commercial instrument (Pekin Elmer,DSC-7) under high purity dry nitrogen at a flow rate of 20ml/min. Thenthe characteristic temperatures were obtained: glass transition temper-ature (Tg), temperature corresponding to the onset of the crystallization(Tx) and crystallization temperature Tc.

2.3. Heat treatment

Three as-cast Zr56Al16Co28 bulkmetallic glasses (BMGs)were heatedin an oven protected in a nitrogen atmosphere at varying temperatures(660 K, 800 K and 973 K) and labeled as Zr–660 K, Zr–800 K and Zr–973 K respectively. For each sample, the annealing time lasted 5 h tomake themicrostructure fully relaxed. As-cast andheat-treated sampleswere cut into rectangular shapes (10 × 3 × 1.5 mm3) for hemolytic testand in vivo experiments.

2.4. XRD

The structures of the metallic glass in the as-cast state and heat-treated samples were detected by X-ray diffraction (XRD) at room tem-peraturewith Cu Kα radiation produced using a commercial device (D8,Bruker AXS GmbH, Germany). The XRD is operating at reflection modeat 40 kV and 40 mA with a beam size of 1.5 mm × 10 mm.

2.5. TEM

High-resolution transmission electron microscopy (TEM, JEOL2010 F, equipped with an energy-dispersive X-ray (EDX) detector)was conducted to characterize the atomic structures of Zr56Al16Co28bulk metallic glass. Samples with thickness around 15 μm were carriedout by ion beammilling at 4 keV at an incident angle of 5°. Microstruc-tures at atomic level were characterized by selected area electrondiffraction (SAED) and Fourier transform of high-resolution TEMmicrographs.

2.6. DMA

Mechanical spectroscopy experiments (also termed dynamic me-chanical analysis (DMA)) were performed on the Zr56Al16Co28 BMG

using a forced-vibration equipment. The equipment parameters wereenclosed elsewhere [23]. DMA experiments were performed at a fixedfrequency of 0.3 Hz and heating with a constant heating rate at 5 K/min.Experimental samples with the dimension of 30 mm in length × 2 mmin width × 1 mm in thickness were prepared. All the experiments wereperformed in a high vacuumatmosphere to avoid oxidation. Strain ampli-tude was lower than 10−4.

2.7. Electrochemical analysis

The corrosion behavior of the samples was evaluated by electro-chemical measurements. Electrolyte used was simulated body fluid(Ringer's solution) with pH of 7.5 at 310 K open to air, which was pre-pared from reagent grade chemicals and deionizedwater. The composi-tion of Ringer's solution (g/L) is the following: 6.5 g NaCl, 0.42 g KCl,0.25 g CaCl2 and 8.4 g NaHCO3. Electrochemical measurements wereconducted in a three-electrode cell using a platinum counter electrodeand a saturated calomel reference electrode (SCE). Potentiodynamic po-larization curves were measured with a potential sweep rate at1 × 10−3 V/s in Ringer's solution after immersing the specimens for36,000 s.

2.8. Hemolytic test

20 ml of healthy human blood was taken from a female volunteerand treated with 1 ml of 2% potassium oxalate solution. The anticoagu-lant sample was diluted with normal saline (4:5 ratio by volume).Metallic samples (10×3×1.5mm)were dipped in a standard tube con-taining 10ml of normal saline and incubated at 37 °C for 30min respec-tively. Then 0.2ml of diluted bloodwas added to this standard tube andthe mixtures were incubated for another 60 min at 37 °C. Similarly,normal saline solution was used as a negative control and distilledwater as a positive control. Next, all the tubes were centrifuged for5 min at 750 ×g, and the supernatant was obtained and transferred toa 96-well plate for spectroscopic analysis at 545 nm using a microplatereader (Thermo, Multiskan Spectrum). Each sample was tested repeat-edly three times. The hemolytic rate Z was calculated with the equa-tion [31]:

Z ¼ Dt−Dncð Þ= Dpc−Dncð Þ � 100%

where Z represents hemolytic rate, Dt is the absorbance value for

Fig. 2.High resolution TEM images for Zr56Al16Co28 alloys at different states: (a) as-cast corresponding the selected-area electron diffraction (SAED) pattern (see insets). (b) Sample heat-treated at 660 K. The inset is a SAED pattern. (c) Sample heat-treated at 800 K. (d) Sample heat-treated at 973 K.

47R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

sample, Dnc is the absorbance value for negative control, and Dpc is theabsorbance value for positive control.

2.9. Biocompatibility

According to the randomnumber table, 50 Balb/cmice (6weeks old,animal experiment center of Wuhan University, Wuhan, China) weredivided into the following groups: negative control, as-cast, Zr–660K,Zr–800K and Zr–973K. The metallic sample was implanted subcutane-ously in the dorsal region of each mouse after general anesthesia with2% pentobarbital solution. All the mice were housed in a SPF level labo-ratory. At 1 week, blood samples were collected from the retrobulbar

Fig. 3. Storagemodulus G′ and lossmodulus G″ versus temperature for as cast Zr56Al16Co28 BMG(0.3 Hz). Gu corresponds to the value of the unrelaxed modulus.

venous plexus using EDTA-K2. Hematological analysis was evaluatedusing an automatic blood analysis apparatus. At 4 weeks, the metallicsamples were taken out and the corresponding skin tissue was sec-tioned and stained with hematoxylin and eosin. Similarly,hematological analysis was employed for every mouse. Each metallicsample was washed by an ultrasound apparatus and then the surfacemorphology was observed.

2.10. Surface morphology observation

The metallic plate samples after electrochemical analysis wascleaned by ultrasonic vibration in distilled water for three times,

. Experimentswere carried outwith continuous heating (5 K/min) at a constant frequency

Fig. 4. Polarization curves of as-cast, Zr–660K, Zr–800K and Zr–973K tested in a simulatedbody fluid at 310 K.

48 R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

followed by drying and observing the surface morphology using opticalmicroscopic (KH-7700, Hirox Europe).

3. Results and discussion

3.1. Structure characterize

As-cast Zr56Al16Co28 BMG was analyzed with different scanningcalorimetry (DSC) at a heating rate of 20K/min. Glass transition temper-ature Tg is 741 K, and the temperature corresponding to the onset ofcrystallization Tx is 785 K. Super-cooled liquid region, ΔT, equals to44 Kwhich is similar to the value in a previous study [23]. Two exother-mic peaks are observed: Tc1 at 800 K and Tc2 at 901 K. These tempera-tures are the reference temperatures for heat treatments.

Fig. 1 shows the XRD patterns of as-cast Zr56Al16Co28 BMG and heat-treated samples. A broad halo is observed in the patterns of as-cast andZr–660K indicating the amorphous nature of these materials. Severalnew peaks associating to the crystal scattering appear in the pattern ofZr–800K. The broad halo from amorphous phase also can be clearly

Fig. 5. Surface morphology after corrosion in body simulated flu

observed in the pattern indicating that the amorphous phase and crystalphase both exist after heat treatment at 800 K, which is induced by par-tial crystallization at Tc1. Furthermore, the broad halo disappears in thepattern of Zr–973 K indicating that the sample is fully crystallized abovecrystallization temperature Tc2. After heat treatments, three sampleswith different structures were obtained: amorphous, semi-crystallineand fully crystallized Zr56Al16Co28 alloys.

Fig. 2 shows themicrostructures of as-cast and heat-treated samplesobserved by the high-resolution TEM. As-cast BMG and Zr–660K areconfirmed to be the amorphous phase. The TEM images show thatamorphous phase is homogeneous. Zr–800K is confirmed to be semi-crystalline with crystal inside in the amorphous matrix. Crystallinestructurewith regular atomic orders is observed in Zr–973K, confirmingthat the Zr–973K is fully crystallized.

3.2. Mechanical properties as a function of temperature

Dynamic mechanical analysis was performed at a frequency of0.3 Hz and constant heating rate of 5 K/min for as-cast Zr56Al16Co28bulk metallic glass (Fig. 3). The temperature dependence of the storagemodulus G′ and the loss modulus G″ can be grouped into three distincttemperature domains similar to the previous studies [23,25]: (1) at lowtemperature, thematerial is in the amorphous state. G′ is high and G″ isclose to zero as observed in Fig. 3. (2) A decrease in G′ and an increase inG″ are observed with increasing temperature, and this temperaturerange is associated with the supercooled liquid region for amorphousmaterials. A maximum of G″ is observed at a temperature Tα. This max-imum corresponds to theα relaxation associated to the glass relaxationprocess. (3) Storage modulus G′ and loss modulus G″ are drastically in-creased with increasing temperature due to the crystallization processat high temperatures.

3.3. Corrosion behavior

Fig. 4 shows polarization curves of as-cast and heat-treated samplesin Ringer's simulated body fluid at room temperature. As-cast BMGshows awider passivated potential region indicating good corrosion re-sistance in the simulated bodyfluid. This is similar to the results of a pre-vious study [18]. As-cast and Zr–660K present the similar passivationpotential suggesting that heat treatment at 660 K does not change the

id. (a) As cast; (b) Zr–660K; (c) Zr–800K; and (d) Zr–973K.

Fig. 7. Hemolytic rate of as-cast and heat-treated samples at 660 K, 800 K and 973 Krespectively.

Fig. 6. CPR values of as-cast and samples heat-treated at 660 K, 800 K and 973 Krespectively.

49R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

corrosion resistance. The passivation potential decreases with increas-ing heating temperature for Zr–800K and Zr–973K. Fig. 5 shows thesurface morphology after electrochemical corrosion. As-cast BMG andZr–660K show a relatively small corrosion region whereas Zr–800Kand Zr–973K exhibit larger corrosion areas.

To quantitatively study the corrosion behavior of as-cast and heat-treated samples, the penetration rate (CPR, μm/year) is calculated byFaraday's law [8,32]:

CPR ¼ 0:327 Micorrð Þ=mρ

where M is the atom-fraction-weighted value of atomic weight, icorr(mA/cm2) is the corrosion current density which is determined byextrapolative procedures of polarization resistance method, m is theion valence and ρ is density.

Fig. 6 shows the CPR values of as-cast and heat-treated samples. TheCPRs increasewith increasing heating temperature. As-cast sample pre-sents the lowest CPR which is 0.072 μm/year whereas Zr–973K showsthe highest CPR of 0.502 μm/year. The corrosion rates are extremelylow suggesting that Zr56Al16Co28 alloys possess excellent corrosion re-sistance. Furthermore, CPRs of the heated samples are much higherthan that of as-cast, indicating that as-cast possesses the best corrosionresistance.

3.4. In vitro hemolysis assay

Fig. 7 shows the results of the in vitro hemolytic test. The hemolyticrates of all the Zr56Al16Co28 alloys are below 5%, suggesting that the as-cast and heat-treated samples will not lead to severe hemolysis regard-less of the microstructural states. This result confirms that the in vitrodegradation of Zr56Al16Co28 alloys has no destructive effect on erythro-cytes, indicating that the Zr56Al16Co28 alloys possess excellent biocom-patibility. Further, the hemolytic rate slightly increases after heatingabove 800 K suggesting that the Zr56Al16Co28 BMGs in the amorphousphase has better hemocompatability than crystallized counterparts.

3.5. In vivo test

The samples were implanted subcutaneously in mice to study thetissue compatibility of the Zr56Al16Co28 alloys. Fig. 8 shows the hemato-logical analysis results at 1 week and 4 weeks after implantation. Theactivities of WBC and Neu were not up-regulated after 1-week and 4-week implantations, suggesting a potentially lower risk of persistent in-flammation (white blood cells: 2.3–13.1 × 109/L; neutrophils: 16.3%–70.2%) [33,34].

The levels of inflammation are also confirmed by hematoxylin andeosin staining of back skin in direct contact with the implanted mate-rials (see Fig. 9). As shown in Fig. 9, under an optical microscope, forsamples heat-treated under 800K, no dysplasia is present in the skin tis-sue, cells adjacent to the implanted materials are neatly arranged withclear structures. At 4 weeks, only a small number of inflammatorycells (dark purple-stained tissue macrophages) have infiltrated the re-gional tissue beside the implanted materials. Whereas, a larger numberof inflammatory cells are present for samples heated above 800 K.

4. Discussion

Heat treatment provides an effective way to modify the mechanicalproperties of BMGs. For the as-cast BMGs heated below Tg, the long-range atomic order and the amorphous nature do not change. However,the mechanical properties change after heat treatment below Tg. Theshear modulus (G) of as-cast Zr56Al16Co28 BMG increases from 34 GPato 36.1 GPa after heat treatment at 660 K which had been confirmedby dynamical analysis in the previous investigation [23]. This is mainlyattributed to that the free volume and atomic mobility are decreased

during the annealing process which enhances the stiffness of theBMGs [23,25,29].

Above Tg, the BMG crystallizes due to the thermally-activated move-ments of atoms. As shown in the results of DSC, two temperatures of crys-tallization (Tc1, Tc2) are present. Below Tc1, partial crystallized andnanoscale crystals appear in the BMGs. This crystalline structure stronglyenhances themechanical properties of the BMGs due to the intrinsic stiff-ness of the crystals. This can be confirmed from Fig. 2 that the shearmod-ulus increases with elevated heat temperature. According to the previousinvestigation [23], the shear modulus of Zr–800K is 37.4 GPa, but theshear modulus of Zr–973K reaches 50.3 GPa which is 46% higher thanthat of the as-cast BMG. Indeed, the shear modulus is supposed to de-crease with increasing temperatures due to thermal softening of BMG athigh temperatures [35–37], but the appearance of crystals induce an in-crease in Young's modulus. Furthermore, the BMG is fully crystallizedabove Tc2. Owing to the high intrinsic modulus of crystals, we can seethat the modulus of as-cast BMG strongly increases above Tc2 as shownin Fig. 3.

Regarding the corrosion behavior in a simulated body fluid, as-castand heat-treated samples present high corrosion resistance. This canbe confirmed from the CPR which possesses maximum value of 0.502μm/year. Indeed, the Zr-based alloys are expected to have ahigh corrosionresistance due to the formation of a highly protective Zr oxide layer at thesurface [38–41]. In addition, the corrosion potential results show that thecorrosion resistance decreases after heat treatment. This means that theBMG possesses the best corrosion resistance compared with crystallizedcounterparts. The corrosion resistance of metallic glasses is expected tobe better than those of their crystallized counterparts according to theprevious studies [8,30,42]. For instance, the corrosion penetration rate ofZr–Cu–Ni–Ti–Al glass is 0.6 ± 0.3 μm/year, but the Zr–Cu–Ni–Ti–Al crys-tallized counterpart is 1.3 ± 0.7 μm/year [8]. This tendency is in accor-dance with our results. It was found that the structure difference

Fig. 8. White blood cell content and neutrophil percentage of blood analysis results formice with implantation of as-cast and heat-treated samples at 660 K, 800 K and 973 K re-spectively. Analysis at 1 week and 4 weeks after implantation.

50 R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

between crystalline and amorphous phase is the main reason inducingthis phenomenon. In crystallized samples, numerous grain boundariesor dislocations appear which have a large volume fraction in the bulksamples [42]. The large volume fraction of intergranular boundaries willsignificantly affect the corrosion resistance of crystalline and nano-crystalline materials by local electrochemically-active sites. Bulk metallicglasses are homogeneous, and grain boundaries or other phase precipi-tates are not present. This homogeneity can drastically reduce the proba-bility of localized corrosion. Furthermore, during immersion in solution,

Fig. 9. Hematoxylin and eosin staining of back skin adjacent to the implanted materials af

the surface composition is different to the bulk due to different diffusionrates of the elements. In the previous study, it was found that the surfacefilms of the immersed Zr–Al–Co BMG are enriched in Zr and Al, and defi-cient in Co [2]. It has been reported that preferentially oxidation of Zr andAl in the surface film of the Zr-based BMGs occurswhen they are exposedto air [19]. Immersion in the corrosive solutions results in the change ofthe composition of the surface films. The corrosion behavior of the Zr–Al–Co–Nb BMGs showed that Zr is further concentrated while Al contentdecreases in the spontaneously passivated films on the alloys afterimmersion in chloride medium compared to those exposed to air [19].From an atomistic viewpoint, the supper corrosion resistance can beattributed to enrich the Zr atoms and Zr oxidation films formed on thesurface. This film also benefits biocompatibility through preventing therelease of toxic elements.

Macrophages, as a component of the immune system, are found toactively respond to most of the biomaterial implants in vivo [43–45].Moreover, macrophages will infiltrate continuously to phagocytose un-digested particles or to fuse into foreign body giant cells and remain onthe surface of the implanted materials. Thus, for those non-degradablebiomaterials, it is essential to choose those that cause less macrophageresponses for long-term implantation [46]. Huang [47] reported that alower level of inflammation was induced by Zr-based BMGs comparedto Ti alloys via morphological analysis and proinflammatory cytokinesecretion measurements. And Li [33] showed that Zr-based metallicglasses, as well as commercially pure Ti, did not cause major systemicinflammatory responses in an oral mucosa irritation test. Similarly,Zr56Al16Co28 alloys heated below 800 K present excellent biocompati-bility in our studies, which is confirmed by both in vitro and in vivoexperiments. Whereas for samples heated above 800 K, though theydid not result in great changes of WBC and Neu, they present a largernumber of infiltrating macrophages in the skin tissue slice. Indeed, itwas hypothesized before that both surface chemistry and proteinadsorption capability of the implanted material played significantroles in the inflammatory andwound-healing responses to biomaterialsand medical devices in vivo [46]. In this case, the altered surface chem-istry might result in different levels of macrophage responses to Zr-based BMGs treated with varying temperatures. However, further stud-ies are still needed to clarify the above hypothesis.

ter 4 weeks of implantation. (a) As cast; (b) Zr–660K; (c) Zr–800K;and (d) Zr–973K.

51R. Wang et al. / Journal of Non-Crystalline Solids 411 (2015) 45–52

Additionally, high corrosion resistance in physiological environ-ments was found to be critical to the biocompatibility of implant mate-rials. High corrosion resistance of alloys can decrease the toxic ionsreleasing in corrosive solution, which results to good biocompatibility[2,48,49]. Many authors had pointed out that, when Zr-based alloysare exposed in the physiological environment, an oxide layer canform,which ismainly consisted of ZrO2 [12,50]. ZrO2 is found to possessvery excellent biocompatibility which promotes the adhesion andgrowth of surrounding cells. Notwithstanding, the Al3+ is believed tobe toxic to the viable cells, but the oxide layer can prevent the releaseof this element. This layer could reach to 9 nm according to a previousstudy [41].

Regarding the influence of the heat treatment, the biocompatibilityand corrosion resistance do not change below Tg. However, when theheat treatment temperature is above Tg, the corrosion resistance de-creases significantly. The appearing crystalline components enhancethe modulus and brittleness of the materials. After heat treatmentabove Tg, Zr56Al16Co28 BMGs lose their elasticity after crystallizationdue to the high stiffness and brittleness of the crystals. This could beharmful for the biomedical application of Zr56Al16Co28 BMGs, especiallyfor orthopedic applications.

5. Conclusion

In conclusion, we investigated the mechanical properties, corrosionresistance and biocompatibility of Zr56Al16Co28 BMG in the as-casted,relaxed, and crystallized states by DMA, XRD, electrochemical corrosionand in vivo/vitro biocompatible tests. The mechanical properties ofZr56Al16Co28 BMGs strongly depend on the heat-treatment temperatureand corresponding microstructures. Heat treatment below the glasstransition temperature Tg induce a slight increase in modulus but a sig-nificant increase in modulus above Tg.

The as-cast BMG presents better corrosion resistance than its crystal-lized counterparts. With an increase of heating temperatures, the pene-tration corrosion rate increases. The Zr56Al16Co28 alloys do not lead tohemolysis. As-cast Zr56Al16Co28 BMG has a better hemocompatabilityand biocompatibility than its crystallized counterparts.

Acknowledgment

The authors would like to appreciate Dr. Jichao Qiao for providingthe sample and the discussion on the mechanical properties. The au-thors are grateful to Bijin Xiong for discussion on paper writing.

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