Materials Science and Engineering C 32 (2012) 665–669

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Materials Science and Engineering C

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In-vitro cytotoxicity and in-vivo biocompatibility of as-extrudedMg–4.0Zn–0.2Ca alloy

Yonghua Xia b,1, Baoping Zhang a,⁎,1, Yin Wang c, Mingfang Qian c, Lin Geng c

a National Engineering Laboratory for Carbon Fiber Technology, Institute of Coal Chemistry, Chinese Academy of Sciences, Taiyuan 030001, PR Chinab The First Affiliated Hospital of Harbin Medical University, Harbin 150001, PR Chinac School of Materials Science and Engineering, Harbin Institute of Technology, Harbin 150001, PR China

⁎ Corresponding author. Tel.: +86 351 4196806.E-mail address: zhangbp@sxicc.ac.cn (B. Zhang).

1 These authors contributed equally to this work.

0928-4931/$ – see front matter © 2012 Elsevier B.V. Alldoi:10.1016/j.msec.2012.01.004

a b s t r a c t

a r t i c l e i n f o

Article history:Received 31 May 2011Received in revised form 28 November 2011Accepted 4 January 2012Available online 11 January 2012

Keywords:BiodegradationCytotoxicityImplantMagnesium

In this paper, the in-vitro cytotoxicity and the in-vivo compatibility of Mg–4.0Zn–0.2Ca alloy are studied. Thecytotoxicity of Mg–4.0Zn–0.2Ca alloy is examined by MTT method on osteoblast cells. The in-vivo behavior ofMg–4.0Zn–0.2Ca alloy is investigated on rabbits. It has been found that Mg–4.0Zn–0.2Ca alloy extract has nocytotoxicity on osteoblast cells. Three months after in-vivo experiment, about 35–38% magnesium alloy im-plant has been degraded and a degradation layer which is composed of Ca, P, O and Mg has been formed onthe magnesium alloy implants. Histological analysis showed that new bone is observed around magnesiumimplant without inflammation reaction. In-vitro and in-vivo test indicated that the Mg–4.0Zn–0.2Ca alloyhas good biocompatibility.

© 2012 Elsevier B.V. All rights reserved.

1. Introduction

Magnesium and magnesium alloys are light metals with charac-teristics including low density, high specific strength and high specificstiffness [1]. In particular, the elastic modulus and compressive yieldstrength of magnesium are close to natural bone, which makesthem a promising substitute for biodegradable implant applicationsin the biomaterial field [2,3]. Many clinical cases as well as in-vitroand in-vivo assessments have demonstrated that magnesium alloyspossess good biocompatibility [4–6]. However, a large number ofmagnesium alloys that have been reported are an extension of the en-gineering material, some of these alloys containing aluminum (Al) orrare earth (RE) elements. The addition of aluminum (Al) or rare earth(RE) elements can effectively improve the corrosion resistance andstrength of Mg alloys [7–9]. However, Al is harmful to osteoblastand neuron [10,11], and RE elements are difficult to be absorbedand excreted by human bodies. In addition, it was found that thepart of RE elements can also lead to hepatotoxicity [12]. Apparently,magnesium alloys containing such elements are not suitable for bio-degradable applications. Thus, other magnesium alloys containingharmless alloying elements have attracted much attention. Mg–Ca[13,14], Mg–Zn [15], Mg–2Zn–0.5Mn [16] and Mg–1.2Mn–1.0Zn[17] fall into this category, and the research shows that these magne-sium alloys are potential biodegradable materials with satisfactorymechanical properties.

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Based on the aforementioned considerations, Zn and Ca have beenchosen as alloying elements and successfully improved the corrosionresistance and mechanical properties of magnesium [18–21]. Howev-er, the biocompatibility of Mg–Zn–Ca has not been studied. In thispaper, the feasibility of Mg–Zn–Ca alloys used as bone implant mate-rials is evaluated by both the in-vitro and in-vivo methods.

2. Experiment

2.1. Preparation of samples

High purity Mg–Zn–Ca alloy containing 3.95 wt.% Zn and 0.21 wt.%Ca (simply designated as “Mg–4.0Zn–0.2Ca” in this report) was pre-pared from high purity Mg (99.99 wt.%), purity Zn (99.8 wt.%), and anMg–26.9 wt.%Ca master alloy by extrusion of the cast billets. The castbillets were obtained by the following routes. Melting and alloying op-erations were carried out in a steel crucible under the protection of amixed gas atmosphere of SF6 (0.3 vol.%) and CO2 (Bal.). Purity Zn andmaster alloy were added into the melt pure Mg at 720 °C. The meltalloy was kept for 10 min at this temperature to ensure that all the re-quired alloying elements were dissolved in the melt alloy. After that,the alloy was cooled down to 700 °C, kept for several minutes andthen poured into a steel mold which had been pre-heated at 200 °C.The billets were homogenized, and then water cooled. The extrudedMg–4.0Zn–0.2Ca alloy rods were obtained at a temperature of 270 °C.The extrusion ratio was 16:1, and the extrusion speed was 2 mm s−1.The rod samples for in-vivo animal study were cut from as-extrudedrods and were 2 mm in diameter and 8 mm in length. Implants weresterilized in alcohol for 5 min and then dried by hot air.

Fig. 1. Morphologies of osteoblast cells cultured for 7 days in different extractionmedia: (a) negative control, (b) 100%, and (c) 50%.

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2.2. Cytotoxicity test

Mouse osteoblast cells (MC3T3-E1, Shanghai Maisha Bio-Technology Ltd. Co. China) were adopted to evaluate the cytotoxicityof Mg–4.0Zn–0.2 Ca alloys. The cells were cultured in Dulbecco'smodified Eagle's medium (DMEM), 10% fetal bovine serum (FBS),100 U mL−1 penicillin and 100 mg mL−1 streptomycin at 37 °C in ahumidified atmosphere of 5% CO2. The cytotoxicity tests were carriedout by indirect contact. Extracts were prepared using DMEM serumfree medium as the extraction medium. The surface area of extractionmedium ratio was 1.25 mL cm−2, whichwas conducted in a humidifiedatmosphere with 5% CO2 at 37 °C for 72 h. The supernatant fluid waswithdrawn and centrifuged to prepare the extractionmedium, then re-frigerated at 4 °C before the cytotoxicity test. The DMEM medium wasused as negative controls group and the extraction medium withoutcell was used as error control. Since, according to the research of Fischeret.al [22], in the case of Mg materials, the use of MTT test kits leads tofalse positive or false negative results. Cells were incubated in 96-wellcell culture plates with 5×104 cells/mL medium in each well and incu-bated for 24 h to allow attachment (100 μL of cell solution/well). Themedium was then replaced by 100 μL of extracts. After incubating thecells in a humidified atmosphere of 5% CO2 at 37 °C for 1, 2, 4 and7 days, respectively, cell morphology was observed by opticalmicroscope (Nikon ELWD 0.3 inverted microscope). 3-(4,5-Dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide (MTT)(Sigma) was then dissolved in phosphate-buffered saline (PBS) at aconcentration of 5 mg mL−1. After that 20 μL MTT was added to eachwell and the sampleswere incubated for 4 h (37 °C, 5% CO2, 95% relativehumidity). Subsequently, 100 μL formazan solutions were added toeach well and optical density (OD) measurements were conducted bymicroplate reader (Bio-RAD680) at 570 nm with a reference wave-length of 630 nm. The cell relative growth rate (RGR) was calculatedaccording to the following formula:

RGR ¼ ODtestwithcell−ODwithoutcellð Þ=ODnegative � 100%

2.3. Surgery

Animal tests were approved by the Ethics Committee of the FirstAffiliated Hospital of Harbin Medical University. The in-vivo degrada-tion experiments were performed in the animal laboratory of the hos-pital. A total of 15 adult New Zealand rabbits (6 females), 2.0–2.5 kgin weight, were used. In the experimental group, sodium pentobarbi-tal (30 mg kg−1) was administered to perform anesthesia by intrave-nous injection. The sterile Mg–4.0Zn–0.2Ca alloy rod sample wasimplanted into the femora of the rabbit; the defects module sizewas 2 mm in diameter and 8 mm in length. Ti6Al4V (>99.99%) wasalso studied as a control.

After operation, all animals received a subcutaneous injection ofpenicillin to avoid a wound contamination and were allowed to movefreely in their cages without external support. After operation, five rab-bits were sacrificed randomly at 1, 2 and 3 months, respectively.

2.4. Degradation and histological analysis

The bone samples with magnesium implants were fixed in 2.5%glutaraldehyde solution and then embedded in epoxy resin for micro-structure analysis. The samples were sliced by hard tissue slicer(ZJXL-ZY-200814-1). Samples were made perpendicular to the longaxis of the implant to get a cross-section of the implant and surround-ing bone tissue. The cross-section microstructure was observed by anoptical microscope (OLYMPUS PEM-3) and a scanning electronic mi-croscope (Hitachi S-5500). The residual implant areas were measuredon the cross-section images using analysis software. The ratio of theresidual cross-section area of implants to the original cross-sectionarea (residual area/implant area×100%) was used to assess the in

vivo degradation rate of magnesium alloys. The element distributionsin the residual implants and the degradation layer after 3 months im-plantation were analyzed.

For histological analysis, the bone samples with magnesium im-plants were fixed in 4% formaldehyde solution, dehydrated, andthen decalcified in ethylene diamine tetra acetate. Then, the speci-mens were embedded in paraffin and cut into films with 5 μm inthickness. The films were then stained with hematoxylin and eosin.Histological images were observed on an optical microscope.

2.5. Statistical analysis

A t-test was used to determine whether any significant differencesexisted between the mean values of the cytotoxicity and animal testsof the experiment. The statistical significance was defined as Pb0.05.

3. Results and discussions

3.1. Cytotoxicity

Fig. 1 showed themorphologies of osteoblast cells cultured in differ-ent extracts after 7 days of incubation. It could be seen from Fig. 1 that

Fig. 2. Relative cell viability (% of control) of osteoblast cells after 1, 2, 4 and 7 days in-cubating in different media.

Fig. 4. Optical images of the cross-sections of bones after 3 months post implantation,(a) Mg–Zn–Ca and (b) Ti6Al4V (M, metal; D, degradation layer; N, new bone; B, bone).

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the cell morphologies in different extracts were normal and healthy,which was similar to that of the negative control. Fig. 2 showed the rel-ative cell viability (% of control) of osteoblast cells after 1, 2, 4 and 7 daysof incubation. There was no significant difference between the relativecell viability in the extracts and those in the negative control. Accordingto standard ISO 10993-5:1999 [23], the cytotoxicity of these extractswas Grade 0–1, indicating that the alloy was safe.

The MTT assay is the most common method employed for the de-tection of cytotoxicity or cell viability. However, in the case of Mgma-terials, Fischer et al. [22] found that the use of these test kits mightlead to false positive or false negative results, because the formationof the tetrazolium salt could lead to a change in colors to affect thevalue of absorbency. In fact,the Mg2+ effects on the MTT tests wereconnected with the concentration. Only the Mg2+ at high concentra-tion would cause false results. In this paper, the Mg–4.0Zn–0.2Caalloy has good corrosion resistance, and there are fewMg2+ in the ex-tracts after 72 h extraction. The extracts without cells have muchlower absorbance values compared with those seeded the cells. Inorder to obtain exact results, we approached the problem from calcu-lation of the absorbance values by subtraction of the results withoutcells from the measurements with cells for each value. The resultswe get in this paper very similar to our previous study, in which thecytotoxicity was tested by neutral red assay [24].

3.2. In-vivo degradation

Fig. 3 was a photograph of magnesium alloy and Ti6Al4V implantsafter 3 months implantation. There was no inflammation reaction.The tissue and the bone were healthy. It was showed that Mg–4.0Zn–0.2Ca alloy implant has good biocompatibility. Fig. 4 showed

Fig. 3. OM images of magnesium alloy and Ti6Al4V implants after 3 months implanta-tion (A, Mg–Zn–Ca; B, Ti6Al4V).

the optical images of the cross-section of the magnesium andTi6Al4V6 implants after 3 months implantation, with Ti6Al4V6 usedas control group. It could be seen that the shapes of the magnesiumimplant had been changed from rod shape to irregular shape, indicat-ing the implant was corroded by the body fluid, or the implant de-graded in the body fluid. Meanwhile, a degradation layer or areaction layer could be clearly found on the surface of the alloy im-plant, as indicated by D in Fig. 4a. In addition, newly formed bonewas observed between the degradation layer and bone tissue aroundthe magnesium alloy implants, as shown by N in Fig. 4a, which wouldbe detail characterized in the histological analysis. The degradationrate was calculated according to the ratios of the cross section areaof the residual implant to the original implant. After 3 months im-plantation, about 35–38% Mg–4.0Zn–0.2Ca alloy implant was degrad-ed. The Ti control group, which as we know has not degraded, after3 months implantations the Ti implants keeps the regular roundshape.

Fig. 5a showed a high magnification microstructure of the boneimplant interface after 3 months implantation by SEM. It could beclearly seen that the degradation layer was not dense, and manycracks were found. These cracks are a preparation-artifact, causedby extensive drying of the samples. Before SEM observation, the sam-ples have highest hydration, and the dehydration was performedwhich could cause the cracking of the degraded layer. In order to re-veal the chemical composition of the degradation layer, EDS was usedto analyze the chemical composition of interface. The results wereshown in Fig. 5b. From the analysis results, it could be figured outthat the degradation layer was mainly composed of carbon, oxygen,magnesium, calcium and phosphorous. However, the chemical com-position was not homogeneous through the whole layer. At the posi-tion close to the Mg implant side, higher calcium content and higherCa/P ratio were found. At the position close to the bone side, the cal-cium content was still high, but the Ca/P ratio became much smaller

Fig. 5. (a) SEM microstructure of the interface between magnesium implant and boneinterface after 3 months post implantation; (b) EDS analysis patterns of implant andbone interface after 3 months post implantation.

Fig. 6. Tissue response of the Mg–Zn–Ca alloy and Ti6Al4V implantation at 1, 2 and 3 mont3 months, (d) Ti6Al4V 1 month,(e) Ti6Al4V 2 months, and (f) Ti6Al4V 3 months. (I, implan

668 Y. Xia et al. / Materials Science and Engineering C 32 (2012) 665–669

than that at the position close to theMg implant side. In fact, our previousstudy indicated that the release of the Mg2+, Zn2+ and Ca2+ was veryfew, which is safe during degradation process [24]. Therefore, the Zn2+

released during degradation is not mentioned in this article. The distribu-tion of the Ca2+ was very important in these degrade process, it is notonly from the degraded of the Mg–4.0Zn–0.2Ca implants but also fromthe physiological environment, which determines the Ca/P ratio and thenewbone growth. Therewas also having a sharp change inMg2+ contentat the interface. However, it has few of theMg in the bone, indicating thatMg2+ can be absorbed and excreted by human bodies.

3.3. Histological analysis

Fig. 6 showed the tissue response to the Mg–4.0Zn–0.2Ca alloyand Ti6Al4V pins implantation at 1, 2 and 3 months. It could be clear-ly seen that some lymphocytes were identified in histological tissuein 1 month after operation, but there was no visible evidence of mul-tinucleated giant cells in both implants. After 2 months implantation,in the magnesium group there was an active bone formation, whichwas evident by large number of new disorganized trabeculae. After3 months implantation, new bone tissue was formed around the mag-nesium implant. In comparison with the histological microstructureobtained at the cortical bone near the implantation site, as shown inFig. 6, no difference could be found in the histological microstructurebetween the new bone and the cortical bone. The tissue response inthe Ti6Al4V group showed similar results to those reported indocuments.

Magnesium alloys have attracted much attention as potential bio-degradable bone implant materials due to their biodegradability inthe bioenvironment as well as their excellent mechanical propertiessuch as high strength and an elastic modulus close to that of bone.In this paper, the in-vitro cytotoxicity and in-vivo biocompatibilityof new kind of Mg–4.0Zn–0.2Ca alloy were studied. The cytotoxicitytest indicated that the Mg–4.0Zn–0.2Ca alloy had no cytotoxicity.Rabbit implantation indicated that the Mg–4.0Zn–0.2Ca alloy didnot cause any inflammation reaction. One month after operation, allmagnesium implants were fixed tightly. There was no gap betweenthe bone and the residual implant. Optical images from Fig. 4 andSEM microstructure from Fig. 5 showed clearly that there was a deg-radation layer formed on the surface of the magnesium implants.

hs: (a) Mg–4.0Zn–0.2Ca 1 month, (b) Mg–4.0Zn–0.2Ca 2 months, (c) Mg–4.0Zn–0.2Cat situated; L, lymphocytes; N, new bone; B, bone; O, osteoporosis.)

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Histological images showed that new bone tissue was in contact withthe magnesium implant through this degradation layer. Line scanningin Fig. 5 proved that large amounts of Ca and P were found around themagnesium implant. The histological analyses revealed that thismagnesium-containing calcium phosphate degradation layer couldpromote or accelerate the new bone formation.

In-vivo degradation was a very complex process, so it was difficultto accurately assess the degradation rate of an implant material. Witteet al. [25] found that the corrosion of a magnesium rod in the medul-lar cavity was not homogeneous in all cross-sections by using micro-computed tomography. In our study, the cross-section area of the re-sidual implant was calculated to describe the degradation rate of theMg implant. Due to the inhomogeneous corrosion of a magnesiumrod in the medullary cavity, the calculated degradation rate basedon the images was similar to the real in-vivo degradation rate of mag-nesium implants. In-vivo degradation, compared with other metalimplants, is an ultimate merit of magnesium alloy. After implantationin the rabbit, Mg–4.0Zn–0.2Ca alloy would be reacted with body fluidon the surface and get dissolved in the surrounding body fluid. Atfirst, the released Mg2+, Zn2+ and Ca2+ could be absorbed by the sur-rounding tissues and excreted through the gastrointestinal route andthe kidney. However, with the increasing time of implantation, moreMg2+, Zn2+ and Ca2+ ions are dissolved into the solution, and thelocal pH near the surface of the Mg implants could be >10 [26]. Asa result, an insoluble magnesium-containing calcium phosphatewould be precipitated from the body fluid on the surface of the mag-nesium implant and tightly attached to the matrix, which retardeddegradation. In addition, the corrosion layer on the Mg–4.0Zn–0.2Caalloy contained Mg, Ca and P, which could promote osteoinductivityand osteoconductivity, predicting good biocompatibility of magne-sium. Therefore, it is proposed that theMg2+, Zn2+ and Ca2+ releasedduring degradation should be safe.

4. Conclusions

In this paper, the feasibility of Mg–4.0Zn–0.2Ca alloy used as boneimplant materials was evaluated by in-vitro and in-vivo methods. Thecytotoxicity of Mg–4.0Zn–0.2Ca alloy was examined by MTT methodon osteoblast cells. It is found that Mg–4.0Zn–0.2Ca alloy extractshave no cytotoxicity on osteoblast cells. The in-vivo behavior ofMg–4.0Zn–0.2Ca alloy was investigated on rabbits. The alloy did notinduce inflammation reactions nor affect the new bone formation.Three months after in-vivo experiment, about 35–38% magnesium

alloy implant was degraded, and a degradation layer which was com-posed of Ca, P, O and Mg was formed on the magnesium alloy im-plants. Histological analysis showed that new bone formed aroundmagnesium implant. In-vitro and in-vivo test indicated that the Mg–4.0Zn–0.2Ca alloy had good biocompatibility.

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