In vivo degradation and bone reaction of long-term fixation with a magnesium alloy made by twin-roll casting in a rat femur model

Abstract

BACKGROUND:

The effect of casting parameters on the microstructure and corrosion resistance of Mg alloys is still limited, especially in clinical animal experiments.

OBJECTIVE:

We prepared a new magnesium rare earth alloy (Mg-Re, where Re is Ce or La) by vertical two-roll casting and Mg-A by further rolling. The microstructure characteristics, degradation behavior, and bone reaction of the two alloys were studied.

METHOD:

Ti, Mg-Re, and Mg-A alloy plates were implanted in a rat femur model, and their degradation behavior was observed 48 weeks later.

RESULTS:

In vivo experiments showed no significant changes around the femur in the Ti group, excluding external factors that may cause bone remodeling and lead to new bone formation. Mg-A induces more new bone formation than Mg-Re, which meets the necessary conditions to prevent pathological fracture. The specimen staining and sectioning showed that the liver and heart of rats implanted with magnesium alloys had no pathological changes and the cell structure was normal, similar to that of rats without a magnesium alloy.

CONCLUSION:

Mg-A alloy has good healing potential as a biodegradable implant material.

Keywords

Mg-Re sheet in vivo amorphous degradable bone response

1. Introduction

Bone transplantation is one of the most frequently performed tissue transplantation procedures worldwide, second only to clinical blood transfusions. Metal materials play a very important role in treating orthopedic injuries or replacing bone tissue in modern medicine [1]. Magnesium alloys have higher mechanical strength and fracture toughness, and are more suitable for load-bearing applications. Compared with stainless steel, titanium, and cobalt-chromium alloy, Mg metal has better biocompatibility, does not produce toxic metal ions in the degradation process, can be absorbed by the human body, and does not require a second operation to remove, greatly reducing medical costs and patient pain. In addition, Mg is an essential element to maintaining daily activity, accounting for approximately 0.05% of the total body weight, more than half of which exists in the skeleton. Various in vitro and in vivo experiments have proven that Mg has good biocompatibility and biosafety [2–4]. However, pure Mg corrodes too quickly under normal physiological conditions, losing mechanical integrity before bone tissue can properly heal, and producing hydrogen faster than the host tissue can process [5,6]. Therefore, although magnesium as an orthopedic implant has had some success, there is still room for development.

In the past ten years, Mg alloys containing rare earth elements have attracted attention. The addition of rare earth elements such as Y, La, Ce, Nd, Sm and Gd improves the plasticity and strength of the Mg alloy, and reduces the anisotropy and tension/compression asymmetry of the magnesium alloy compared with AZ31 and other common Mg alloys [7–10]. In addition to improving the ductility and strength, Mg alloys containing rare earth elements also have high creep resistance and corrosion resistance [11–13]. In addition to the alloy composition, the corrosion resistance of Mg is also affected by its internal microstructure characteristics such as the grain boundary (GB) [14–16], precipitated phase [17–19], grain orientation [20–23], and dislocation [24]. Most biodegradable Mg alloys are deformed alloys, and their mechanical properties are superior to casting alloys because of the fine grain strengthening effect produced by plastic deformation [25,26].

Previous studies prepared magnesium rare earth (Mg-Re) alloy sheets using a vertical two-roll continuous casting machine (VTRC). Microscopic characterization experiments show that the resulting crystal structure is amorphous. Moreover, Mg-Re has better corrosion resistance than intermediate alloys [27]. However, the effect of casting parameters on the microstructure and corrosion resistance of Mg alloys is still limited, especially in clinical animal experiments. Understanding the corrosion behavior of magnesium alloys in different casting methods is of great significance for expanding the application of Mg-based biomedical implants.

On the basis of previous studies, a Mg-Re alloy was prepared by the VTRC process, and an Mg-A alloy was prepared by rolling the Mg-Re alloy. The microstructures, in vivo degradation behaviors, and bone reactions of Mg-Re (Ce, La), Mg-A, and titanium implants were observed in a rat femoral model for up to 48 weeks. The long-term effects of the new Mg-Re alloy on biological systems and local vital organs were further analyzed.

2. Methods

2.1. Material preparation and characterization

In this experiment, high purity Mg ingot (99.99%), AZ31 ingot, Mg-10LA ingot, and Mg-20Ce ingot (wt.%) were used as raw materials. Smelting occurred in a resistance furnace protected by high purity argon gas. The AZ31 sheet was prepared by Mg ingot and AZ31 ingot smelting, and the Mg-Re sheet was prepared by Mg ingot, AZ31 ingot, Mg-10La, and Mg-20Ce smelting. First, the Mg ingot and AZ31 ingot were placed in the resistance furnace and heated to 720 °C for 1.5 h to completely melt the compound. Mg-10La and Mg-20Ce ingots were then added to the melt and the temperature was reduced to 700 °C and held for 1.5 h, producing the Mg-Re alloy. A rectangular 50 × 100 × 10 mm³ Mg-Re alloy ingot plate was hot rolled at 250 °C, reducing the total thickness of the composite sample by 88.0% to produce the Mg-A sample.

As shown in Fig. 1, the I process uses VTRC technology to cast Mg-Re (Ce, La) alloy, and the II process uses additional 250 °C hot rolling to produce another Mg alloy, Mg-A.

Fig. 1.

Schematic diagram of the magnesium alloy sheet manufacturing process. A new Mg-rare earth alloy (Ce, La) was prepared by vertical twin-roll casting (VTRC) and additional rolling to obtain another Mg-A alloy.

2.2. Implant method

Sixteen rats (Wistar; male, 30 weeks of age, 370 ± 15 g) were selected from the University of Tokyo Medical School. All animal experiments were reviewed by the Animal Care and Use Committee of Saitama Institute of Science and Technology (Approval No. 2020-1). The rats were randomly divided into 4 groups. Rectangular Mg-Re and Mg-A tablets with a thickness of 4 mm × 13 mm and 0.5 mm were implanted into the femoral fixation model of rats in groups 1 and 2, respectively. The third group was high purity titanium of the same size, the fourth group was the same age rats without implants as control. All Mg alloy implants were fixed with AZ31 screws and titanium implants were fixed with titanium screws. The implants were disinfected with ethylene oxide before surgery. The implantation process is shown in Fig. 2. Daily clinical observations were performed after surgery. Postoperative pain and distress were observed for expressions of behavioral abnormalities, food, and water intake. Intramuscular injection of buprenorphine hydrochloride (0.01–0.05 mg/kg; three times a day; one week) and carprofen (5 mg/kg; weekly; four weeks) analgesics were used to inhibit postoperative pain. Penicillin was administered to prevent postoperative infection (0.2 mg/kg; three times a day; one week).

Fig. 2.

The implant surgery was performed in Wistar rats: (a) anesthetized rats; (b) mark the corresponding implant location; (c) drill holes in the femur of rats; and (d) fixation of alloy plates with bone screws.

The femur and viscera of rats with high purity titanium and the Mg alloys were implanted 48 weeks after operation. Micro-CT, histological, and physiological analysis were performed. The alloy sheets were explanted and the microstructure was analyzed.

2.3. In vivo corrosion assessment

Electron backscatter diffraction (EBSD) electron microscopy (SEM, Gemini300, Germany & EBSD, Oxford C-Nano, UK) was used to characterize the sample microstructure including GB, grain orientation, texture, and residual strain. Channel 5 analysis software was used to analyze the EBSD results. EBSD measurement samples were prepared by conventional grinding and diamond polishing (0.5 μm), and Ar⁺ ion milling was performed in a precision ion polishing system (M4000PLUS). Imaging was performed under a liquid nitrogen cooling system at an operating acceleration voltage of 3 kV and incident angle of 40°.

2.4. Micro-CT analysis

After implantation for 2, 8, 16, 32, and 48 weeks, 2D cross-sections and 3D models of the thigh bone and metal plates were taken using micro-CT (R. m-Ct. Rigaku, Tokyo, Japan) to evaluate the formation of new bone.

2.5. In vivo degradation tests

After 48 weeks of implantation, the cross-section of the implant was observed by SEM and the structure of the degradation layer was analyzed.

2.6. Histological observation

Autopsies were performed 48 weeks after implantation. The histological sections of the heart and liver were evaluated using 4 μm conventional paraffin-embedded formalin staining and eosin (H&E) staining. Liver sections were stained by PAS staining, also known as periodate schiff staining and glycogen staining, to reveal glycogen and other polysaccharides.

2.7. Serum diagnosis

After surgery, 10 ml of blood from the heart of rats in each group was taken. The blood was centrifuged and 3 ml of the serum supernatant was sent to Yeast Industry Co., Ltd., Japan for serum diagnosis to categorize the physiological health of the rats.

3. Results and discussion

3.1. Microstructures

The microstructure evolution of Mg-Re and Mg-A magnesium alloys along RD, TD, and ND directions was studied using EBSD. Figure 3 shows the inverse pole figure (IPF) diagrams of the RD, TD, and ND samples. The IPF figure shows the microstructures of the Mg-Re alloy prepared by VTRC and the Mg-A alloy hot rolled from Mg-Re along the RD to the true strain level shown in the figure. The compression direction RD is the rolling direction. The compression direction TD is the horizontal direction. The compression direction ND is the rolling direction. The colors in the figure indicate the direction of the compression axis relative to the local lattice frame according to the IPF triangle. Compared with Mg-Re, Mg-A showed extrusion deformation during the manufacturing process and grain growth and recrystallization during solidification.

Fig. 3.

EBSD micrographs of (a,c,e) an unrolled magnesium alloy Mg-Re and (b,d,f) hot rolled magnesium alloy Mg-A.

As shown in Figs 4a and 4b, the low-angle grain boundaries (LAGB) of Mg-Re and Mg-A are 38.3% and 35.8%, respectively. Figure 4c shows that the average particle size of Mg-Re was 25.5 μm. Mg-A shows dynamic recrystallization after hot rolling at 250 °C; however, the grain size was not uniform. As shown in Fig. 4d, the average particle size decreases rapidly to about 11.7 μm. Previous studies have shown that local orientation errors within grains lead to higher corrosion rates. The larger the LAGB value is, the larger the local orientation error is. In the absence of local orientation error, grain orientation has limited influence on the corrosion rate [28]. After hot rolling at 250 °C, the GB orientation deviation of Mg-A is less than that of Mg-Re; therefore, the corrosion rate of Mg-A is less than that of Mg-Re.

Fig. 4.

The distributions of (a,b) grain boundary misorientation and (c,d) grain size.

3.2. Bone response and histological examination

Figure 5(a) shows micro-CT images and 2D cross-sectional images of the thigh bone of rats implanted with the Mg alloys and titanium. During the corrosion process of the Mg-Re plate during implantation, a small amount of new bone formed between the bone and Mg alloy lamina 4 weeks after surgery, and further growth was observed at 12 weeks. After 24 weeks, the boundary between the plate and bone became blurred because of new bone tissue. After 48 weeks, the Mg alloy plate and bone end gap were nearly completely covered by new bone. In the corrosion process of the Mg-A plate, a small amount of new bone also appeared 4 weeks after surgery and continued to grow. At 48 weeks, the Mg alloy plate and bone end gap were nearly completely covered by new bone. During the corrosion of the titanium plate implants, numerous biological tissues appeared on the Ti screws and Ti plates at 8 weeks and 24 weeks postoperative. No significant new bone tissue was formed continuously between the lamina and bone. After 32 weeks and 48 weeks of surgery, the biological tissue on the titanium screw and plate disappeared and no signs of new bone growth were observed. Four weeks after surgery, the boundary between the Ti plate and thigh bone was more obvious than that between the two Mg alloy plates. On the basis of the increase in corrosion time, the bone growth of the two Mg alloy plates was better than that of the Ti plate.

Fig. 5.

(a) Representative 3D micro-CT reconstructions showing bone response at 4, 12, 24, and 48 weeks postoperative with thin Ti, Mg-Re, and Mg-A implants. Red box is the area of interest. (b) Implants after 48 weeks with hematoxylin and eosin (H&E) staining of decalcified sections showing the morphology of existing and new bone around the implant. The blue box represents the part of bone not in contact with the implant and the stained section of this part. (c) Femur parameters were statistically analyzed according to the horizontal section area of interest. BV/TV: bone volume/tissue volume (^∗ P < 0.05, ^∗∗ P < 0.01).

Figure 5(b) shows the stained section of rat thigh bone after 48 weeks, with bone pink and bone marrow purple. Starting from the central part of the fixed metal (red dotted box in Fig. 5a), some of the bones in the fixed Mg alloy changed shape significantly, forming a mature new bone structure, while the bones and marrow did not change at all in the thigh bone of the unfixed Mg alloy plate (blue dotted box in Fig. 5a). No significant changes in bone or marrow were observed in the fixed titanium plates.

Figure 5(c) shows that a large mature new bone structure was formed at the interface of the two Mg alloys, and the direction of bone formation was toward the implant. The periosteum formation and femur volume increase of new bone were similar between the two Mg alloy groups. At 4 weeks, the increased volume of femur implanted with the Mg-Re plate was greater than that with the Mg-A plate, and at 12 weeks, the increased volume of the femur with the Mg-A plate was greater than that with the Mg-Re plate. Previous studies confirmed the beneficial effect of Mg²⁺ on new bone formation, promoting bone cell adhesion and bone tissue growth [29–32]. According to the SEM analysis of EBSD in this study, the degradation rate of Mg-A is less than that of the Mg-Re plate because of the higher corrosion resistance of the Mg-A plate. With the plate degradation, Mg²⁺ stimulates local tissue, forming new bone. Therefore, the femoral volume around the implanted Mg-Re plate was greater than that of the Mg-A plate and titanium. However, with Mg base material degradation, excessive ion release can inhibit the formation of new bone tissue [33] from implantation through 12 weeks. The increase in femur volume after implantation of the Mg-A plate was greater than that with the Mg-Re plate and titanium plate. In this study, the surgery was performed when the bones were completely healthy. Therefore, whether Mg-Re has a therapeutic effect on the clinical healing of a fracture or bone injury needs further study. No significant changes were observed near the femur after implantation of the nonabsorbable titanium, which rules out external factors that may cause bone remodeling by altering the mechanical load of the implant.

According to these research results, there is a relationship between different casting methods and the properties of the Mg-Re alloys, including biological conditions. The rolled Mg-Re alloy can produce larger grain size and lower orientation deviation, which improve the corrosion resistance of an Mg alloy. In vivo results confirmed that the Mg-A alloy had better bone formation ability and corrosion resistance over long term implant.

3.3. In vivo degradation

Figure 6 shows the representative cross section of the alloy sample and the corresponding SEM image. Characteristic corrosion layers of the two Mg alloys were observed as double layers. Degradation layers of different thickness were observed in the Mg alloys. The inner thickness of the Mg-Re alloy was approximately 30 μm and the outer thickness was approximately 45 μm. In addition, there were microcracks in both the inner and outer layers. Microcracks may be caused by accelerated dissolution of the degradation products. After rolling, the inner thickness of the Mg-A alloy was approximately 10 μm, and the outer thickness was 40 μm. The double layer structure was dense and uniform, and no obvious microcracks were observed.

Fig. 6.

SEM images of alloy plates implanted in rat thigh bone after 48 weeks: profiles of the degradation layer of (a) the Mg-Re alloy and (b) Mg-A alloy.

Mg and its alloys dissolves in body fluids based on following equations: $\begin{eqnarray}\displaystyle \text{} & \text{} & \displaystyle \text{Mg}\rightarrow \text{Mg}^{2+}+2\text{e};\quad \text{anodic reaction}.\nonumber\\ \displaystyle \text{} & \text{} & \displaystyle 2\text{H}_{2}\text{O}+2\text{e}\rightarrow \text{H}_{2}+2\text{OH}^{-};\quad \text{cathodic reaction}.\nonumber\\ \displaystyle \text{} & \text{} & \displaystyle \text{Mg}+2\text{H}_{2}\text{O}\rightarrow \text{Mg(OH)}_{2}+\text{H}_{2};\quad \text{Overall reaction}.\nonumber\end{eqnarray}$

Mg(OH)₂ changes MgCl₂ under the action of the cathodic ion (Cl⁻), which can increase the PH of the solution and result in further dissolution of Mg [34,35]. Furthermore, Ca²⁺ and PO₄ ³⁻ in body fluids reacts with OH⁻ to form Ca₁₀(PO₄)(OH)₂ [36,37]. Previous research [38–40] report studied Mg and C, O, P, Ca and other elements in the corrosion layer. It is proved that these degradation products can promote newly formed bone and show good biocompatibility. Therefore, it could be concluded that the degradation product of the Mg-RE alloy is acceptable and good. However, the dissolution mechanism of the degradation layer remains to be further studied. Mg and other elements such as C, O, P and Ca in corrosion layer were studied.

3.4. Histological examination after implantation

Figure 7 shows the histological appearance of the liver and heart after 48 weeks in vivo. The size and morphology of cells and nuclei in the experimental groups and control group have no significant difference. The rats implanted with Mg alloy for a long time had no liver injury and normal liver cell structure, similar to the rats without a Mg alloy. Compared with the control group, the experimental group showed no morphological changes and no edema of the myocardium and nucleus after HE staining. No inflammation or tissue damage was observed in the three experimental groups.

Fig. 7.

Representative HE staining and PAS staining of the liver and representative HE staining of the myocardium 48 weeks after transplantation.

3.5. Serum test results

According to Table 1, there were no significant differences in total protein, albumin, total cholesterol, triglyceride, and glucose between the control group and titanium alloy group. The concentrations of sodium, potassium, chloride, calcium, and magnesium in the serum of the four groups were measured. The concentration of Mg ions in the Mg alloy group were slightly greater than those in the other groups. Elevated creatinine levels indicate decreased kidney function. Slightly higher concentrations of creatinine were observed in rats with rolled Mg alloys, which may be associated with higher concentrations of Mg ions. The AST (glutamic oxalacetic transferase) and ALT (glutamic-pyruvic transaminase) values of the control group, titanium alloy group, and two Mg alloy groups were similar, AST and ALT are indicators of early myocardial infarction, hepatocyte necrosis, degeneration, cirrhosis, liver cancer and other diseases, the liver function of rats in these groups did not change significantly because of the alloy implantation.

Table 1
Serum test results of rats 48 weeks after implantation of the Mg alloys, titanium, and control

TP (g/dL) ALB (g/dL) CRE (mg/dL) Na (mEq/L) K (mEq/L) Cl (mEq/L) Ca (mg/dL) Mg (mg/dL) AST (IU/L)

Control group 6.2 4.2 0.32 138 5.1 98 9.8 2.1 59

Ti 6.5 4.4 0.33 145 4.6 104 9.7 2.1 72

Mg-A 6.3 4.3 0.37 138 5.4 97 10.1 2.5 57

Mg-Re 6.2 4.3 0.35 140 4.6 100 10 2.2 54

ALT (IU/L) LDH (IU/L) AMY (IU/L) r-GT (IU/L) T-CHO (mg/dL) TG (mg/dL) HDL-C (mg/dL) T-BIL (mg/dL) GLU (mg/dL)

Control group 36 288 1389 3 > 100 361 28 0.07 373

Ti 31 290 1492 3 > 92 338 29 0.06 303

Mg-A 36 277 1436 3 > 112 292 34 0.06 317

Mg-Re 34 251 1499 3 > 121 342 30 0.08 336

	TP (g/dL)	ALB (g/dL)	CRE (mg/dL)	Na (mEq/L)	K (mEq/L)	Cl (mEq/L)	Ca (mg/dL)	Mg (mg/dL)	AST (IU/L)
Control group	6.2	4.2	0.32	138	5.1	98	9.8	2.1	59
Ti	6.5	4.4	0.33	145	4.6	104	9.7	2.1	72
Mg-A	6.3	4.3	0.37	138	5.4	97	10.1	2.5	57
Mg-Re	6.2	4.3	0.35	140	4.6	100	10	2.2	54
	ALT (IU/L)	LDH (IU/L)	AMY (IU/L)	r-GT (IU/L)	T-CHO (mg/dL)	TG (mg/dL)	HDL-C (mg/dL)	T-BIL (mg/dL)	GLU (mg/dL)

Control group	36	288	1389	3 >	100	361	28	0.07	373
Ti	31	290	1492	3 >	92	338	29	0.06	303
Mg-A	36	277	1436	3 >	112	292	34	0.06	317
Mg-Re	34	251	1499	3 >	121	342	30	0.08	336

4. Conclusion

The microstructure, corrosion behavior, and in vivo bone reaction of Mg-Re alloys prepared by VTRC and hot rolling, as well as the effect of long-term fixation on systemic circulation, were discussed. Through the systematic analysis, the following conclusions are drawn:

Mg-A has less local orientation error and higher corrosion resistance than Mg-Re.

Mg-A creates more bone tissue than Mg-Re.

An in vivo implant degradation test showed that the degradation layer presented a two-layer structure.

Staining of the rat heart and liver samples showed that the corrosion reaction of the Mg alloy plates did not cause an abnormal cell morphology.

The serum biochemical test results of the Mg alloy groups were consistent with that of the control group, except for higher concentrations of Mg ions and creatinine. The long-term implantation of the Mg alloys will not harm the circulation or immune system of the organism, indicating no harm to humans.

Footnotes

Acknowledgements

The authors thank Ashleigh Cooper, PhD, from Liwen Bianji (Edanz) () for editing the English text of a draft of this manuscript.

Conflict of interest

None to report.

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