Biodegradable Zn-0.05Mg-(0.5, 1 wt.%) Ag alloy was manufactured by indirectly extruding the alloy ingot at 200 °C with an extrusion ratio of 16:1. Dynamic recrystallization took place during the extrusion process, leading to the formation of equiaxed crystals with twins in both cross-sectional and longitudinal direction. There was no detectable Ag-related phase present except the Mg2Zn11 in the alloys. Tensile strength was increased with an increase in Ag content, reaching 202 MPa when Ag content is 1 wt.%. As-extruded Zn-0.05Mg-0.5Ag showed better corrosion performance with a low corrosion current density of 2.2 A/cm2 and low corrosion rate of 0.15 mm/year. The antibacterial property improved for both Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli) by addition of Ag. The antibacterial rates were more than 99% when Ag content is up to 1 wt.%. The biodegradable Zn-Mg-Ag alloys with high antibacterial behavior show great potential in medical devices.
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Introduction
Zn, as an essential element in human nutrition, dominates a vital role in cell proliferation and growth, wound healing as well as in the immune and nervous systems. Zn also supports normal growth, wound healing, and a proper sense of taste and smell. It interacts with a wide range of organic ligands and plays roles in the metabolism of RNA and DNA, signal transduction, and gene expression. It also regulates apoptosis. The oxide of Zn shows good antibacterial performance under the visible and ultraviolet radiation. Zn and its alloy have been becoming the most potential materials in bio-absorbable implant among the biodegradable metals in that they possess moderate mechanical properties and corrosion behaviors between Mg and Fe (Ref 1, 2). Zinc has a standard corrosion potential − 0.76 V, which is between Fe (− 0.44 V) and Mg (− 2.37 V). Therefore, the number of research work in Zn alloys has been rapidly increasing during recent years, which is regarded as the excellent candidate for the biodegradable for vascular stents. However, the strength and ductility of pure Zn are poor, which could not meet the clinical requirements. Element alloying is commonly chosen to tailor the properties of Zn. Similar as Zn and Ca in Mg alloys, Mg and Ca also could enhance the strength and ductility of Zn (Ref 3). Addition of minor 0.4 wt.% Mg in as-cast Zn improved the ultimate strength, yielding strength and elongation from 29.7, 27.5 MPa, and 0.62% to over 100, 90 MPa, and 30% (Ref 4). Zn alloyed with 1.0 wt.% Mg achieved the best combination of ultimate tensile strength and ductility (UTS ~ 250 MPa and ε ~ 12%) (Ref 5). Addition of Mg also is beneficial to biological behaviors of Zn alloys. Gong et al. reported that hot extrusion of as-cast Zn-1.0 wt.%Mg could improve the uniformity of biodegradation (corrosion rate down to 0.12 mm/year) and exhibit good biocompatibility (Ref 5). Li et al. obtained the enhanced compressive yielding strength from (102.92 ± 6.73) MPa for pure Zn up to 300-400 MPa for the alloying 1.0 wt.% Ca or 1.0 wt.% Sr after hot rolling and hot extrusion processes (Ref 6). Copper, as an essential trace element for human being, enhanced mechanical strength and elongation of Zn alloys together with antibacterial effects. The tensile YS and UTS and elongation values of Zn-4.0 wt.%Cu alloy are about 227 ± 5 MPa, 270 ± 0.5 MPa, and 50.6 ± 2.8%, respectively. This high elongation value could facilitate the micro-tubes processing for stent fabrication (Ref 7). Very high YS and UTS values from 213.7 to 426.7 MPa and from 257 to 440.5 MPa were achieved by combination alloying of Mg and Cu (Ref 8). Therefore, Mg, Ca, Sr, and Cu are favorable elements to improve the mechanical properties of Zn alloys.
Besides elemental alloying, hot working can also improve the mechanical properties and biodegradation behaviors because it can refine grains. Hot-rolled Zn-Mg alloys showed better ductility than their as-cast counterparts due to the inhibition of twinning by the refined Mg2Zn11 intermetallic phase and the finer grain size (Ref 4). The Zn-0.8 wt.%Mg alloys hot extruded in a hydraulic press at an extrusion ratio and rate of 10:1 and 2 mm/min owned the best combination of tensile mechanical properties (tensile yield strength of 203 MPa, ultimate tensile strength of 301 MPa and elongation of 15%), which will satisfy the requirement of the mechanical loading for the vascular stent if the elongation increases up to over 20% (Ref 9). 0.1wt.% Mn addition into as-rolled Zn-1.0 wt.%Mg enhanced the tensile strength and elongation up to 298 MPa and 26%, respectively (Ref 10). A significant enhancement of yield strength, tensile strength, and elongation was achieved from 65, 84 MPa, and 1.3% for as-cast Zn-3.0 wt.%Mg alloy to 205, 220 MPa, and 6.3% after 2-pass equal channel angular pressing (ECAP) technology, respectively. The grain size decreased from 48 μm for the as-cast alloy to 1.8 μm after homogenization and ECAP processes, leading to the reduced corrosion rate (Ref 11). A super-large elongation as much as 168% for Zn-15 wt.%Al alloys has been achieved through backward extrusion at 200 °C because of the refinement of Zn-rich phase (Ref 12). The yield strength, tensile strength, and elongation reached 200-300 MPa, 300-400 MPa, and 30% were recently reported for as-cast Zn-0.8 wt.%Mg alloy followed by subsequent extrusion and drawing processes (Ref 13). Therefore, elements alloying and hot working are most effective methods to improve the properties of Zn-based biodegradable alloys.
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It is important for the medical devices to possess antimicrobial properties, which is able to avoid the unnecessary infections to the patients. Metal ions Cu2+ and Ag+ have been conventionally used as antibacterial agents because of their excellent antimicrobial properties (Ref 14-16). Therefore, Cu or Ag alloying is one of the important ways to produce antibacterial metal materials. Research on Cu-containing antibacterial Ti alloys has shown strong antibacterial ability (Ref 17, 18). Both Cu and Ag are regarded as ideal alloying element since they are able to improve mechanical properties and keep good biocompatibility. Silver has been used for burns and the healing of wounds. Ag-containing materials have been successfully used for dental implants (Ref 19), and many studies have been performed based on its antibacterial properties. Silver ions or silver nanoparticles are able to kill bacteria and prevent them from adhering to the surface of implants (Ref 20-22). Besides enhancing the antibacterial rate of Zn alloy, Ag has also been proved to be able to enhance the tensile strength of Zn alloys. High yield strength and ultimate tensile strength of 236 MPa and 287 MPa were attained in Zn-7.0%Ag alloy. It is worth noting that Zn-7.0%Ag displayed superplasticity over a wide range of strain rates from 5 × 10−4 to 1.0 × 10−2 s−1. The elongations reached 325 and 418% at the strain rate of 5.0 × 10−4 s−1 (at 200 °C) for the Zn-5.0Ag and Zn-7.0Ag alloys, respectively (Ref 23).
Mg has super-strengthening effect to strengthen Zn alloys. A significant increase in tensile strength for Zn-0.02 wt.%Mg from 167 to 455 MPa was achieved by multi-drawing process (Ref 24). Benefit from the reduction in friction, allowing extrusion of larger billets, enhanced speed, and an increased ability to extrude smaller cross sections, indirectly extruded Zn-0.05 wt.%Mg alloy showed a good combination of high ultimate tensile strength up to 235 MPa and the elongation to fracture to 26%. It had a degradation rate that is about 0.15 mm/year and good biocompatibility and antibacterial behavior (Ref 25). The medical implants should possess good antibacterial performance along with load-bearing mechanical strength, biodegradation, and biocompatibility. The hybridization of Ag and Mg has seldom been reported in the published papers. In order to improve the antibacterial property of Zn-Mg alloy and investigate the strengthening effect of Ag on Zn-Mg alloy, Ag was incorporated into Zn-0.05Mg alloy, and then, the alloy was subjected to hot working treatment technique. Zn-Mg-Ag alloys were prepared by melt-blending of Ag with Zn-Mg master alloy followed by subsequent steel mold casting method. The alloys were subjected to homogenization treatment at 340 °C for 4 h prior to indirect extrusion at 200 °C with an extrusion ratio of 16:1. The microstructure, mechanical properties, corrosion, and antibacterial behaviors were investigated in as indirectly extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys, which can enrich the experimental database and provide practical guidance for the Zn-based biodegradable alloys.
Materials and Method
Materials
Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloy ingots were fabricated by melting commercial pure Zn (99.95 wt.%), pure Ag (99.99 wt.%), and Zn-50 wt.%Mg master alloy in a resistance furnace in air according to the designed compositions. Then, the melt was poured into a steel mold. The ingots were homogenized at 340 °C for 4 h under protection of graphite powders prior to extrusion. The indirect extrusion was carried out at 200 °C with an extrusion ratio of 16:1. The chemical composition of the extrudate was determined by inductively coupled plasma atomic emission spectrometer (ICP-AES, Optima 4300DV, PE, USA), and it showed that the extruded rod contained 0.056 wt.% for the top part and 0.055 wt.%Mg for the bottom parts, respectively, indicative of uniform distribution of Mg in the as-extruded rods.
Methods
Microstructure Characterization
The microstructures of the extruded Zn-0.05Mg and Zn-0.05Mg-(0.5, 1 wt.%) Ag alloy rods were characterized by optical microscopy (OM) (OLYMPUS GX-71) and scanning electron microscopy (SEM) (JEOL JSM-6510A). Samples for OM and SEM were firstly ground with SiC paper up to 2000 grits and then mechanically polished with 0.5 μm diamond pastes. Finally, these samples were etched in a mixture of 10% hydrochloric acid and 90% alcohol and washed immediately using alcohol for the microstructure observation on the microscope. The microstructure was analyzed on x-ray diffractometer (XRD) (Philips PW3040/60) with Cu-Kα radiation at scan rate of 3°/min and scan range from 30° to 100°.
Mechanical Testing
Dog bone-shaped tensile specimens of 5 mm in diameter and 25 mm in gauge length were prepared by machining the extruded rods. The tensile tests were performed at a constant cross-head speed of 1.5 mm/min at room temperature on SHIMADZU AG-X100 kN materials testing machine. Average and standard deviation of ultimate tensile strength (UTS), tensile yield strength (TYS), and elongation to break were determined by three tests for each group.
Electrochemical Corrosion Testing
Electrochemical potentiodynamic polarization tests were performed according to the previous publication (Ref 25), which were carried out on an electrochemical work station (Germany ZAHNER ENNIUM) with a scan rate of 1 mV/s in simulated body fluid (SBF) with pH 7.40 at 37 °C. The SBF (simulated body fluid) was standardized with the compositions of 8.035 g/L NaCl, 0.355 g/L NaHCO3, 0225 g/L KCl, 0.231 g/L K2HPO4·3H2O, 0.311 g/LMgCl2·6H2O, 39 mL/L HCl (1 mol/L), 0.292 g/L CaCl2, 0.072 g/LNa2SO4, 6.118 g/L TRIS, and 0-5 mL/L HCl (1 mol/L), according to ISO/FDIS 23317:2007 (E). Experiments were operated with a typical three-electrode electrochemical system: a sample as the working electrode (WE), a platinum sheet as the counter electrode (CE), and a saturated calomel electrode (SCE) as the reference electrode with a potential of 0.242 V/SHE. All the samples were ground by SiC paper up to 2000 grits and then mechanically polished with 0.5 μm diamond pastes, ultrasonically cleaned in alcohol, and dried by warm flowing air.
Immersion Testing
The disk-shaped samples with diameter of 11.5 mm in and thickness of 3 mm were chosen for immersion test. They were ground with SiC papers up to 2000 grits, cleaned in ethanol, and dried by warm flowing air prior to the immersion test. The weight of the sample was carefully recorded before and after the immersion tests. The immersion test was carried out in SBF with pH 7.40 at 37 °C, and the fluid was refreshed every 2 days. Samples were taken out from SBF after immersion for 1, 2, 4, 8, and 14 days. And then they were cleaned with 200 g/L of chromic acid (CrO3) solution in the ultrasonic cleaning machine to remove the corrosion products. The in vitro corrosion rate (CR) was calculated by Eq 1 according to the weight loss (Ref 25):
$${\text{CR}} = K\frac{\Delta W}{A \times T \times \rho }$$
(1)
where K stands for the time conversion coefficient (8.76 × 104), W is the weight loss after immersion (g), A is the sample area exposed to solution (cm2), T is the exposure time (h), and ρ is the density of the material (g/cm). Average and standard deviation were obtained after three measurements were taken from each group. The surface morphology after immersion tests was examined on a scanning electron microscope (JEOL JSM-6510A) equipped with an energy-dispersive x-ray spectroscopy (EDS).
Antibacterial Testing
Antibacterial test was measured by co-culture of the material with Staphylococcus aureus (S. aureus) or Escherichia coli (E. coli) (provided by the Microorganism Laboratory of the Chinese Medical University). Zn-0.05Mg, Zn-0.05Mg-0.5Ag, Zn-0.05Mg-1Ag samples with diameters of 10.0 mm and heights of 3.0 mm were ground with SiC papers up to 1200 grit, ultrasonic cleaned in acetone, absolute ethanol, and distilled water for 10 min each, followed by sterilization with ethylene oxide.
The sample was put in one well of 24-multi-well culture plate. 0.8 mL of bacterial suspension with a concentration of (5-10) × 105 cfu/mL, Escherichia coli ATCC25922 (E. coli) or Staphylococcus aureus ATCC25923 (S. aureus) was dropped into each well and then incubated at 37 °C for 12 h and 24 h, respectively. Then the co-cultured bacterial suspensions were diluted 40-fold with PBS. 0.05 mL of this diluted bacterial suspension was spread evenly onto the nutrition agar plate and incubated at 37 °C for another 24 h for bacterial colonies count. The antibacterial rate (AR) was determined by the following equation (Ref 25):
$${\text{AR}} = \frac{N1 - N2}{N1}$$
(2)
where N1 and N2 are the bacterial colony numbers for the control alloy (Ti-6Al-4 V) and the test materials (Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag), respectively.
Results and Discussion
Morphology
The cross- and longitudinal section microstructures of indirectly extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys are shown in Fig. 1. The crystals are mainly equiaxed, indicating that the dynamic recrystallization takes place during the extrusion process. Some twins can be observed in both cross-sectional and longitudinal images of the samples. The crystals usually slide along the {0001} basal plane when the Zn-based alloys are extruded. Then the slides are retarded when the deformation accumulates to some extent, leading to the formation of the twins. At the same time, the twins change the orientation of the crystals and facilitate the further sliding process. A small amount of new phase with a different contrast in the image appears in Zn-0.05Mg alloy, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag. The grain size along cross section increases from 23 μm for Zn-0.05Mg to 42 μm for Zn-0.05Mg-0.5Ag, and 27 μm for Zn-0.05Mg-1Ag alloy. The grain size along longitude direction shows similar trend. The longitude grain size of Zn-0.05Mg is measured to be 23 μm, increasing up to 40 μm and 30 μm for the alloy containing 0.5 and 1 wt.% Ag. It is worth to mention that the crystal sizes for Zn-0.05Mg-0.5Ag are larger than those of Zn-0.05Mg alloy, which is different from the reported Ag-added Zn alloys (Ref 26). The reason for this abnormal phenomenon has not been yet understood. Ogura et al. reported Ag addition in Al-Zn-Mg alloy did not refine the grains because Ag atoms preferentially trapped vacancies migrating not only grain boundaries but also dispersoids and then form fine precipitates in grain interiors and in the vicinity of grain boundaries (Ref 27). However, the grain in Zn-0.05Mg-1Ag is significantly refined, which is much smaller than that in An-0.05Mg-0.5Ag alloy. The average grain size in Mg-Zn alloy decreases to about 5 μm with 1% Ag, and to about 2 μm with 2% Ag additions (Ref 28).
Fig. 1
Microstructures of as-extruded Zn-Mg alloys: (a) and (b) the cross section and longitudinal section of Zn-0.05Mg; (c) and (d) the cross section and longitudinal section of Zn-0.05Mg-0.5Ag; (e) and (f) the cross section and longitudinal section of Zn-0.05Mg-1Ag alloy
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Microstructure
The phase constituent of the alloys was investigated by x-ray diffraction testing. As shown in Fig. 2, Mg2Zn11 phase appears in both Zn-0.05Mg and Zn-0.05Mg-(0.5, 1) Ag alloys. Besides, the existence of AgZn3 phase is not able to be distinguished due to its peak overlapping with that of Mg2Zn11. Mostaed et al. did not observe the second phase in as-extruded Zn-2.5Ag alloy (Ref 23) due to the high solubility of Ag in Zn according to Zn-Ag binary phase diagram (Ref 29). Combining XRD patterns and morphologies of the alloys, we believe that there is no detectable AgZn3 phase in Zn-0.05Mg -0.5Ag and Zn-0.05Mg-1Ag alloys.
Fig. 2
XRD patterns of as-extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys
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Mechanical Properties
The mechanical properties of as-extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys are evaluated by stress–strain curves. As shown in Fig. 3, the ultimate tensile strength (UTS) of Zn-0.05Mg-0.5Ag does not show any enhancement accompanying with a decrease in tensile strain. When Ag content is 1 wt.%, the yielding strength and ultimate strength increase up to 170 and 202 MPa, higher than the values of 140 and 182 MPa for Zn-0.05Mg alloy, and 160 and 183 MPa for Zn-0.05Mg-0.5Ag alloy, respectively. However, the tensile strain decreases from 9.9% for Zn-0.05Mg and 7.7% for Zn-0.05Mg-0.5Ag, and 6.9% for Zn-0.05Mg-1Ag alloys at strain rate of 10−3 s−1. The strengthening effects on mechanical properties of alloys are usually achieved by minor addition of alloying elements via solid solution and/or grain refinement. Liu et al. got an enhanced tensile strength from 151 to 209 MPa and 241 MPa by addition of 0.1 wt.% Sr and Ca into Zn-1.5Mg alloy (Ref 2). Wang et al. obtained a hot-worked Zn-5.0Mg-1.0Fe (wt%) alloy micro-tube with a high strength of more than 220 MPa and tensile strain of over 20% (Ref 30). Ag is mainly solid solution in Zn matrix without forming discernible compounds. According to the morphology of Zn-0.05Mg-0.5Ag, no obvious grain refining is seen. Therefore, the mechanical strength is not significantly increased despite Ag addition into Zn-0.05Mg alloy. When Ag content is 1 wt.%, the average grain size of Zn-0.05Mg-1Ag alloy is much smaller than that of Zn-0.05Mg alloy, leading to the increase in the mechanical strength. It can be seen that the grain size of Zn-0.05Mg-1Ag alloy is almost same as that of Zn-0.05Mg alloy. The enhanced mechanical strength of Zn-0.05Mg alloy at high Ag content is mainly contributed to the solid solution strengthening effect of Ag in Zn matrix. Yue et al. reported an adverse effect of Fe addition on the mechanical properties in Zn-3Cu alloy due to the dynamic recrystallization and the formation of brittle phase of FeZn13 (Ref 31). As an alloying element, the strengthening effect of Ag is not strong as other elements such as Mg, Cu, and Fe. According to the published and our present results, it is necessary to optimizing the content of Mg and Ag to achieve the highly improved mechanical properties.
Fig. 3
Tensile stress vs. strain curves (a) and the UTS, TYS, and elongation of as-extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys
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Corrosion Properties
Figure 4 shows potentiodynamic polarization curves of the indirectly extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys in simulated body fluid at 37 °C, and the curve of pure Zn is shown for comparison. As seen from Fig. 4, corrosion potential (Ec) of Zn-based alloys containing Mg and Ag shifts to the negative side. It is indicative of a more cathodic potential of Mg2Zn11 (Ref 32). The corrosion potential (Ec), pit potential (Ep), and current density (Ic) of the alloys are tabulated in Table 1.
Fig. 4
Potentiodynamic polarization curves of the indirectly extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys in simulated body fluid at 37 °C: The curves of pure Zn are shown for comparison (Ref 25)
Table 1
Corrosion parameters of as-extruded pure Zn, Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys
Alloys
Ec, V
Ep, V
jc, µA/cm2
Zn
− 0.85
...
31.6
Zn-0.05Mg
− 0.90
− 0.79
10.0
Zn-0.05Mg-0.5Ag
− 0.93
− 0.80
2.0
Zn-0.05Mg-1.0Ag
− 0.93
− 0.79
7.1
×
Table 1 shows that the Ec of pure Zn-0.05Mg moves from − 0.90 toward − 0.93 V with an addition of 0.5 wt.% Ag. The corrosion potential is steady at − 0.93 V when the content of Ag reaches 1 wt.%. The alloy with more negative Ec indicates its tendency to experience more significant corrosion damage. Pitting corrosion tends to take place on passive alloys in the presence of halogen species such as chloride, bromide, and iodide ions and hydrogen fluoride. These halogen species attack the defects of passive films and locally break the film. The sensitivity to pitting corrosion of the metal can be estimated based on the pitting potential obtained by electrochemical measurements (Ref 33). A slight variation of Ep first increasing from − 0.79 to − 0.80 V and then decreasing down to − 0.79 V happens for the Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys, respectively. There is no much difference in Ep for these alloys, demonstrating that they have close pitting corrosion sensitivity (Table 2).
Table 2
EDS results of as-extruded Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys
Different from pure Zn, Zn-Mg and Zn-Mg-Ag alloys show breakdown potentials in the anodic polarization curves, indicating the formation of protective films on the alloy samples. Zn-Mg and Zn-Mg-Ag alloys exhibit obvious passivation region (PR). The passivation currents show a decrease with an addition of Ag, indicating the stronger ability to resist corrosion in the specific environment. Zn ions will be firstly released into the solution and form Zn hydroxide When Zn alloy is in contact with SBF solution. ZnO is then formed via dehydration of Zn hydroxide, additionally making contribution to the antibacterial property of materials. As for Mg, it can be dissolved and form corrosion products in the form of MgCO3 (magnesite) (Ref 34).
The corrosive current density shows apparent decrease with the addition of 0.5 wt.% Ag and then increase with Ag content increasing up to 1 wt.%. Compared with the value 10.0 μA/cm2 for Zn-0.05Mg alloy, the corrosion current density reduces by 78.0 and 19.0% by adding 0.5 and 1 wt.% Ag, respectively, indicative of the improvement of anticorrosion property of Ag-containing alloys. Khripta et al. have demonstrated that the corrosion behaviors are affected by the surface roughness, grains, microstructures, and second phases in the alloys via severe plastic deformation treatment (Ref 35, 36). The increase in corrosion current density for Zn-0.05Mg-1Ag is probably due to the local corrosion at the grain boundaries due its small grain sizes. Sikora-Jasinska et al. concluded that the Zn-Ag alloys with high Ag content showed a tendency to more localized corrosion than the solid solution at lower Ag content (Ref 23). Pit potential could show the tendency to pit corrosion of metals. However, it is not able to predict that the pit corrosion takes place in the practical application. By combining the results of corrosion potential and pit potential, the information about pit corrosion trend can be obtained by using corrosion potential as a reference. Generally, the easier the metal is pit corroded, the smaller difference between corrosion potential and pit potential. As a result, Zn-0.05Mg-0.5Ag alloy is the most difficult to be pit corroded among the three alloys.
Weight Loss After Immersion Test
The dependence of corrosion rate on the immersion time in SBF at 37 °C for the indirectly extruded alloys is shown in Fig. 5. The corrosion rates of all the alloys are not steady at the beginning of immersion and then get stabilized at the value of 0.15-0.2 mm/year or 2.5-3.5 g/m2/day after 4 days’ later. As for the alloys immersed for 1 day, the corrosion rate of the alloys shows a big difference, with the highest rate of 4.2 g/m2/day or 0.25 mm/year for Zn-0.05Mg-1Ag alloy and the lowest rate of 1.0 g/m2/day or 0.1 mm/year for Zn-0.05Mg alloy. The corrosion rates of the tested alloys are very close to the value of 3.0 g/m2/day or 0.18/year. There is only small difference (0.05 mm/or 1.0 g/m2/day) in the corrosion rate for these alloys until immersion of 14 days. The corrosion rate of Zn-0.05Mg and Zn-0.05Mg-0.5Ag alloy immersed for 14 d is almost reached the same value of 3 g/m2/day or 0.15 mm/year. This corrosion rate is close to the as-rolled Zn-1.0Mg alloy in report publication (Ref 37). However, the corrosion rate of Zn-0.05Mg-1Ag is a little higher than throughout the whole immersion process than those of Zn-0.05Mg and Zn-0.05Mg-0.5Ag alloys, indicating that the high Ag content will accelerate the corrosion of Zn-0.05Mg alloy.
Fig. 5
Correlation between the corrosion rate of the as-extruded pure Zn and Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys and the immersion time: The curves of Zn are shown for comparison (Ref 25)
×
Surface morphology of indirectly extruded Zn-Mg-based alloy samples after immersion for 14 days in SBF solution at 37 °C is shown in Fig. 6. No apparent pores are presented on the corroded surface of the alloys, indicative of uniform corrosion process. Some bright aggregated corrosion product (marked by the arrows) can be observed. Obvious interconnected cracks shown in Fig. 6(a) and (c) and round aggregates with size of a few micrometers appear on the surface of Zn-0.05Mg alloy. Some obvious large aggregates are seen for the Zn-0.05Mg-0.5Ag even though the aggregates turn small, while the cracks are hardly seen on the surface of the alloys. It might be attributed to the coarse grain for Zn-0.05Mg-0.5Ag alloy, larger than those for Zn-0.05Mg and Zn-0.05Mg-1Ag alloys, as measured from the morphology observation. The smallest aggregates are uniformly distributed on the corroded surface of Zn-0.05Mg-1Ag alloy. However, some thin cracks take place, meaning its poor corrosion resistance. The addition of Ag facilitated to form a more dense and protective film, prohibiting the formation and growth of corrosive products. According to EDS results, the corrosion products are determined to be (Zn, Mg, Ca)3 (PO4)2 or (Zn, Mg, Ca)CO3, which are introduced from the corroded alloy and SBF.
Fig. 6
Surface morphology of indirectly extruded Zn-Mg-based alloy samples after immersion for 14 days in SBF solution at 37 °C: (a) Zn-0.05Mg; (b) Zn-0.05Mg-0.5Ag; (c) Zn-0.05Mg-1Ag
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Antibacterial Tests
Postoperative infection is a critical issue in surgical implants. Thus, the implants are required to be antibacterial, which can prevent the infection. Silver has been used in burns and wound healing due to its high antibacterial activity in a lot of chemical states. Figure 7 illustrates the bacterial colonies after they are incubated with Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys. No bacterial can be seen in all three samples, indicating their excellent antibacterial ability. Zn is easy to be oxidized and its oxide (ZnO) showed good antibacterial capability (Ref 38). Petrini et al. found that the coexistence of ZnO with TiO2 reduced the viability of five streptococcal oral strains on titanium oxide surfaces modified with zinc (Ref 39).
Fig. 7
Morphologies of E. coli (above) and S. aureus (below) cultures with controlled alloy (a), Zn-0.05Mg(b), Zn-0.05Mg-0.5Ag(c), Zn-0.05Mg-1Ag (d) at 37 °C for 24 h: Controlled alloy (a) and Zn-0.05Mg(b) are shown for comparison (Ref 25)
×
Silver (Ag) has been widely used to improve the antibacterial ability of implants. Ag ion released from the alloy can interact with bacterial cell wall membranes, destroy the wall membrane, and eventually kill the bacteria (Ref 40). The antibacterial property of Zn-Mg alloy is improved for both Staphylococcus aureus and Escherichia coli by addition of Ag. As seen from Fig. 8, the antibacterial rate in Staphylococcus aureus culture increases from 98.01% for Zn-0.05Mg to 98.68% for Zn-0.05Mg-0.5Ag, 99.34% for Zn-0.05Mg-1Ag, respectively. The Escherichia coli antibacterial rate also reaches 98.26 and 99.42% when Ag content increases from 0.5 to 1 wt.%, about two percent higher than that of the alloy free of Ag. The antibacterial ability usually increases with the increase in Ag content (Ref 41, 42). Xie et al. observed the 100% antibacterial rate until the concentration of Ag reached 3.5 wt.% in Zn-Ag scaffold (Ref 43). Therefore, it is effective for Ag to enhance the antibacterial rate of the alloys.
Fig. 8
Antibacterial rate of control, Zn-0.05Mg, Zn-0.05Mg-0.5Ag, and Zn-0.05Mg-1Ag alloys after co-cultured 24 h with E. coli and S. aureus
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Conclusions
The biodegradable Zn-0.05Mg alloy with 0.5 and 1 wt.% Ag was manufactured by indirectly extruding technology. There was no new Ag-related phase present except the Mg2Zn11 phase in the as-extruded alloys. The yielding strength and ultimate strength of Zn-0.05Mg alloy were enhanced by 21.4 and 11.0% when 1.0% Ag was added while the elongation to break exhibited a decrement by adding Ag. The anticorrosion performance was improved by addition of Ag. The Zn-0.05Mg-0.5Ag showed the lowest corrosion density of 2.2 μA/cm2 and low corrosion rate of 0.15 mm/year. The addition of Ag improved the antibacterial ability of Zn-0.05Mg alloy, in which the antibacterial rate in Staphylococcus aureus culture and Escherichia coli culture reached over 99%. The Zn-Mg-Ag alloys with reasonable mechanical strength and good antibacterial performance are promising for biodegradable medical devices.
Acknowledgments
This project was supported by the National Natural Science Foundation of China (81672139), Postdoctoral Science Foundation of China (No.: 194016), Doctoral Research Starting Foundation of Affiliated Zhongshan Hospital of Dalian University (No.: DLDXZSYY-BK201703), and Doctoral Research Starting Foundation of Dalian University (No.: 20152QL002).
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