Mechanical properties, in vitro degradation behavior, hemocompatibility and cytotoxicity evaluation of Zn–1.2Mg alloy for biodegradable implants

Abstract

Zn, which is a promising alternative candidate as a biodegradable implant material, possesses excellent biocompatibility and biodegradability. However, the insufficiency of its strength and hardness has largely limited its application. Nevertheless, adding alloying elements and mechanical forming by extrusion have generally enhanced its mechanical properties. In the present work, Zn–1.2Mg alloy has been designed and treated by extrusion. Experimental results demonstrated that the studied alloys were composed of a matrix of Zn and a precipitated phase of Mg₂Zn₁₁, and the grain size became smaller and more homogeneous after extrusion. The as-extruded alloy exhibited much higher yield strength (YS 219.61 MPa), ultimate tensile strength (UTS 362.64 MPa), elongation (21.31%) and hardness (96.01 HV). The corrosion rates of the as-extruded alloy were higher compared with those of the as-cast alloy and reached values of 0.19 mm per year in electrochemical tests and 0.11 mm per year after exposure in Hank's solution for 30 days, respectively. Moreover, the as-extruded alloy displayed excellent hemocompatibility (hemolysis rate of 1.85%, superior thromboresistance and no signs of thrombogenicity). The viability of human osteosarcoma HOS cells and MG63 cells cultured in diluted extracts of the alloy exceeded 70%, which demonstrated no potential cytotoxicity and tolerance in cellular applications.

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1. Introduction

Metals are widely used as biomedical materials for orthopedic implants and cardiovascular interventional devices, owing to their higher strength and toughness compared with polymers and ceramic materials.^1–3 Nevertheless, traditional metallic biomaterials with high corrosion resistance such as stainless steels, cobalt–chromium alloys, tantalum and titanium or its alloys are generally used as permanent implants in patients.^2–4 Once these are implanted in a patient's body, metal ions might be released from partially implanted devices for a long time, causing a local inflammatory response in some organs or tissues. Hence, patients required reoperation after they were cured or the implants were no longer needed, so that the intense pain and extra expense of a second surgical procedure might be involved.³

Biodegradable metals with good mechanical properties, excellent degradability and biocompatibility have recently been suggested as a class of promising metallic biomedical materials for orthopedic implants, attracting a great amount of interest.^4–11 Magnesium and its alloys have attracted much attention as biodegradable metals for implants owing to their similar mechanical properties to those of the body's natural bone and the biocompatibility of magnesium ions.^4–7 Nevertheless, one of the most significant risks of Mg-based biodegradable devices is that they are corroded so rapidly. It has generally been acknowledged that Mg-based biodegradable metals with a very low corrosion potential (−2.36 V_SCE), which exhibit a rapid corrosion rate in physiological environments, cannot offer effective support and their mechanical integrity is compromised prior to recovery after implantation in the body.^4,7,10

Compared with Mg-based alloys, Zn and its alloys might be considered to be interesting candidates owing to their higher electrode potential and similar mechanical properties.¹² Furthermore, Zn element is generally considered to be an essential element for humans, playing an important role in the metabolic processes of the human body.¹³ After these factors were considered, Zn and its alloys were suggested as new members of the family of biodegradable materials. However, preliminary studies have demonstrated that pure Zn had poor mechanical properties and hardness (the tensile strength of as-cast pure Zn was only 20 MPa, its elongation was 0.3% and its hardness was approximately 25 HV^9,14), which might not offer sufficient support for implanted devices. Therefore, as an indispensable strengthening process, alloying has been adopted to improve their mechanical performance. Many novel types of zinc alloys have been designed and developed, with the main concentration on Zn–Mg,^9,14–19 Zn–Ca,¹⁹ Zn–Sr,¹⁹ Zn–Mg–Ca,^20,21 Zn–Mg–Sr,^20–22 Zn–Ca–Sr²⁰ and Zn–Mg–Mn alloys.²³

As an effective alloying element in zinc alloys, Mg was found to prominently improve their mechanical performance, and the addition of 0.93 wt% or 1.55 wt% Mg to pure Zn provided superior tensile strength and hardness.^9,14 Nevertheless, the maximum elongation of Zn–Mg alloys was approximately 1.5%, which might be suggested as indicating poor ductility for implanted devices. However, advanced processing technology such as extrusion,^16,18–20 rolling^19,20,22,23 and high-pressure torsion (HPT)²⁴ could be employed to greatly improve the mechanical properties of Zn-based biodegradable metals, which were significantly superior to those of the well-known Mg-based biodegradable metals that have been developed.^23,25 In the present work, Zn–1.2Mg alloy has been designed and, in consideration of its effective composition between Zn–1Mg alloy and Zn–1.5Mg alloy, its development as a biodegradable metal and its feasibility for implants were evaluated, in particular after hot extrusion. Its microstructure, mechanical properties, in vitro degradation behavior, in vitro hemocompatibility and cytotoxicity were tested to investigate its feasibility as a new biodegradable material.

2. Materials and methods

2.1 Materials preparation

Zn–Mg alloy with the nominal composition of Zn–1.2Mg was melted at 470–520 °C under a protective atmosphere of CO₂ with pure Zn (99.99%, Huludao Zinc Industry Co., China) and pure Mg (99.99%, Henan Yuhang Metal Materials Co., China) as the raw materials. After holding and stirring for about 30 min, the molten metal was cast into a steel mold at about 430 °C. Thereafter, ingots were hot-extruded. In the process of extrusion, the samples were preheated to 250 °C for 2 h and then extruded at an extrusion ratio of 36 : 1 at 250 °C. The actual chemical composition of the alloy ingots determined by inductively coupled plasma atomic emission spectroscopy (ICP-AES, Varian 715) was Zn–1.22% Mg (wt%).

Specimens for microstructure characterization and in vitro measurements were cut from the ingots with a geometric size of 10 mm × 10 mm × 2.5 mm. However, samples cut from the extruded bars had a thickness of 2.5 mm and a diameter of 10 mm parallel to the extrusion direction. Besides, specimens for microstructure characterization were also obtained along the direction of hot extrusion. All the specimens were polished to 3000 grit, followed by being ultrasonically cleaned in acetone, absolute ethanol and distilled water and then dried in the open air.

2.2 Microstructure characterization

An X-ray diffractometer (XRD, X'Pert Pro) with Cu Kα radiation was employed for the identification of the constituent phases of the studied alloy. Diffraction patterns were generated in the range of 10–90° at a scanning speed of 4° min⁻¹.

Evaluation of the microstructure of the studied alloys was carried out using a scanning electron microscope (SEM, Quanta 200) with an energy-dispersive spectrometer (EDS). All the samples for microscopic observation were polished and etched with an acid solution.

2.3 Mechanical properties testing

Samples for mechanical testing were machined according to ASTM E8/E8m-11.²⁶ The mechanical tests were carried out using an Instron 5969 universal materials testing machine with a crosshead speed of 1.0 mm min⁻¹. At least four measurements were made for each group. The fracture morphology was investigated by scanning electron microscopy (SEM, Quanta 200).

The Vickers microhardness was determined using a HMV-2T hardness tester with an applied load of 100 g and a dwell time of 15 s. At least three parallel specimens were selected for each group and six fields for collection were made for each sample.

2.4 In vitro degradation measurement

The in vitro corrosion resistance of the alloy was determined by a potentiodynamic polarization test and an immersion test. The measurement of the corrosion resistance was carried out in Hank's solution.⁹

2.4.1 Electrochemical tests. A three-electrode cell was used for measurements on an electrochemical workstation (Parstat 2273). The container was filled with 200 mL Hank's solution and the temperature was controlled at 37.5 ± 0.5 °C using a water bath. A platinum electrode was used as the auxiliary electrode, a saturated calomel electrode (SCE) as the reference electrode, and specimens of the studied alloy as the working electrode, respectively. All samples were connected to a copper sheet and the area of the working electrode exposed to the electrolyte was 0.283 cm². The open-circuit potential (OCP) of each sample was continuously monitored for 3600 s in Hank's solution to achieve stabilization. Then, potentiodynamic polarization tests were conducted using a scan rate of 1 mV s⁻¹ and a scan range from −1600 to −400 mV. The corrosion current density was calculated by extrapolating the polarization curve and the corrosion rate (V_corr) was calculated according to ASTM G102-89.²⁷ The surface morphology of the specimens after electrochemical testing was further observed by SEM.

2.4.2 Immersion tests. Immersion tests were carried out in Hank's solution according to ASTM G31-72.²⁸ The ratio of the surface area to the solution volume was 1 cm² : 25 mL. After periods of immersion, the samples were removed from Hank's solution, rinsed with distilled water and then dried at room temperature. The corrosion products were removed by chemical reagents according to ISO 8407:2009.²⁹ For the studied alloy, a solution of 200 g L⁻¹ CrO₃ and 10 g L⁻¹ AgNO₃ was used for this purpose. The surface morphology of the studied samples after immersion in Hank's solution for 30 days was observed using SEM. The corrosion rate was calculated according to ASTM G31-72.²⁸

2.5 Blood compatibility test

The in vitro hemocompatibility was investigated by hemolysis rate tests, measurements of dynamic clotting time and observations of platelet adhesion. Healthy blood obtained from volunteers and collected by Fuxing Hospital (Capital Medical University, Beijing) was mixed with sodium citrate (3.8 wt%) in a ratio of 9 : 1 and diluted with physiological saline (0.9 wt%) in a volume ratio of 4 : 5. Measurements of the hemolysis rate and dynamic clotting time were carried out in these diluted blood samples.

2.5.1 Hemolysis rate tests. Specimens of the studied alloy were immersed in centrifuge tubes containing 10 mL physiological saline and incubated at 37 ± 0.5 °C for 30 min. Then, 0.2 mL diluted blood was added to these tubes and the mixtures were incubated at 37 ± 0.5 °C for 60 min. Thereafter, the specimens were removed and all the mixture solutions were centrifuged at 3000 rpm for 5 min. After this period, the supernatant from each tube was transferred to a cuvette using a pipette, where the absorbance was calculated using an ultraviolet spectrophotometer (Agilent 8453) at 545 nm. Physiological saline solution was used as a negative control and ultrapure water as a positive control. An average of three measurements were made for each group. The hemolysis rate was calculated as follows:

Hemolysis = (D_t − D_nc)/(D_pc − D_nc) × 100%

where D_t is the optical density of the tested group and D_nc and D_pc are the optical densities of the negative and positive control groups, respectively.

2.5.2 Dynamic blood clotting tests. A series of specimens in the study were prepared at the bottom of a 24-well plate and the plate was placed in a thermostatically controlled environment at 37 °C for 15 min. Then, 100 μL diluted blood was dropped onto the surface of a specimen, followed by the addition of 10 μL 0.2 M CaCl₂ solution, and the solution was mixed uniformly. The plate containing blood was kept in the thermostatically controlled environment at 37 °C for a given time. In our experiments, periods of 15, 30, 45, and 60 min were selected. Then, the specimen was put into a centrifuge tube containing 5 mL distilled water for 5 min, and then the absorbance of the solution was recorded with an ultraviolet spectrophotometer (Agilent 8453) at 545 nm. For each specimen, the average absorbance was obtained from three measurements.

2.5.3 Platelet adhesion tests. Platelet-rich plasma (PRP) was prepared by centrifuging fresh whole blood containing sodium citrate (3.8 wt%) for 10 min at a rate of 1000 rpm. Specimens of the studied alloy were immersed in PRP, incubated at 37 °C for 60 min, and then rinsed with physiological saline to remove non-adhered platelets. The adhered platelets were fixed in 2.5% glutaraldehyde solution for 60 min at room temperature, followed by dehydration in a gradient mixture of ethanol/distilled water (50%, 60%, 70%, 80%, 90%, 95%, and 100%) for 10 min each, and then dried in the air. The morphology of the adhered platelets on the experimental alloy plates was observed by SEM.

2.6 Cytotoxicity tests

For cytotoxicity tests, specimens of the studied alloy were sterilized by Co60 γ-radiation sterilization technologies. Human osteosarcoma HOS cells and MG63 cells obtained from the Cell Resource Center of the Shanghai Institutes for Biological Sciences were used for the evaluation of cytotoxicity via an MTT assay. HOS cells were cultured in RPMI-1640 medium and MG63 cells were cultured in DMEM (Dulbecco's modified Eagle medium). Both media contained 10% fetal bovine serum (FBS, Zhejiang Tianhang Biotechnology Co., Ltd), 100 U mL⁻¹ penicillin and 100 μg mL⁻¹ streptomycin and were cultured in 96-well flat-bottomed cell culture plates for 24 h at 37 °C in a humidified atmosphere of 5% CO₂ to allow attachment. An extraction medium was prepared using medium with a surface area/extraction medium ratio of 1.25 cm² mL⁻¹ for 48 h and then serially diluted to concentrations of 75%, 50%, 25% and 12.5% with fresh medium. Then, 100 μL extracts was then replaced with medium from the plates and incubated for 3 days. Next, 10 μL MTT (3-(4,5-dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazolium bromide) was added to each well and incubated for 4 h at 37 °C. After that, 150 μL DMSO (dimethyl sulfoxide) was added. The absorbance was measured by a microplate reader (Bio-Rad iMark) at 570 nm with a reference wavelength of 630 nm. The medium was used as a negative control and the medium with 5% DMSO as a positive control.

2.7 Statistical analysis

SPSS 18.0 software was used to analyze the obtained data. The values from mechanical testing, measurements of in vitro corrosion and evaluation of in vitro hemocompatibility were averaged (n > 3). Student's t-test was performed to determine statistical significance and differences were considered to be statistically significant at values of p < 0.01.

3. Results and discussion

3.1 Microstructures

Fig. 1 shows the XRD patterns of as-cast Zn–1.2Mg alloy and as-extruded Zn–1.2Mg alloy. It can be seen that both patterns are similar in shape. The result shows that Zn–1.2Mg alloy is composed of a matrix of Zn with a hexagonal close-packed (hcp) crystal structure and a precipitated phase (Mg₂Zn₁₁).³⁰

Fig. 1 X-ray diffraction patterns of (a) as-cast Zn–1.2Mg alloy and (b) as-extruded Zn–1.2Mg alloy.

Fig. 2 shows the microstructure of the studied alloys. As we can see from Fig. 2(a), a typical micrograph of the studied samples consists of dendrites, primary grains and eutectic mixtures.^9,14,20 Quantitative analysis by EDS (as shown in the inset) shows that Zn and Mg elements are predominantly present as eutectic mixtures. After hot extrusion (shown in Fig. 2(b), perpendicular to the extrusion direction), the grain sizes became more homogeneous and smaller, and the volume fraction of eutectic mixtures (as shown in the inset, proved by EDS quantitative analysis) decreased. The XRD results for the as-extruded alloy in Fig. 1 also demonstrate grains with smaller sizes, whereas peaks with smaller FWHM (full width at half maximum) values from the Zn matrix were detected. Nevertheless, as we can see from Fig. 2(c) (parallel to the extrusion direction), strip grains with more eutectic mixtures precipitated along the grains are visible, and the results of EDS quantitative analysis of the eutectic mixtures (as shown in the inset of Fig. 2(c)) indicate that Zn and Mg exist predominantly in the precipitated phase. Obviously, the grain pattern has a fibrous structure along the extrusion direction, which indicates that the grains are distorted and elongated. The grain shape and boundary are adjusted during hot working, which results in the occurrence of cracking and stretching of dendrite walls along the extrusion direction.

Fig. 2 SEM images showing micrographs of (a) as-cast Zn–1.2Mg alloy, (b) as-extruded Zn–1.2Mg alloy perpendicular to the extrusion direction, and (c) as-extruded Zn–1.2Mg alloy parallel to the extrusion direction. The insets show the results of EDS performed on alloy samples marked by the red frames, respectively.

3.2 Mechanical properties

Table 1 shows the tensile properties and microhardness of samples of the studied alloys. It is obvious that as-cast Zn–1.2Mg alloy has a quite low yield strength (YS), ultimate tensile strength (UTS), elongation and microhardness of 116.54 ± 1.20 MPa, 129.56 ± 6.38 MPa, 1.44 ± 0.63% and 92.68 ± 6.72 HV, respectively. This might be attributed to more brittle eutectic mixtures, which act as regions of stress concentration, facilitate the growth of fracture cracks and display no plastic strain during loading for the as-cast alloy. However, after hot extrusion, the tensile properties and microhardness are significantly enhanced, which suggests the effectiveness of thermal deformation on the improvement in the mechanical properties. It is more encouraging that the ultimate tensile strength and elongation of as-extruded Zn–1.2Mg alloy are three times greater than those of the as-cast alloy, increasing to 362.64 ± 4.87 MPa and 21.31 ± 2.31%, respectively, which indicates superior mechanical properties to those of previously reported magnesium alloys for biomedical applications.^23,25 Furthermore, Kubásek et al.,¹⁵ Gong et al.¹⁸ and Li et al.¹⁹ prepared Zn–1Mg alloys in the extruded state with much lower values of mechanical properties compared with those in the present study.

Table 1 Data for tensile properties and microhardness of the studied alloy

Zn–1.2Mg alloy	Tensile properties			Microhardness (HV)
	YS (MPa)	UTS (MPa)	Elongation (%)
a Indicates p < 0.01 when compared with the as-cast alloy.
As-cast	116.54 ± 1.20	129.56 ± 6.38	1.44 ± 0.63	92.68 ± 6.72
As-extruded	219.61 ± 15.42^a	362.64 ± 4.87^a	21.31 ± 2.31^a	96.01 ± 6.58

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Fig. 3 presents the typical stress–strain curves and fracture surfaces of the studied alloy samples. As we can see from Fig. 3(a), it is worth noting that the typical stress–strain curve of as-extruded Zn–1.2Mg alloy displays a wide yielding plateau, and the elongation ratio is greater than 20%, which might be attributed to the occurrence of plastic deformation. As shown in Fig. 3(b), there is an obvious style of cleavage fracture with cleavage planes and coarse grain boundary delamination on the fracture surface of as-cast Zn–1.2Mg alloy, which might be because of the metallographic microstructure filled with coarse dendrites, primary grains and eutectic mixtures (as shown in Fig. 2(a)). Nevertheless, the fracture surface exhibits notable differences and the feature of ductile fracture with shear lips and a fibrous region corresponds to the elongation of 21.31% measured for the as-extruded alloy. A quite different fracture morphology is observed, which is mostly due to the occurrence of cracking and stretching of dendrite walls along the extrusion direction and the adjustment of the grain shape and boundary during hot working, which is confirmed by the metallographic microstructure, with a homogeneous and smaller grain shape and size (as shown in Fig. 2(b)).

Fig. 3 (a) Typical stress–strain curves of the studied alloys and typical fracture surfaces of (b) as-cast Zn–1.2Mg alloy and (c) as-extruded Zn–1.2Mg alloy after tensile tests at room temperature.

3.3 Corrosion behavior

Table 2 presents the corrosion parameters obtained from the potentiodynamic polarization tests and immersion tests. The corrosion rates of the as-extruded alloy were higher compared with those of the as-cast samples and reached values of 0.19 mm per year (V_corr), 0.11 mm per year (V after exposure to Hank's solution for 30 days) and 0.09 mm per year (V after exposure to Hank's solution for 90 days). It is obvious that the corrosion rate of the studied alloy increased by at least 30% after extrusion. This could be related to the metallographic microstructure of the as-extruded alloy, in which the Mg₂Zn₁₁ phase is distributed relatively uniformly among the grain boundaries and in the matrix, and has been involved in the formation of galvanic micro-cells, which then accelerated the corrosion of the substrate.⁹ Moreover, the corrosion rate of the studied alloy from the potentiodynamic polarization tests was significantly higher than that from the immersion tests. The reason for this is that when the potentiodynamic polarization tests were carried out, the surface of the sample had undergone a strong polarization effect. In particular, precipitation led to strong galvanic corrosion forces under the influence of the polarization effect. Furthermore, one last thing to note is that an overdose of Zn may cause systemic toxicity and the optimal daily dose is 15 mg,³¹ although zinc is an essential element for humans. As the experimental data show, the corrosion rate of the extruded alloy was about 0.11 mm per year, i.e., 0.21 mg per cm² per day. The studied alloys are considered for implant applications, and the typical model of implants is assumed to have a shape 4 mm in diameter and 25 mm in length. In this case, implants in the body may provide a Zn dose of 1.53 mg per day, which is far below the suggested intake of 15 mg per day; thus, Zn–1.2Mg alloy could be considered as a biodegradable metal with appropriate corrosion resistance.

Table 2 Corrosion rates of the studied alloy obtained from potentiodynamic polarization tests and immersion tests

Zn–1.2Mg alloy	Electrochemical tests			Immersion tests V (mm per year)
	Ecorr (V)	Icorr (μA cm^−2)	Vcorr (mm per year)	30 d	90 d
As-cast	−1.18 ± 0.01	7.68 ± 0.19	0.12	0.08 ± 0.01	0.07 ± 0.01
As-extruded	−1.20 ± 0.01	12.38 ± 1.83	0.19	0.11	0.09 ± 0.02

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Fig. 4 shows the surface morphologies of the studied alloy after electrochemical testing and immersion testing. The surfaces of the studied alloy are severely corroded with localized pitting (as shown in Fig. 4(a) and (b)), which is probably due to attack by the anode and then the corrosion is exacerbated by a galvanic corrosion reaction. Moreover, the surfaces of the studied alloy are covered with precipitates after electrochemical testing, and the EDS results indicate the presence of the elements C, O, P, Cl, Ca and Zn. The presence of these elements has also been detected on the surface of the studied alloys after exposure in Hank's solution, which is due to the corrosion products containing zinc phosphates, zinc carbonates and zinc hydroxide.³² Furthermore, another study also found similar corrosion products when Zn–Mg alloys (with a Mg content of 1–16 wt%) were immersed in a highly reactive chloride-containing solution, which consisted of hydrozincite (Zn₅(OH)₆(CO₃)₂) and traces of simonkolleite (Zn₅(OH)₈Cl₂·H₂O).³⁰ Fig. 4(c) and (d) illustrate the surface morphologies of the studied alloys after removal of the corrosion products. It is evident that the surfaces display severely corroded morphologies with a larger corroded area, which might be caused by the presence of high concentrations of chloride ions in Hank's solution. The highly reactive ions disrupted the equilibrium of oxide film on the surfaces, after which a more active secondary phase was exposed to Hank's solution, resulting in a galvanic corrosion reaction.

Fig. 4 SEM micrographs of the surface morphologies of (a) as-cast Zn–1.2Mg alloy and (b) as-extruded Zn–1.2Mg alloy and the inset showed the EDS results corresponding to the sample after electrochemical measurements, and surface morphologies of (c) as-cast Zn–1.2Mg alloy and (d) as-extruded Zn–1.2Mg alloy with the removal of corrosion products after immersion in Hank's solution for 30 days.

3.4 Hemocompatibility

Fig. 5 shows investigations into the hemocompatibility of the studied alloy. As shown in Fig. 5(a), the as-cast alloy induced a hemolysis rate of 1.62% when in contact with the diluted blood. However, after hot extrusion the hemolysis rate increased, displaying a higher hemolysis percentage of 1.85%. However, the hemolysis rates for the studied alloy are far lower than 5%, which is a judging criterion for the blood compatibility of biomaterials, according to ASTM F756-00.³³ The hemolysis rate represents the proportion of red blood cells disrupted by the sample when in contact with blood. Therefore, it is suggested that the in vitro degradation of Zn–1.2Mg alloy has no destructive effect on erythrocytes. The blood clotting behavior of the studied alloy is shown in Fig. 5(b). Obviously, the absorbance values of the hemolyzed hemoglobin solution decrease with time. The slower the optical density declines with time, the longer is the clotting time. It is clearly indicated that the studied alloys have changed slightly. Moreover, the higher absorbance values correlate with improved thromboresistance of the material.³⁴ That is to say, the studied alloy exhibits better blood compatibility. Typical SEM images with the adhesion of platelets after incubation in PRP for 1 h are shown in Fig. 5(c) and (d). The platelets on the surface of the experimental specimens retained a nearly round shape and remained isolated, showing an early stage of activation, which means that there are no signs of thrombogenicity with the studied alloy. In brief, the studied Zn–1.2Mg alloy displays superior blood compatibility.

Fig. 5 Hemocompatibility of the studied alloys: (a) hemolysis percentages and (b) blood clotting profiles on the alloy surfaces. SEM micrographs of platelets adhering to as-cast Zn–1.2Mg alloy (c) and as-extruded Zn–1.2Mg alloy (d).

3.5 Cytotoxicity

Fig. 6 illustrates the viability of HOS cells and MG63 cells cultured in the extraction medium at concentrations of 100%, 75%, 50%, 25% and 12.5% for 3 days. It can be seen that the cell viability is positively influenced by the concentration of the extracts. For HOS cells, a higher extract concentration led to a significant reduction in cell viability and induced cytotoxicity, where the relative cell viability was below 10%. In contrast, for MG63 cells high extract concentrations had a slight negative impact on cell viability, and extracts with concentrations of 50% and 25% even contributed to an increase in cell viability, which exceeded 100% of that of the negative control group. It could be inferred that a certain level of Zn ions might promote cell attachment and proliferation, which in turn resulted in increases in bone healing and new bone formation.¹⁹ In addition, Jun Ma et al.³⁵ evaluated the cellular behavior of HCECs exposed to extracellular Zn²⁺, and the results revealed that Zn²⁺ promoted cell viability, proliferation, spreading, and migration at low concentrations, but had an inhibitory effect at high concentrations. Furthermore, compared with the as-cast alloy, as-extruded Zn–1.2Mg alloy gave rise to a lower relative cell viability, which was due to its higher corrosion rate resulting in a higher concentration of Zn ions (as shown in Table 2). Nevertheless, according to the recommendation of Jiali Wang et al.,³⁶ use of the current ISO 10993 standards^37,38 for predicting the potential health risks of degradable Mg-based biomaterials via cytotoxicity tests is not desirable or justified, and their research recommended a minimum sixfold to a maximum tenfold dilution of extracts for the in vitro cytotoxicity testing specified in part 5 of ISO 10993 (ref. 37) of potential orthopedic implants. Moreover, as a new biodegradable metal, zinc could be subjected to the same principle. In the current study, all the viability values for cells cultured in extracts with a concentration of 12.5% (sevenfold dilution of extracts) exceeded 70%,³⁶ which thereby demonstrated no potential cytotoxicity and tolerance in cellular applications.

Fig. 6 Viability of HOS cells and MG63 cells after incubation in extracts of the studied alloys: (a) as-cast and (b) as-extruded Zn–1.2Mg alloy.

4. Conclusions

In order to overcome the low strength and hardness of pure Zn, alloying with Mg has been adopted, and, in consideration of its effective composition between Zn–1Mg alloy and Zn–1.5Mg alloy, Zn–1.2Mg alloy has been developed as a biodegradable metal. Moreover, hot extrusion treatment was also employed as an effective method for improving its mechanical properties, and Zn–1.2Mg alloy was treated by extrusion to further modify its properties. Following this, its microstructure, mechanical properties, in vitro degradation behavior, in vitro hemocompatibility and cytotoxicity were systematically studied to investigate its feasibility as a new biodegradable metal. The YS, UTS, elongation and hardness were largely improved after hot extrusion (219.61 MPa, 362.64 MPa, 21.31% and 96.01 HV, respectively). The as-extruded alloy displayed appropriate corrosion resistance, and its corrosion rate reached 0.19 mm per year in electrochemical tests and 0.11 mm per year after exposure in Hank's solution for 30 days. Moreover, the results of evaluation of hemocompatibility revealed that the as-extruded alloy exhibited excellent hemocompatibility (low hemolysis rate, superior thromboresistance and no signs of thrombogenicity). Moreover, the viability of HOS cells and MG63 cells cultured in diluted extracts of the alloy exceeded 70%, which demonstrated no potential cytotoxicity and tolerance in cellular applications. Based on these findings, as-extruded Zn–1.2Mg alloy has been shown to be a potential candidate as a biodegradable implant material owing to its superior mechanical performance, appropriate corrosion resistance and in vitro biocompatibility.

5. Statement

The research into biocompatibility was conducted in accordance with the Guide for Biological evaluation of medical devices (GB/T 16886, revised 2005) published by AQSIQ (General Administration of Quality Supervision, Inspection and Quarantine of the People's Republic of China) and SAC (Standardization Administration of the People's Republic of China). All experimental protocols were approved by the Ethics Committee of Fourth Military Medical University.

Acknowledgements

This work was supported by the research funds of Lepu Medical Technology (Beijing) Co., Ltd, the National Natural Science Foundation of China (No. 51271199 and No. 51501223), and the National High-tech R&D Program (863 Program, No. 2015AA033702).

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Footnote

† These two authors contributed equally to this work.

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