Research on biodegradable Mg-Zn-Gd alloys for potential orthopedic

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Research on biodegradable Mg-Zn-Gd alloys for potential orthopedic implants: In vitro and in vivo evaluations Hongwei Miao, Dandan Zhang, Chenxin Chen, Lei Zhang, Jia Pei, Yun Su, Hua Huang, Zhongchang Wang, Bin Kang, Wen-Jiang Ding, Hui Zeng, and Guangyin Yuan ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.8b01563 • Publication Date (Web): 22 Jan 2019 Downloaded from http://pubs.acs.org on January 24, 2019

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ACS Biomaterials Science & Engineering

Research on biodegradable Mg-Zn-Gd alloys for potential orthopedic implants: In vitro and in vivo evaluations Hongwei Miaoa, c#, Dandan Zhangb,#, Chenxin Chena, Lei Zhanga, Jia Peia, Yun Sub, Hua Huanga*, Zhongchang Wangc, Bin Kangd, Wenjiang Dinga, Hui Zengd, Guangyin Yuana*. a

National Engineering Research Center of Light Alloy Net Forming and Key State Laboratory of

Metal Matrix Composites, Shanghai Jiao Tong University, School of Materials Science and Engineering, Shanghai, 200240, China; bDepartment

of Ophthalmology, Shanghai Ninth People’s Hospital, Shanghai Jiao Tong

University, School of Medicine, 200011, Shanghai, China. cInternational dDepartment

Iberian Nanotechnology Laboratory, Braga, 4715-330, Portugal.

of Orthopedics, Peking University Shenzhen Hospital, Shenzhen, 518036, China

* Corresponding

author: TEL: +86-21-34203051; FAX: +86-21-34202794; E-mails:

[email protected] (H. Huang); [email protected] (G. Yuan) #

Hongwei Miao and Dandan Zhang contributed equally to this work.

Abstract Various kinds of biodegradable Mg alloys have been developed in recent years due to their appropriate mechanical properties, biodegradation and good biocompatibility. In this study, Mg-2.0Zn-xGd alloys (x= 0.5, 1.0, 1.5 and 2.0 wt.%) were prepared. Hot extrusion was applied in order to refine the microstructure and improve the degradation resistance. The microstructure, mechanical properties and in vitro degradation behavior of Mg-2.0Zn-xGd alloys were 1

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investigated firstly. The as-extruded Mg-2.0Zn-1.0Gd alloy exhibits excellent mechanical properties (UTS: 338 MPa, YS: 284 MPa, elongation: 24 %) and low in vitro degradation rate (0.24 mm/year) with uniform degradation morphology, and then this alloy was selected for further assessments. The cytotoxicity of as-extruded Mg-2.0Zn-1.0Gd alloy to MC3T3-E1 cell is found to be grade 0~1, indicating a good biocompatibility. The in vivo experiment shows that the in vivo degradation rate of this alloy is about 0.31 mm/year after 30 days implantation in cranial defect of Sprague-Dawley rats. All of these indicate a promising prospect of Mg-2.0Zn-1.0Gd alloy as biodegradable applications, especially as orthopedic implants. Keywords: Biodegradable; Magnesium alloy; Mechanical properties; Degradation behavior; Cytotoxicity.

1. Introduction Compared with the most extensively used permanent orthopedic implant materials such as stainless steels and titanium alloys, Mg alloys exhibit many outstanding advantages and have been researched extensively as biodegradable materials in recent years

1-2.

Firstly, Mg alloys

possess suitable mechanical properties and their elastic modulus (40-45 GPa) is similar to that of human bone (10-30 GPa), which can eliminate the stress shielding effect. Secondly, Mg alloys are biodegradable and this can avoid a second surgery for implant removal. In addition, there have been a lot of studies which confirm that Mg alloys have good biocompatibility and some positive clinical effects. For instance, Zhang et al.

3

has reported that Mg can promote

CGRP-mediated osteogenic differentiation, which shows the therapeutic potential of this ion in 2


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orthopedics. Lee at al. 4 has reported a long term clinical study of Mg-Ca-Zn alloy in vivo and the results shows that Mg can facilitate early bone healing and exhibit potential for further application in orthopedic implants. The orthopedic applications require to maintain mechanical integrity over 12-18 weeks during the healing of bone tissue 1, which means the Mg alloys used for orthopedic applications must possess good biocompatibility, excellent mechanical properties and outstanding corrosion properties simultaneously. However, the mechanical properties of pure Mg are not enough for many biomedical applications 5, the traditional commercial Mg alloys, such as AZ31 and AZ91, contain Al element which is harmful to neurons and associated with dementia and Alzheimer’s disease. To meet the requirements of clinical applications, lots of Mg alloys, such as Mg-Zn and Mg-RE

6

2

alloys, and so on, have been designed. These studies confirmed that Zn and

Rare-earth (RE) elements are suitable for alloying of biomedical Mg alloys. Zn is essential for human body and the recommended daily intake for adults is 8-11 mg/day 1, which is considered non-toxic below the appropriate amount. Although some of the RE elements are toxic, such as Pr and Ce, others like Gd and Nd are acceptable at low concentrations 6. Recently, the MAGNEZIX screw (made by MgYREZr, which is similar to WE43 and has Gd addition) demonstrates good biocompatibility and osteoconductive properties in hallux valgus surgery, and has obtained the CE mark for medical devices 1. Hence, Mg-Zn-Gd alloy system is expected to have good biocompatibility. Actually, the Mg-Zn-Gd-based alloys have been developed and investigated extensively as structural materials for industry applications 7. Huang et al. has reported that the nano-scale I-phase which can precipitate during extrusion is beneficial to both strength and 3



elongation

8-10.

Good mechanical properties and weakened basal texture after deformation are

beneficial for the fabrication and application of medical implants. Therefore, a new type of biomedical Mg-Zn-Gd alloy could be obtained by further optimizing the Zn and Gd contents to tailor the degradation properties. Lots of studies confirmed that a little amount of Zn addition could improve both the tensile strength and corrosion resistance of Mg alloys. Usually, the Zn content for ternary biomedical Mg alloys is around 2.0 wt.%

11-12,

superabundant Zn usually leads to serious localized

corrosion13. RE elements are also proved to have a beneficial effect on the corrosion resistance and mechanical properties of Mg alloys 14. Therefore, in this study, Mg-2.0Zn-xGd (x= 0.5, 1.0, 1.5 and 2.0 wt.%) alloys with different Gd addition were designed in order to optimize a new kind of biodegradable Mg alloy for orthopedic implants. The microstructure, mechanical properties, and cytotoxicity of the as-extruded alloys were characterized, and the degradation behaviors were investigated both in vitro and in vivo.

2. Experimental section 2.1 Materials preparation Alloys were produced by melting high purity Mg (99.99%), high purity Zn (99.995%), and Mg-30 wt.% Gd master alloy. The raw materials were melted in a resistance furnace under the protection of mixed atmosphere of SF6 (1 vol.%) and CO2 (99 vol.%). After melting pure Mg, Mg-Gd master alloy was added into the melt at 760 °C. Zn was then added into the melt at 4


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700 °C. The melt was held at 730 °C for 15 min and then poured into a steel mould which had been preheated to 200 °C. The chemical compositions of the alloys were determined by inductively coupled plasma (ICP-AES, iCAP6300, USA) analyzer and the results were listed in Table.1. The ingots used for extrusion were machined into billets with a dimension of 60 mm in diameter and 60 mm in height, and then solid solution treated at 470 °C for 16 hours. All the billets were preheated at 250 °C for 1 hour and then extruded at this temperature with an extrusion ratio of 9:1. And the extrusion speed was about 2 mm/s. As a counterpart, pure Mg (99.99 %) was also extruded under the same conditions. Table 1 Chemical composition of the alloys Analyzed composition (wt.%) Alloy Zn

Gd

Mg

Mg-2.0Zn-0.5Gd

2.21

0.51

Bal.

Mg-2.0Zn-1.0Gd

2.33

0.84

Bal.

Mg-2.0Zn-1.5Gd

2.15

1.71

Bal.

Mg-2.0Zn-2.0Gd

2.17

2.12

Bal.

2.2 Microstructure characterization and tensile test The X-ray diffractometer (XRD, PW3040/60) with Cu-Kα radiation at 40 KV and 100 mA was used for phase analysis. The step was fixed to 10 °/min and the measurement angle ranged from 20° to 80°.The optical microstructure (OM) was characterized by microscopy (Leica 5



MEF4M). For OM observations, the as-extruded specimens were polished and then etched with an etchant of 1 g oxalic acid, 1 ml nitric acid, 1 ml acetic acid and 150 ml distilled water. The volume fraction of the secondary phase was measured using imageJ software and the linear intercept method was applied to measure the average grain size for the as-extruded samples. The grain orientation of the as-extruded specimens was revealed by electron backscatter diffraction (EBSD) using a scanning electron microscope (SEM, FEI NOVA), and the test surface is parallel to the direction of the extrusion. Tensile test was carried out on a Zwick/Roell Z100 testing machine at room temperature. The dog-bone specimens (15mm in gauge length) for tensile test were cut along the extrusion direction by electric spark machine. At least three samples were tested for each alloy and all the tests were carried out at an initial strain rate of 1×10-3 /s. 2.3 In vitro degradation tests In order to evaluate the in vitro degradation properties, immersion test was conducted at 37± 0.5 °C for 10 days in Hank’s solution, which was composed of 8.0 g/L NaCl, 1 g/L glucose, 0.4 g/L KCl, 0.14 g/L CaCl2, 0.35 g/L NaHCO3, 0.1 g/L MgCl2·6H2O, 0.1208 g/L Na2HPO4·12H2O, 0.06 g/L KH2PO4, 0.06 g/L MgSO4·7H2O. The samples for immersion test were cut from the as-extruded rods with a diameter of 15 mm and a height of 3 mm, then grinded to 3000 grit paper. The ratio of the solution volume to the specimen surface area is 40 ml/cm2. The immersion solution was renewed every 24 hours in order to keep a relative stable pH value. The samples used for observation of corrosion morphology were sputtered with Au in order to 6


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improve the conductivity. After observing the degradation morphology with corrosion products by SEM and EDS, the samples were immersed in a chromic acid consisting of 200 g/L CrO3 to remove the degradation product and rinsed with distilled water, ethanol, finally dried in warm flowing air. Then the degradation morphology without corrosion product was observed by optical microscope, SEM and a 3D-laser profilometry. The degradation rate (R) was calculated by the following formula according to ASTM-G31-72: R=

𝐾𝑊

(1)

𝐴𝑇𝐷

where K is a constant, W is the mass loss in g, A is the area in cm2, T is the time of exposure in hours, and D is the density in g/cm3. To eliminate the error, at least three samples were tested for each condition. 2.4 Cytotoxicity assessments Osteoblastic cells MC3T3-E1 (Cell Bank, Chinese Academy of Sciences) were cultured in alpha-modified Eagle’s medium (α-MEM, Gibco, USA), supplemented with 10% fetal bovine serum (FBS, Gibco, USA) and 1 % penicillin &streptomycin (Gibco, USA) at 37 °C with 5 % CO2, with fresh medium replaced every 2 days. Cytotoxicity was determined by indirect contact method. The extracts were prepared by using growth medium consisting of α-MEM supplemented with 10 % fetal bovine serum and 1 % penicillin &streptomycin, with the sample surface area to solution media ratio of 1.25 cm2/ml according to ISO 10993-5. After 72 hours incubation at 37 °C with 5 % CO2, the extraction media was diluted to 50 % and 10 % concentration. The pH value and osmolality of these extracts were measured with a pH meter 7



(FE20, Mettler Toledo, Switzerland) and a freezing point osmometer (Osmolality 3000, Gonotec). Cells were seeded in a 96 well plate at 2×104 cells/ml, 100 μl for each well, and incubated for 24 hours to allow attachment. Then the medium was replaced by 100 μl extraction media. After 1, 3 and 5 days of incubation, the extracts were replaced by 100 μl Cell Counting Kit-8 (CCK8, Beyotime Biotech, China) solution (a ratio of 1/10 in α-MEM). The optical density (OD) measurements were measured at 450 nm with a microplate reader (iMARK, Bio-Rad, USA). The morphology of the cells after incubation in extraction media for 1 and 3 days was observed by cell staining. The extraction media were removed from the incubator and gently washed with Dulbecco phosphate buffer saline (D-PBS, HyClone, USA) which was pre-warmed to 37 °C. Cells were fixed in 4 % paraformaldehyde solution for 15 min and gently rinsed with D-PBS. Then the adhering cells were permeabilized in 0.5 % TritonX-100 in D-PBS and gently rinsed with D-PBS. Cells were subsequently stained with Alexa Fluor 488 phalloidin for 30 min, and after rinsed twice with D-PBS, the nuclei were stained with DAPI for 10 min and rinsed twice with PBS. 2.5 In vivo degradation tests All procedures were performed in accordance with the NIH guidelines for the care and use of laboratory animals (NIH Publication No. 85e23 Rev. 1985) and approved by the Animal Experimental Ethic Committee of Ninth People’s Hospital affiliated to Shanghai Jiao Tong University School of Medicine. Twelve Sprague-Dawley rats (SD rats, female, 6-8 weeks old) 8


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were selected and the surgery was carried out under sterile conditions, all the animals were anesthetized using pentobarbital sodium (3.5 mg/100ml) through intraperitoneal injection. Then a sagittal scalp incision was made, followed by the 8 mm critical-sized cranial defect created by a trephine, the as-extruded Mg-2.0Zn-1.0Gd alloy and as-extruded pure Mg plates with a diameter of 8 mm and a thickness of 1 mm were implanted into the defect in order to evaluate the in vivo degradation behavior. The samples were harvested at 10 days and 30 days post implantation and scanned by a micro-computed tomographic imaging system (μCT, GE Explore Locus SP micro-CT, USA). Three-dimensional images were reconstructed to evaluate the in vivo corrosion behavior. The surface of samples was observed by SEM (Phenom XL) and the corrosion products were studied by EDS. After being embedded in resin and grinded on grit paper, the longitudinal section with corrosion products and Mg matrix was observed by SEM and the elements distribution was studied by EDS line scanning.

3. Results 3.1 Microstructure characterization and tensile test Fig.1 shows the XRD results of Mg-2.0Zn-xGd alloys. Peak analysis reveals that there are mainly α-Mg, I-phase, W-phase and Mg2Zn3 phase in these alloys, which is correspond to the previous reports about Mg-Zn-Gd alloys 9, 15-16.

9



Fig. 1 XRD results of Mg-2.0Zn-xGd alloys

Fig. 2 (a) shows the optical micrographs of the as-extruded Mg-2.0Zn-xGd alloys along the extrusion direction. In this study, the alloys were hot extruded, which means the dynamic recrystallization and the growth of grain would happen during the extrusion process. When the content of Gd is below 1.0 %, due to the lack of secondary phases, the recrystallization was not fully occurred, so the grain size of Mg-2.0Zn-0.5Gd is larger than that of Mg-2.0Zn-1.0Gd. With the increasing of Gd content, the non-recrystallization region decreases and there can hardly find non-recrystallization region when the Gd content is over 1.0 wt.%. At the same time, the alloys with higher Gd content possess much faster speed of dynamic recrystallization, which means there were much more time for the growth of grain. As a result, the average grain size decreases from about 6.7 μm to about 4.9 μm firstly and then increases to about 7.2 μm (Fig 2 (b)). The volume fraction of the secondary phases increases steadily, as shown in Fig. 2 (c), and a typical black band structure is formed when the Gd content is higher than 1.0 wt.%, which should be avoided due to its adverse effect on mechanical properties 8, 17.

10


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Fig. 2 Optical microstructure images and analyzed results including average grain size and secondary phase of the as-extruded Mg-2.0Zn-xGd alloys: (a) optical micrographs; (b) average grain size; (c) volume fraction of secondary phases.

Fig.3 shows the EBSD results of the as-extruded Mg-2.0Zn-xGd alloys. The pole figure shows that the texture is weakened gradually with the addition of Gd. As a comparison, the texture intensity is 27.81 for Mg-2.0Zn-0.5Gd alloy, while it is 12.73, 8.64 and 5.79 for the 1.0Gd, 1.5Gd and 2.0Gd containing alloy, respectively. 11



Fig. 3 EBSD results of the as-extruded Mg-2.0Zn-xGd alloys: (a) Mg-2.0Zn-0.5Gd；(b) Mg-2.0Zn-1.0Gd; (c) Mg-2.0Zn-1.5Gd; (d) Mg-2.0Zn-2.0Gd. (RD=extrusion direction, TD=radius direction)

The mechanical properties of the as-extruded Mg-2.0Zn-xGd alloys are shown in Fig. 4. The yield strength (YS), ultimate strength (UTS) and elongation increase firstly and then decrease steadily with increasing the Gd addition. The as-extruded Mg-2.0Zn-1.0Gd alloy shows the best mechanical properties, and the UTS, YS and elongation reached 338 MPa, 284 MPa and 24%, respectively, which may be related to the refined and uniform microstructure without obvious black band structure after extrusion.

Fig. 4 Mechanical properties of the as-extruded Mg-2.0Zn-xGd alloys

3.2 In vitro degradation tests The calculated degradation rate of the as-extruded Mg-Zn-Gd alloys according to the 10 12


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days’ mass loss test is presented in Fig. 5. The Mg-2.0Zn-0.5Gd alloy shows the slowest degradation rate, i.e. 0.15 mm/year, then the degradation rate increases steadily with further addition of Gd and it reaches 0.24 mm/year, 0.42 mm/year and 1.19 mm/year for the 1.0Gd, 1.5Gd and 2.0Gd containing alloy, respectively.

Fig. 5 In vitro degradation rate of the as-extruded Mg-2.0Zn-xGd alloys

Fig. 6 shows the SEM images of sample surfaces after immersion in Hank’s solution for 10 days. The cracks on the surfaces are results of the shrinkage of degradation products during drying. As a result of the highest degradation rate, there are plenty of porous degradation products on the surface of Mg-2.0Zn-2.0Gd alloy. According to the EDS results, the degradation products consists of C, O, Mg, Ca and P, which implies the existence of hydroxides and hydroxyapatite

18.

There is no significant difference among the element composition of

degradation products for all alloys, which means the degradation products of these alloys are similar to each other.

13



Fig. 6 Surfaces observation and EDS results of degradation products for the as-extruded Mg-2.0Zn-xGd alloys after immersion in Hank’s solution for 10 days

In general, uniform degradation behavior is especially important for degradable Mg alloys as biomaterials. The degradation morphologies after removing the degradation products are shown in Fig. 7. There are few serious corrosion pits on the surface of Mg-2.0Zn-0.5Gd and Mg-2.0Zn-1.0Gd samples, while the larger corrosion pits can be easily found when the addition 14


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of Gd is over 1.0 wt.%, as shown in Fig. 7 (a). Besides the large corrosion pits, other areas, which are marked by frame lines, are also observed by SEM and 3D-laser profilometry. It can be seen that there are plenty of uniformly distributed tiny corrosion pits on the surface of Mg-2.0Zn-0.5Gd and Mg-2.0Zn-1.0Gd alloys, and the depth of corrosion pits is very shallow and uniform, indicating uniform degradation mode. As contrast, there are lots of visible larger corrosion pits on the surface of the Mg-2.0Zn-1.5Gd alloy and Mg-2.0Zn-2.0Gd alloy, which can lead to serious localized corrosion, and the degradation is very deep and uneven.

15



Fig. 7 Surfaces observation after removing degradation products: (a) SEM observation; (b) local observation by 3D-laser profilometry

3.3 Cytotoxicity assessments Due to the excellent mechanical properties and good degradation resistance, the as-extruded Mg-2.0Zn-1.0Gd alloy (Hereafter, donated as ZG21 in figures) was chosen for further cytotoxicity assessment. The as-extruded pure Mg under the same production process was chosen as a control group. Fig. 8 (a) shows the cell viabilities of MC3T3-E1 cells after 1, 3 and 5 days incubation in Mg-2.0Zn-1.0Gd alloy and pure Mg extracts. There are no obvious differences between different culture periods. Both 100% extracts were cytotoxic to MC3T3-E1 cells, and the cell viability increased significantly after dilution. The cytotoxicity of Mg-2.0Zn-1.0Gd alloy and pure Mg are both evaluated to be grade 0-1 under the condition of 50 % and 10 % extracts, meeting the clinical requirement 19. In addition, both 10% extracts show 16


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stimulatory effect on the growth of MC3T3-E1 cells. Both the pH value and osmolality decrease with the dilution, and the extracts of the as-extruded Mg-2.0Zn-1.0Gd alloy have higher pH value and osmolality compared with the extracts of pure Mg, as shown in Fig. 8 (b) and (c). Fig.8 (d) exhibits the morphologies of MC3T3-E1 cells after 1 and 3 days incubation in different extracts. In comparison with the negative control, there is no obvious difference in cell morphology and cell density between the cells cultured in 50 % and 10 % extracts.

17



Fig. 8 In vitro cytotoxicity assessments: (a) the viability and (d) morphologies of the MC3T3-E1 cells cultured in 10%, 50% and 100% extracts of as-extruded pure Mg and Mg-2.0Zn-1.0Gd alloy; (b) and (c) the pH value and osmolality of extracts. (ZG21 represents the Mg-2.0Zn-1.0Gd alloy)

3.4 In vivo degradation tests The surgical photos and results of reconstruction of micro-CT are shown in Fig. 9 (a), from the degradation morphology of in vivo test, it can be found that both the pure Mg and Mg-2.0Zn-1.0Gd alloy exhibit a basic uniform degradation morphology after 10 days of 18


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implantation. And after implantation for 30 days, the serious localized corrosion pits could be found both in pure Mg and Mg-2.0Zn-1.0Gd alloy, and the Mg-2.0Zn-1.0Gd alloy shows faster degradation rate compared with pure Mg. The in vivo degradation rate calculated by volume loss according to the micro-ct results is shown in Fig. 9 (b), which is 0.11 ± 0.04 mm/year in the first 10 days and reaches 0.31 ± 0.01 mm/year after 30 days of implantation for the Mg-2.0Zn-1.0Gd alloy.

19



Fig. 9 In vivo degradation test: (a) surgical photos and reconstruction of micro-CT results; (b) in vivo

degradation rate; (c) surface morphology and longitudinal section observation. (ZG21 represents the Mg-2.0Zn-1.0Gd alloy)

Further observation of degradation morphology is shown in Fig. 9 (c), the degradation morphology of pure Mg and Mg-2.0Zn-1.0Gd alloy after implantation for 10 days shows no significant difference between each other, which is similar to that of Mg-2.0Zn-1.0Gd samples which were immerged in Hank’s solution for 10 days, and the degradation products became thicker after 30 days of implantation. The EDS results of degradation products are also in agreement with the in vitro tests except for the existence of N element, which may be caused by 20


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some organic matter of the body. The line scanning results of the longitudinal section are shown in Fig. 9 (c), it can be found that there are mainly O and Mg in the whole degradation layer, only little amount of C, P and Ca could be found. The amount of Zn and Gd is so small due to the low alloying addition. The degradation layer becomes thicker with the addition of alloying elements and the increase of implantation time, which is corresponded well with the micro-CT results.

4. Discussion 4.1 Mechanical properties The degradable biomaterials are expected to possess favorable balanced mechanical properties at the early stage after implantation. It has been suggested that alloys with the yield strength of above 200 MPa and the ductility no less than 10 % are acceptable for biomedical applications

18.

In general, thermal mechanical processes are used to improve the mechanical

properties and corrosion resistance of Mg alloys according to the previous studies

20.

In this

study, extrusion was chosen as the thermal mechanical processes method to strengthen the Mg-2.0Zn-xGd alloys. The as-extruded Mg-2.0Zn-1.0Gd alloy exhibits good combination of yield strength (284 MPa) and elongation (24 %), which are both better than those of the other studied as-extruded Mg-Zn-Gd alloys and meet the clinic requirements. This can be explained as follows: (i) compared with the as-extruded Mg-2.0Zn-0.5Gd alloy, the average grain size of the as-extruded Mg-2.0Zn-1.0Gd alloy was smaller due to the increased recrystallization and 21



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restricted grain growth by the tiny secondary phases. According to the Hall-Petch relationship, the yield strength increases with the grain refinement. In addition, there are more dispersed tiny secondary phases in the as-extruded Mg-2.0Zn-1.0Gd alloy after extrusion, which can provide second phase strengthening and then contribute to the strength. As for elongation, grain refinement can also contribute to the improvement of elongation for Mg alloys 20-21, because the non-basal slip systems can be activated due to the grain boundary compatibility effect

22.

In

addition, the weakened basal texture of the as-extruded Mg-2.0Zn-1.0Gd alloy compared to Mg-2.0Zn-0.5Gd alloy, as shown in Fig. 3, may also contribute to the improvement of elongation. (ii) With further addition of Gd, the grain size begins to increase and the basal texture intensity is weakened at the same time, as shown in Fig. 3. Consequently, all these lead to a decrease of the strength for the alloys containing more than 1.0 wt. % Gd. The decrease of elongation is caused by the larger secondary phases in the as-extruded Mg-2.0Zn-1.5Gd alloy and Mg-2.0Zn-2.0Gd alloy, which can cause stress concentration during plastic deformation and promote the initiation and propagation of crack 16. 4.2 In vitro cytotoxicity In general, in vitro cytotoxicity tests are used to evaluate the biocompatibility of the alloys that used for biomaterials. As for the degradable biomaterials, the toxicity can be mainly caused by the following three factors: the metal ion itself, the released ion concentration and the degradation products 23. Firstly, as for the metal ion, there are three elements, Mg, Zn and Gd, in the studied Mg-Zn-Gd alloys. As an important constituent of human body, Mg is crucial for bone 22


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health and about 60 % of Mg in the human body is stored in the bone matrix 3, it can be absorbed or excreted by human body. As a result, Mg has good biocompatibilty and suitable for orthopedic implants application. Zn is also one of the most abundant nutritionally essential elements in human body, the human requirement for Zn is about 8-11 mg/day

1-2.

The

biocompatibility of Gd is acceptable at low concentrations 6 and the Gd containing stents (AMS) have already been implanted into patients and no patient showed allergic or toxic reaction 1. Secondly, the Zn and Gd content of Mg-2.0Zn-1.0Gd is 2.0 wt.% and 1.0 wt.% respectively in this study, and the daily release of Zn and Gd is far lower than the limited value according to the corrosion rate calculated by the mass loss test. This means the daily trace release of Zn and Gd elements is safe enough. Thirdly, the in vitro cytotoxicity assessments also prove that the Mg-2.0Zn-1.0Gd alloy is acceptable as a biomaterial. The MC3T3-E1 cells showed reduced viability in 100% extract of both pure Mg and Mg-2.0Zn-1.0Gd due to the high pH value and osmolality, while the cytotoxicity of 50% and 10% extract was found to be Grade 0-1. The same phenomenon was found in other research, and Willbold et al.

14

found that the viability of MCT3-E1 cells was

reduced in 100 % extract of Mg-Nd alloy due to the high pH value and osmolality resulting from the high ion concentration, while the 50% and 10% extract showed acceptable biocompatibility. Wang et al.

19

has reported that the recommended dilution range for the extracts to perform

cytotoxicity evaluation of Mg-based materials was between 6 and 10 times. Therefore, the cytotoxicity results of 50 % and 10 % extracts of the Mg-2.0Zn-1.0Gd alloy in this study are more reasonable. 23



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Collectively, these results show that the biocompatibility of Mg-2.0Zn-1.0Gd alloy is good and acceptable as biomaterial. 4.3 In vitro and in vivo degradation behavior Except for the mechanical properties and cytotoxicity assessments, degradation behavior is also need to be considered for the degradable biomaterials. In this study, the as-extruded Mg-2.0Zn-0.5Gd and Mg-2.0Zn-1.0Gd alloys exhibit lower degradation rate and more uniform degradation morphology compared with the as-extruded Mg-2Zn-1.5Gd and Mg-2Zn-2.0Gd alloys. This can be explained as follows: the total amount of secondary phases increases with Gd addition and there are more inhomogeneous distributed large size secondary phases when there are over 1.0 wt.% Gd addition. On the one hand, secondary phase can lead to galvanic corrosion due to the difference of electro potential compared with α-Mg

24-25.

Consequently, the

degradation resistance decreased with the increase in the total amount of alloying elements. On the other hand, the cathode-to-anode area ratio is one of the factors that determine the rate of galvanic corrosion

26.

In this study, the inhomogeneous distributed large secondary phases in

as-extruded Mg-2.0Zn-1.5Gd and Mg-2.0Zn-2.0Gd alloys can lead to serious galvanic corrosion and finally form the localized corrosion pits as shown in Fig. 7. It is worth noting that as for the as-extruded Mg-2.0Zn-0.5Gd and Mg-2.0Zn-1.0Gd alloys, the degradation rate calculated by mass loss test is only about 0.15 mm/year and 0.24 mm/year respectively, which is much lower than previously reported Mg alloys for biomedical application, such as AZ91D (2.8 mm/year) 23 and WE43 (0.71 mm/year)

27.

After being implanted into SD rats, the degradation morphology 24


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and degradation rate of Mg-2.0Zn-1.0Gd alloy shows no obvious difference compared with pure Mg. The thicker degradation layer is caused by the addition of Zn and Gd elements, which can form secondary phases and accelerate the degradation. The calculated in vivo degradation rate (0.31 mm/year) for Mg-2.0Zn-1.0Gd is a little faster than the in vitro degradation rate (0.24 mm/year), and this can be explained as follow: the samples were implanted into cranial defect of SD rats and because the cranium is very thin, most surface of the samples were exposed to the in vivo fluid, cells and proteins, which may accelerate the degradation rate. In addition, it has been reported that the Mg alloy was corroded fastest in the head, at an intermediate level in the back, and slowest in the femur

28,

which may also explain the faster corrosion rate in vivo compared

with the corrosion rate in vitro in this study. As for the degradation products, there are mainly C, O, Mg, Ca and P elements for both in vitro and in vivo experiments. Mg in aqueous solution dissolves according to the following equations: Anodic reaction: Mg →Mg 2 + +2e Cathodic reaction: 2H2O + 2e → H2 +2OH ― The chloride ion (Cl-) can transform Mg(OH)2 into soluble MgCl2

2, 29

,this can cause an

increase in OH- ions which can elevate the pH of the solution and can also lead to further dissolution of Mg due to the damage of Mg(OH)2 protection layer. In addition, when there are PO4- and Ca2+ in the solution, they can react with OH- to form hydroxyapatite Ca10(PO4)6(OH)2 (HA)

2, 29-30.

Previous studies

31

have confirmed that these degradation products show good

biocompatibility and even can promote the osteoinductivity. Thus, it can be concluded that the 25



biocompatibility of degradation products of Mg-Zn-Gd alloy is good and acceptable as biomaterial. Taken together, the Mg-2.0Zn-1.0Gd alloy shows good biocompatibility from the results of in vivo degradation tests and in vitro cytotoxicity assessments, and it could be further applied for potential orthopedic implants.

5. Conclusions The microstructure, mechanical properties, cytotoxicity and in vitro and in vivo degradation behavior of Mg-2.0Zn-xGd (x=0.5, 1, 1.5, 2 wt.%) alloys with different Gd addition were investigated systematically. The following main conclusions can be drawn: (1) The Mg-2.0Zn-xGd alloys are mainly composed of α-Mg matrix, I-phase, Mg2Zn3 phase and W-phase. The grain size decreases firstly when the Gd addition increases from 0.5 wt.% to 1.0 wt.% and then increases with further addition of Gd. The volume of the secondary phases increased with the increase of Gd concentration. (2) The basal texture is gradually weakened with the addition of Gd. After extrusion, the as-extruded Mg-2.0Zn-1.0Gd alloy exhibits the best mechanical properties, (ultimate strength of 338 MPa, yield strength of 284 MPa and elongation of 24%) which is attributed to the finer grain size, dispersed tiny secondary phases, and weakened texture. (3) The in vitro degradation rate of as-extruded Mg-2.0Zn-xGd alloys is 0.15 mm/year, 0.24 mm/year, 0.42 mm/year and 1.19 mm/year corresponding to the x value of 0.5, 1.0, 1.5 and 2.0, respectively. This is caused by the coarser microstructure and more secondary phases which can 26


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lead to more serious galvanic corrosion. The uniform degradation morphology can be found when the addition of Gd is lower than 1.0 wt.%, and further addition of Gd leads to localized corrosion. (4) The in vivo degradation products and degradation morphology of Mg-2.0Zn-1.0Gd alloy is similar to those of pure Mg, and the results of the in vitro cytotoxicity tests also show no significant difference between Mg-2.0Zn-1.0Gd alloy and pure Mg. Overall, after the addition of 2.0 % Zn and 1.0 % Gd to pure Mg, the mechanical properties raised dramatically. The as-extruded Mg-2.0Zn-1.0Gd alloy shows the excellent mechanical properties, good biocompatibility and degradation resistance, indicating a promise biodegradable orthopedic implants material.

Acknowledgements The authors gratefully acknowledge the support by the National Natural Science Foundation of China (No. 51501110, No. 51728202), the Science and Technology Commission of Shanghai Municipality

(No.

17XD1402100),

Shenzhen’s

Three

Renowned

Project

(No.

SZSM201612092), the Medical Support Project of Science and Technology Commission of Shanghai

Municipality

(No.

15411951200),

the

Shanghai

Jiao

Tong

University

Medical-engineering Cross Fund (No. YG2016QN03), the 111 Project (Grant No. B16032).

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For Table of Contents Use Only

Research on biodegradable Mg-Zn-Gd alloys for potential orthopedic implants: In vitro and in vivo evaluations Hongwei Miao, Dandan Zhang, Chenxin Chen, Lei Zhang, Jia Pei, Yun Su, Hua Huang, Zhongchang Wang, Bin Kang, Wenjiang Ding, Hui Zeng, Guangyin Yuan.

31



Fig. 1 XRD results of Mg-2.0Zn-xGd alloys


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Fig. 2 Optical microstructure images and analyzed results including average grain size and secondary phase of the as-extruded Mg-2.0Zn-xGd alloys: (a) optical micrographs; (b) average grain size; (c) volume fraction of secondary phases.



Fig. 3 EBSD results of the as-extruded Mg-2.0Zn-xGd alloys: (a) Mg-2.0Zn-0.5Gd；(b) Mg-2.0Zn-1.0Gd; (c) Mg-2.0Zn-1.5Gd; (d) Mg-2.0Zn-2.0Gd. (RD=extrusion direction, TD=radius direction)


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Fig. 4 Mechanical properties of the as-extruded Mg-2.0Zn-xGd alloys 249x175mm (150 x 150 DPI)



Fig. 5 In vitro degradation rate of the as-extruded Mg-2.0Zn-xGd alloys


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Fig. 6 Surfaces observation and EDS results of degradation products for the as-extruded Mg-2.0Zn-xGd alloys after immersion in Hank’s solution for 10 days





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Fig. 9 In vivo degradation test: (a) surgical photos and reconstruction of micro-CT results; (b) in vivo degradation rate; (c) surface morphology and longitudinal section observation. (ZG21 represents the Mg2.0Zn-1.0Gd alloy)


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Fig. 9 In vivo degradation test: (a) surgical photos and reconstruction of micro-CT results; (b) in vivo degradation rate; (c) surface morphology and longitudinal section observation. (ZG21 represents the Mg2.0Zn-1.0Gd alloy)


Research on biodegradable Mg-Zn-Gd alloys for potential orthopedic

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