Biodegradable and Bioactive Orthopedic Magnesium Implants with

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Cite This: ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

Biodegradable and Bioactive Orthopedic Magnesium Implants with Multilayered Protective Coating Yixuan Li,‡ Jing Gao,† Liyun Yang,† Jian Shen,† Qing Jiang,‡ Chi Wu,§ Dan Zhu,*,† and Yifeng Zhang*,‡

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†

Jiangsu Key Laboratory of Bio-functional Materials, School of Chemistry and Materials Science, Nanjing Normal University, Nanjing 210023, China ‡ Department of Sports Medicine and Adult Reconstructive Surgery, Drum Tower Hospital, Medical School of Nanjing University, Nanjing 210093, China § Department of Chemistry, The Chinese University of Hong Kong, Shatin, N.T., Hong Kong ABSTRACT: Biocompatible and biodegradable magnesium (Mg) and its alloys possess excellent potential as orthopedic implant biomaterials, but they corrode too quickly inside the human body to maintain such functions as supporting and fixation. In this study, we have successfully prepared a composite biomedical material with Mg as the substrate and a multilayered coating on the metal, which contains a chemical conversion layer of fluoride or phosphate, an adhesion layer, and a layer of biodegradable polylactic acid (PLA). The recorded weight loss and the release of ionic magnesium (Mg2+) in the simulated body fluid (SBF) demonstrate that the corrosion rate of the protected Mg has been reduced to 1/ 10 that of the unprotected bare Mg. The corrosion rate of the protected Mg can be adjusted with either the chemical conversion type or the thickness of the PLA layer. The bare and the protected Mg has been manufactured as an implant for bone defect repair of Sprague−Dawley rats. The X-ray and micro-CT images show an improved anticorrosion of the protected Mg in vivo, and the histological analysis indicates that the Mg2+ ions released at an appropriate rate are beneficial for bone growth. The preliminary tests prove that the biodegradable Mg has been effectively protected by the innovative multilayered coating, and the implant has met the required biological functions such as biocompatibility and promotion of bone growth. KEYWORDS: orthopedic implant, biodegradability, magnesium implant, protective surface modification/coatings, anticorrosion, bone defect repairing

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INTRODUCTION

medical functions, including strength and biocompatibility over the whole process of treatment, especially in the short term after the surgery. After the implantation, corrosion resistance is required, and degradation and decay rate shall be low enough for the required mechanical support. After the treatment is accomplished, the alloy shall degrade completely, while the alloying elements must be nontoxic to cells or organisms. Another convenient way to improve the corrosion resistance of magnesium metal is surface modification or surface treatment, which usually refers to a method to obtain a layer on the surface of the substrate without destroying the mechanical properties. Such a layer may be organic or inorganic, may arise from deposition or plating, or may be chemically converted from the metal itself. In this paper, in order to distinguish the physical polymer coating from others, we wish to identify the

Biocompatible and biodegradable magnesium (Mg) and its alloys are potential orthopedic implant materials because their mechanical properties are similar to those of natural bones.1−3 Moreover, magnesium ions (Mg2+) can promote bone growth because they adjust enzymatic activities and adenosine triphosphate storage in the cells, make the osteogenitor cells more attachable and differentiated, and, at the same time, enhance the mineral deposition.4,5 However, the degradation of Mg or Mg alloys inside the body is too quick to meet the mechanical supporting requirements after an orthopedic operation.6−8 In order to avoid such weaknesses including the fragility of magnesium products and the rapid degradation in the body, there are mainly two strategies: one is alloying magnesium with other metals,2,9 and the other is the surface protective treatment of magnesium.10−13 Both will improve the anticorrosion property of magnesium. In the first case, the ideal biomedical alloy shall maintain its mechanical and © XXXX American Chemical Society

Received: April 15, 2019 Accepted: June 26, 2019 Published: June 26, 2019 A

DOI: 10.1021/acsabm.9b00313 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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ACS Applied Bio Materials chemistry or electrodeposition of bioactive or inert ceramics, anodic oxide films, etc., as surface treatment, and the polymer film coated on the metal surface as surface coating. There are many related technologies in industry, too; however, some of the treatments do harm to human health, such as surface treatment of chromate or aluminum plating. Some treatments cannot meet the biodegradable requirements, such as anodic oxidation or microarc oxidation to form a ceramic layer, a titanium layer, or most of the organic coatings, which are difficult to degrade in the human body environment. In addition, some protection coatings are theoretically workable but are not able to achieve actual protection. For example, this suggests the idea of coating the surface of a magnesium/ magnesium alloy implant with a layer of biocompatible and biodegradable polymers, e.g., polylactic acid (PLA), since they can theoretically delay the corrosion of the metal and prolong the service life of implanted devices. Also, they themselves can degrade over the long term, and the degradation products do no harm; they are already used as biomaterials or biomedical scaffolds. However, in the coating of PLA onto the metal surface, there is weak adhesion between the polymer and the metal, and in the physiological environment, with the degradation of PLA some defects appear in the film which make the electrolytic body fluid penetrate when in contact with the metal, causing the degradation of magnesium. The metal degradation produces hydrogen, leading to more and more crack defects of the film and the film peeling off the metal surface. In summary, though, with the development of biomedical materials and apparatus, theoretically there are many methods for surface protection coatings on magnesium/ magnesium alloys; there has not yet been such an ideal surface treatment reported that can fully meet the requirements of effective protection through a simple process which can be commercialized, can achieve a controlled corrosion rate and mechanical decay, and ultimately can completely be degraded into some harmless products. In our study we have provided a biomedical magnesium or magnesium alloy implant material or device with multiple protective layers on the surface. The protective layers can effectively delay the mechanical decay, improve the corrosion resistance of the magnesium in the physiological environment, and guarantee the biomedical service of the material or device made of magnesium. With some technical adjustment, such multilayered protection also works for magnesium alloy implants; a series of in vitro and in vivo results of the magnesium alloy of WE43 will be published in our later work. WE43 is a kind of Mg alloy, with the other elements Y 3.7− 4.3%, Nd 2.4−4.4%, and Zr 0.4−1%; it is a high strength casting alloy and can be used in biomedical implants. It can be used in biomedical implants. The multiple protective layers include three layers, as shown in Figure 1. The first layer attached to the surface of the Mg substrate is a chemical conversion coating with a thickness of several nanometers, which can be obtained through a fluoride or a phosphate treatment. The second layer is a cyanoacrylate adhesive, which was approved by the U.S. Food and Drug Administration (FDA) as a tissue adhesive in 1998.14 The third layer is the biodegradable polymer film via dip-coating of polymer solution and then evaporating the solvent. The biocompatible and biodegradable PLA has been chosen, which is also FDA approved, and these materials can be easily implemented on an implant with a complicated shape with a desired thickness. The composite of PLA with magnesium salt can make bioactive and

Figure 1. Schematic of the section view of the Mg/Mg alloy pin with multilayered protections: (1) metal substrate, (2) chemical conversion layer, (3) adhesive layer, and (4) degradable polymer layer.

biodegradable materials.15 Such multiple protective coatings can effectively reduce the initial corrosion rate and the mechanical decay of the metal matrix in the physiological environment, improve its corrosion resistance, and effectively extend its service life. The in vitro experiments have shown that the corrosion rate of Mg samples can be reduced to 1/10 that of unprotected samples. A rat tibia defect model has been used for the surgeries of the implantation of untreated bare Mg and surface-protected Mg, as well as for the control of titanium (Ti), and the degradation behavior and healing process have been evaluated and compared to each other. The results have shown that the treated Mg metal possesses retarded corrosion and bone-growing stimuli in bone defect healing. The biomedical implant prepared by the above treatments includes the metal matrix, chemical conversion coating, adhesive layer, and biodegradable polymer film with an adjustable thickness of 10−300 μm. The preparation process is simple and applicable, and these materials may find applications in a variety of degradable biological implant devices such as vascular stents, bone nails or connections, hemostatic clamp, or screw, etc.

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EXPERIMENTAL SECTION

Materials. Mg metal plates or wire with a purity of 99.95% have been purchased from Taichen Co., China. PLA (AI-1001, MW 100 000 Da) with a density of 1.25 g/cm3 has been purchased from Shenzhen Esun Industrial Co., Ltd. The metal has been fully polished and ultrasonically cleaned with acetone, ethanol, and water before further treatment or coating. The chemicals hydrofluoric acid, phosphoric acid and phosphate salts, ethanol, and acetone, etc., are analytical reagents and have been used without further purification. Chemical Conversion Coating. The chemical conversion is a protected coating layer on the surface of the metal containing metal compound and is produced by electrochemical or chemical reaction of metals.16−18 In our work, considering the biotoxicity of hexavalent chromium, a chrome-free solution has been targeted, and the fluoride or phosphate treatment has been adopted. In the fluoride conversion, the cleaned metal is put in the aqueous bath of hydrofluoric acid with a concentration of 10−40 wt % at room temperature for 24 h. Hydrofluoric acid is highly corrosive and is a contact poison with the potential for deep, initially painless burns and ensuing tissue death. One should be very cautious when dealing with hydrofluoric acid. During phosphate conversion, in an aqueous bath containing potassium permanganate (100 g/L) and diammonium hydrogen B


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ACS Applied Bio Materials phosphate (20 g/L), whose pH is adjusted to 3.5 by using phosphoric acid (75 wt %), the metal is treated at 35 °C for 3 min. The chemically treated metal is rinsed in water and dried under nitrogen flow for further coating. PLA Coating. A solution of PLA in chloroform with a concentration of 4% (w/v) was prepared by full dissolution via chloroform reflux above the glass transition temperature of PLA. The chemically treated metal plate or wire is dip-coated with a thin layer of the adhesive cyanoacrylate, then dip-coated with the PLA solution, and baked under an infrared heater. The residue solvent is removed completely in a vacuum oven at temperatures of 40−50 °C for days. The thickness of the obtained coating films is estimated with the weight difference before and after the PLA coating, the surface area of the metal, and the density of the PLA (1.25 g/cm3). The thickness is in the range 50−150 μm and can be adjusted with the coating times. The metal plate and wire after surface treatment has multiple protective layers and is ready for further tests or implantation. Characterization and Tests. The microscopic morphology of the chemically treated and untreated metal surfaces has been imaged by scanning electron microscopy (SEM, JSM-5610LV, JEOL, Japan). Each sample has been platinum plated before its measurements. The secondary electron image (SEI) has been used to image the surface morphology and to detect the chemical components. The adhesion between the PLA coating film and the metal substrate has been estimated by the standard grid test (GB/T 9286-1998 or ISO 24091992) of paint coatings. A cutter has been used to cut through the coating film on the plate with the grid of 1 × 1 mm2; an adhesive tape is pasted at the surface for minutes and is torn from the surface. The film grids peeled off with the tape from the metal surface are recorded. The weight loss rates of the samples in the simulated body fluid (SBF) are used to evaluate the in vitro degradation of the pristine or protected metals or alloy. For preparing 1000 mL of SBF, NaCl 8.035 g, NaHCO3 0.355 g, KCl 0.225 g, K2HPO4·3H2O 0.231 g, 1 M HCl 39 mL, Na2SO4 0.072 g, and Tris 6.118 g were used. The samples with total apparent surface area of about 15 cm2 (five Mg rods with diameter of 1.2 mm and lengths of 72, 74, 76, 78, 80 mm, respectively) have been completely immersed in 150 mL of Kokubo solution19 at an experimental temperature of 37 °C for a period of testing time from 1 week to 4 weeks. The corrosive product at the metal surface is removed with an aqueous solution of 15 wt % CrO3 for seconds and completely dried before weighing. The corrosion rates obtained from the weight loss have large errors. We have collected data from different batches of samples, and samples of different shapes and sizes. The amounts of dissolved Mg2+ ions in the SBF have also been determined by using an atomic adsorption spectrometer (Avanta, GBC); the wavelength used for Mg 2+ concentration detection is 285.2 nm. The unprotected bare Mg and the protected ones, fluorinated, phosphated, as well as those both chemically treated and PLA-coated, are placed in Kokubo’s SBF; the corrosion is continued for 1 week. After a period of corrosion time from 5 min to 150 h, 2 mL of the solution is removed and diluted for the atomic absorption measurement. The concentration of Mg2+ ions in the samples solution is measured, and the corrosion rate of the unprotected or protected metal samples can be deduced. The corrosion degree has been characterized with electrochemical analysis at the electrochemical workstation. The electrolyte solution of 0.9% NaCl has been used; a three-electrode system has been established, with the platinum electrode as the auxiliary electrode, the saturated calomel electrode as the reference electrode, and the metal or surface-protected magnesium as the working electrode. The surface area of the working electrode is 1 × 1 cm2; the nonworking part has been coated with epoxy resin. From the Tafel polarization curve, the polarization potential (E) and the corrosion current density (i) of the metal have been recorded at a potential scanning speed of 1 mV/min. In Vivo Animal Study. The animal experiments were approved by the Institutional Animal Ethics Committee of Nanjing University according to the Animal Care Guidelines for the Care and Use of Laboratory Animals in China. The metal wire with the diameter of 1.5 mm has been cut into 2.5 mm lengths, polished clean with both ends polished smooth. The prepared tiny rods have been sterilized with

75% ethanol and irradiated under ultraviolet, and they were then ready for implantation. Sprague−Dawley rats of 5−6 weeks weighing 140−160 g were used in the implantation for bone defect repair and further study. The rats have been anaesthetized via intraperitoneal injection of chloral hydrate (0.3 mL/100 g). After exposure of the plate 5 mm below tuberositas tibiae via a medial approach, a hole of 1.2 mm in diameter and 0.8 mm in depth was created by using a hand drill in both tibias, and the metal pins were implanted. There have been 68 rats used and randomly grouped into four groups, with 17 rats in each group. Rats in group I have been implanted with titanium as a control. Those in group II are for Mg implantation, with those in group III for phosphate treatment and PLA-coated Mg (noted as mMg-P+PLA), and those in group IV for the fluoride treatment and PLA-coated Mg (noted as mMg-F+PLA). The thicknesses of the PLA coating films in groups III and IV are about 70 μm. In each group, rats are sacrificed at 0, 2, 4, 8, and 12 weeks, respectively, after the surgery. X-ray Imaging. X-ray imaging has been used to monitor implant degradation and defect healing. All samples receive X-rays by MX-20 (Faxitron, USA) immediately after dissection. Micro-CT Analysis. Micro-CT (Skyscan 1176, Kontich, Belgium) has been used to monitor the degradation rate of implants and new bone formation surrounding the implants. The samples have been scanned by using a voxel size of 12 μm, with a source voltage of 65 kV and a source current of 385 μA. After reconstruction, the images are reoriented by using the Skyscan Data Viewer software and the region of interest (ROI) is analyzed with the Skyscan CTAn software. Threedimensional reconstruction is performed by using Skyscan CTvox software. According to the absorption coefficient, the metal material of magnesium can be segmented from the surrounding new bone. We have selected the threshold range from 85 to 115 as Mg implants and another threshold range from 120 to 255 as new bones. Morphometric parameters include total TV, BV, and BV/TV. As for the quantitative analysis, 41 slides of 2D images have been used for the surrounding bone formation in the bone marrow without contacting the cortical bone and 166 slides for implant degradation. Histological Analysis. Specific fluorescence bone markers have been administered to evaluate the amount of new bone surrounding the implant. The animals are injected subcutaneously with calcein green (Sigma; 10 mg/kg) 2 weeks and with xylenol orange (Sigma; 90 mg/kg) 1 week, respectively, before the rats are sacrificed. Tibias have been dehydrated in an ethanol gradient series (70, 85, 90, and 100), embedded in MMA without decalcification, and cut into sections by using a diamond bandsaw (Exakt, Norderstedt, Germany). The tissue sections are polished to 50 μm to evaluate the new bone formation rate under a confocal microscope (Leica TCS SP5, Leica, Germany). Statistical Analysis. Data are presented as means ± standard deviation. The unpaired two-tailed Student’s test has been used for comparison between two groups. One-way ANOVA with relevant post hoc tests was used for multiple-group comparisons. The value of P < 0.05 is considered to be of statistical significance and is denoted as “*”, and P < 0.01 is “**”. GraphPad Prism software (Version 6.01) is used for the above statistical analysis.

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RESULTS AND DISCUSSION The appearances of the metal rods before and after chemical conversion treatment have been shown in Figure 2. The surface of the metal rods after polishing and cleaning is smooth and with a silver-white metallic luster. The metal sample treated with 16% hydrofluoric acid also has a shiny silver-like surface, while the phosphate-treated sample is golden colored. The results of the chemically treated and PLA-coated Mg samples, namely, mMg-F+PLA and mMg-P+PLA, are shown in Figure 2e,f. The resulting color of the treated metals depends on treatment time and temperature; e.g., if the sample is treated with 40% hydrofluoric acid, the metal color becomes brown, while if the sample is phosphate-treated at a higher temperature, it changes to black. It changes with alloy C


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ACS Applied Bio Materials

Figure 4. Results of the grid tests for PLA-coated metals of Mg of (a) untreated, (b) phosphate-treated, and (c) fluoride-treated.

Table 1. Number of Coated PLA Grids among 100 Grids Left on the Substrate of (a) Untreated Metal, (b) phosphate-Treated, and (c) Fluoride-Treated Metals

Figure 2. Results of the metal rods: (a) unpolished, (b) polished and cleaned, (c) fluoride-treated, (d) phosphate-modified, (e) fluoridetreated and dip-coated with PLA films, and (f) phosphate-modified and dip-coated with PLA films.

(a) Mg (b) mMg-P (c) mMg-F

elements, too: the 16% HF fluoride-treated WE43 is brown. The PLA film itself is transparent. The coated PLA films are transparent, with a thickness of about 70 μm, which has been estimated with the weight difference before and after the PLA coating, the surface area of the coating, and the density of the PLA (1.25 g/cm3). For example, the rod of 1.3 mm in diameter and 50 mm in length weighed 0.1095 and 0.1277 g, respectively, before and after PLA coating, so the deduced thickness of the PL coating is (0.1277 g − 0.1095 g)/1.25 g/cm3/(3.14 × 0.13 cm × 5 cm) = 0.0071 cm = 71 μm. Figure 3 shows the SEM observations of the metal surface before and after chemical fluoride or phosphate treatment. It can be observed that (a) the polished metal surface is smooth, (b) the fluoride-treated surface shows some uniform tiny particles, and (c) the phosphate coating has given the metal surface some scaly matter grown on the substrate. To evaluate the adhesion between the PLA and the metal substrate, the grid tests have been performed following the standard method of GB/T 9286-1998 or ISO 2409-1992. The 10 × 10 grids have been cut through the film with each grid having a size of 1 × 1 mm2; a tape has been pasted on the cut film and then peeled off in the direction of the cutting lines. Some film grids are peeled off with the tape. Some are left on the substrate, and the results are shown in Figure 4. The numbers of the film grids peeled off or left have been counted, values which are listed in Table 1. The grids have been painted red before tape pasting so that the grid numbers can be clearly counted. Parallel samples have been tested so that the errors have been derived.

sample 1

sample 2

sample 3

average

43 45 88

50 48 83

45 56 86

46 ± 4 50 ± 5 85 ± 3

The unprotected or protected Mg rods with 1.2 mm diameter and 70 mm length have been put into the Kokubo’s SBF for different periods of time. After careful removal of the oxide layer on the metal and after the sample was dried, the PLA was dissolved in chloroform first if the sample was PLAcoated, the weight of the metal rod was weighed, and the mass change per unit area per day of the sample before and after corrosion was deduced and termed as the average corrosion rate, ν (mg/cm2/day), which is defined as follows. W − W1 ν= 0 (1) At Here, the following abbreviations apply: A is the surface area of the specimen (cm2), t is the corrosion time (day), W0 is the original weight of the specimen (mg), and W1 is the weight after the corrosion and the corrosion product has been removed (mg). It can be observed from the weight loss rate shown in Figure 5a that the corrosion resistance of the magnesium metal sample after chemical conversion treatment and PLA coating is improved, the weight loss rate of the fluorinated treatment and PLA coating (mMg-F+PLA) is the lowest, and the corrosion resistance is the best. During the fluoride treatment, we have changed the concentration of the HF from 10 to 40 wt %; the two blue lines in Figure 5a are the corrosion rate of Mg treated with 10% and 40% HF, respectively, and the gray area is the error region. The results shown in the lower blue line, indicating better anticorrosion, are for mMg-F treated with 16 wt % HF, so the concentration of 16 wt % has been adopted for further fluoride treatment.

Figure 3. Microscopic observations with SEM of the metal surfaces: (a) untreated metal, (b) fluoride-treated, and (c) phosphate-modified. D


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Figure 5. Corrosion rates recorded for (a) the weight loss and (b) the release of Mg2+ into the SBF of unprotected bare Mg, fluoride-treated Mg (mMg-F), PLA-coated Mg (Mg+PLA), and both fluorated and PLA-coated Mg (mMg-F+PLA).

Figure 6. (a) Corrosion rates of PLA-coated Mg (Mg+PLA) with different coating thicknesses, recorded from weight loss in SBF in 1 week. (b) The Young’s modulus of the metal rods of unprotected Mg and protected Mg (mMg-F+PLA) after corrosion in the SBF for 1−4 weeks.

Figure 7. Tafel curves recorded for (a) untreated Mg metal and metals with phosphate and fluoride treatment, and (b) the untreated and treated metals with PLA coating.

the unprotected pristine Mg, which is due to the breakage of the PLA coating film and the metal substrate contacts with the solution, accelerating the corrosion rate of the metal. In our observation in the in vitro corrosion experiment, the coated PLA is almost peeled off from the metal surface in the first week. In both graphs of weight loss and Mg release, the mMgF+PLA sample shows excellent anticorrosion, and the corrosion rate has been reduced to 1/10 or 1/20 that of pristine Mg in the first week or one month.

Figure 5b shows the corrosion rate determined from the atomic adsorption measurement; the total amount of Mg2+ ions released into the SBF per unit surface area of Mg+PLA, mMg-F, and mMg-F+PLA, respectively, has been tested. In the first 30 h the corrosion rate of Mg+PLA is lower than that of the mMg-F, which is due to protection of PLA on the metal surface. After that, the corrosion rate of Mg+PLA accelerates abruptly, until at the end of the first week the corrosion rate of the Mg sample with PLA coating is almost as rapid as that of E


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The thicker the coated PLA film is, the slower the corrosion is. We would not coat the PLA thicker than 100 μm, because the thick coating film is easily damaged by the metal forceps in the surgery. The mechanical decay of the unprotected and protected (mMg-F+PLA) magnesium rods in SBF with corrosion time from 0 to 4 weeks is demonstrated in Figure 6b. The Young’s moduli have been obtained from the threepoint test of the rods. The mechanical decay of the protected Mg is slower than that of the unprotected sample, and at the end of the 28 days, the Young’s modulus of the unprotected Mg has dropped lower than 10 GPa while that of the protected sample is maintained above 20 GPa. In summary, the fluoride treatment and the PLA coating have effectively retarded both the weight loss and the mechanical decay of magnesium. To explore the anticorrosion mechanism of chemically treated or/and PLA-coated magnesium, the corrosion rate has been evaluated with the corrosion current density, the corrosion potential, and the corrosion depth in an electrochemical test.20−22 Figure 7a shows the relation between the corrosion potential and corrosion current for the untreated magnesium (Mg) and fluoride-treated (mMg-F) and phosphate-treated (mMg-P) magnesium, while in Figure 7b such a relation is shown for the PLA-coated magnesium (Mg+PLA) and both PLA-coated and fluoride or phosphate-treated magnesium (mMg-P+PLA, mMg-F+PLA). The amount of metal uniformly corroded from an anode or electroplated on a cathode in an aqueous solution in a time period can be determined by using Faraday’s equation from general chemistry.

Table 2. Corrosion Penetrating Rate of the Unprotected and Protected Specimens

Mg mMg-P mMg-F Mg+PLA mMg-P+PLA mMg-F+PLA

CPR (mm/a) from weight loss

CPR (mm/a) from current density

icorr (μA/cm2)

2.5 0.58 0.42 0.69 0.44 0.21

0.30 0.06 0.01

13.2 2.5 0.40 4.0 × 10−4 1.2 × 10−4 1.2 × 10−4

w It = M nF

(2)

Here, the following abbreviations apply: w is the weight corroded or weight loss (g) of the metal, t is the corrosion time (s), I is the current flow (A), M is the atomic mass of the metal (g/mol), and n is the number of eletrons produced or consumed in the process per atom. F is the Faraday constant (96 500 C/mol, 96 500 A s/mol). So, the corrosion rate expressed by the weight loss can be determined as in eq 3. ν= Figure 8. (a) Metal pins used in the rat tibiae defect model, Ti, Mg, mMg-P+PLA, and mMg-F+PLA as shown from top to bottom. (b) Xray images of pins implanted in rat bone tibia defects, with an enlarged image of a local part. (c) X-ray observations of the tibia defects and the implants 0, 2, 4, 8, and 12 weeks after the surgery.

Micorr w ItM 1 = = At nF At nF

(3) 2

Here, icorr is the current density (A/cm ), directly associated with the electrochemical corrosion reactions, so eq 2 indicates that the corrosion rate is proportional to corrosion current density. It can be observed from the Tafel diagram shown in Figure 7a that the corrosion current of the fluoride-treated Mg (mMg-F) or phosphate-treated (mMg-P) is 1 order of magnitude lower than that of Mg pristine metal. The corrosion potential of the mMg-F is higher than that of Mg, indicating that mMg-F possesses a higher corrosion resistance than that of the pristine metal in both thermodynamics and kinetics. However, the corrosion potential of mMg-P is smaller than that of Mg, but the actual corrosion current is smaller than that of Mg because the corrosion current density is the key factor that actually affects the corrosion rate, as we have obtained from eq 3. Figure 7b demonstrates the Tafel diagram of the PLA-coated metals Mg, mMg-F, and mMg-P, noted as Mg+PLA, mMg-F +PLA, and mMg-P+PLA, respectively. It can be observed that the mMg-F+PLA and mMg-P+PLA samples show lower corrosion currents and higher corrosion potential than those of Mg+PLA, indicating that the multilayer-protected samples

On one hand, the fluoride treatment of the chemical conversion coating itself enhances the corrosion resistance of Mg; even if the PLA membrane is damaged, the fluoride conversion film can also achieve effective protection for the magnesium matrix within a certain period of time. On the other hand, the adhesion between the fluoride coating and PLA is largely enhanced too, so that the PLA film is not peeled off from the metal substrate so easily and so soon. The PLA thickness dependent anticorrosion property has also been evaluated; the result is shown in Figure 6a. Four batches of magnesium rods have been tested, with the coated PLA thicknesses averaged as 0 (without coating), 50, 75, and 92 μm, respectively; the corrosion rates of 1 week are reduced from 1.1 mg/cm2/day for the unprotected sample to 0.1 mg/ cm2/day for a sample with a coating thickness above 70 μm. F


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Figure 9. Materials biodegradation and new surrounding bone formation in rat tibia defect evaluated by using micro-CT. (a) Three views of the tibia defects implanted with Ti, Mg, mMg-P+PLA, and mMg-F+PLA. (b) 3D reconstruction images for the implanted Mg (red), mMg-P+PLA (blue), and mMg-F+PLA (purple) and the new bone growth (white) surrounding the implants. (c) The changes of material volumes at 2, 4, 8, and 12 weeks after surgery. (d) BV and (e) BV/TV values of the surrounding bones. *p < 0.05, **p < 0.01, n = 6.

than 0.1 mm/a for the clinical application. Though the corrosion rate in vivo is usually lower than that in vitro, the mechanical decay of the metal might be too fast to meet the clinical requirement. To further investigate the in vivo corrosion resistance of the multilayer-protected metal of magnesium, to prove whether the modified materials can be used as biomaterial implants, the in vivo animal test for the rat bone defect repair has been performed. To determine the bone regenerative capacity, we have implanted four materials including Ti, bare Mg, mMg-P+PLA, and mMg-F+PLA, as shown in Figure 8a from top to bottom, respectively, into rat bone tibia defects, as shown in Figure 8b. After the implantation, the rats have been sacrificed at 0, 2, 4, 8, and 12 weeks; the operated bone with the material is taken out for the material degradation and bone growth evaluation with the X-ray and micro-CT scan. Images of the X-ray are shown in Figure 8c, and we can observe the degradation of the material at the position of the tibial defect. As the image brightness is associated with the density of the metal, the image from titanium is much brighter than that of magnesium and

have greater anticorrosion properties, while between the two multilayer-protected specimens of mMg-F+PLA and mMg-P +PLA, the former possesses a better anticorrosion property. We have another term, corrosion penetrating rate (CPR), defined as the thickness loss of the metal corroded per unit of time, to express the corrosion rate. CPR =

Micorr w = ρAt ρnF

(4)

Here, the following abbreviations apply: w is the weight loss (g), t is the corrosion time (s), and A is the exposed area of the specimen (cm2), if we use the unit μA/cm2 for icorr, and g/cm3 for ρ, the metal density. The unit of CRP shall be millimeters per annual (mm/a). The corrosion penetrating rate of the untreated and the chemically treated Mg has been listed in Table 2, from the corrosion current density from Figure 7 and eqs 2−4. It is known from the reported work23,24 that the corrosion penetrating rate of the biomedical Mg or its alloy shall be lower G


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Figure 10. Sequential fluorescent labeling to determine the new bone formation. (a) Calcein green/xylenol orange assessment of bone remodeling. Scale bars are 200 μm. (b) The mineral apposition rate (MAR) results of Ti, Mg, mMg-P+PLA, and mMg-F+PLA samples on week 8, respectively. *p < 0.05, **p < 0.01, n = 6.

(BV/TV ratios) of the protected mMg-P+PLA and mMg-F +PLA are significantly greater than that of the bare Mg, as shown in Figure 9e. Fluorochrome labeling is a widely used standard technique in skeletal research, which is simple and efficient for the investigation of the dynamics of bone remodeling in combination with plain histology, allowing the determination of the onset time and location of mineralization and the direction and rate of bone formation.25,26 In our work, calcein green and xylenol orange combine with the calcium to label in situ new bone deposition and bone resorption sequentially. The results have been shown in Figure 10a, and the mineral apposition rate (MAR) has been deduced from the distance of the labeled bone formation per day, and shown in Figure 10b. It has been observed in Figure 10 that, in rat tibia over a period of 8 weeks after surgery, MAR significantly increases to 24 μm/ day in mMg-P+PLA and 28 μm/day in mMg-F+PLA, while in the control group it is 3 μm/day in Ti and 6 μm/day in bare Mg, indicating a higher osteoblast activity around the mMg-P +PLA and mMg-F+PLA implant. In addition to the observations from X-ray and micro-CT that the implants of multilayer-protected Mg show a low corrosion rate, the samples also exhibit significant new bone formation after implantation in the rat tibia defect. The new bone formation effect is not supposed to be due to PLA, since PLA has minimal calcification and occasionally has no bone in-growth effect in a long-term follow-up survey.27,28 In previous studies, elemental Mg is crucial to bone formation and metabolism, and the doping of Mg could improve the differentiation of MSCs into osteoblasts.29,30 It is also shown that osteoclast differentiation is inhibited in a conditional medium enriched by Mg degradation products. 31,32 Therefore, only an appropriate amount of Mg2+ release from the degradation of

bone tissue, and the brightnesses of magnesium metal and the bone tissue are similar. The unprotected bare magnesium material could not maintain its integrity after 4 weeks; meanwhile, new bone formation and bone resorption occur due to the release of magnesium ions. On the contrary, we have observed the low degradation rate of and a well-shaped new bone formation around the protected magnesium materials. These observations have been further confirmed and investigated by micro-CT. On the basis of the different material density, volumes of metal and new bone around the implant can be quantitatively separately. Figure 9a shows the sagittal, coronary, and horizontal planes analyzed by special reconstructive software of the bone defect with the materials. In Figure 9a the degradation of the bare Mg implant is apparent at week 4; it continues until week 8 and becomes worse before sacrifice. It is observed in Figure 9a that not only did Ti retain its initial shape all the time, but also mMg-P+PLA and mMg-F+PLA make the implant well-protected until week 12. The bare Mg volume is reduced by 15%, 28%, and 38%, respectively, after 4, 8, and 12 weeks of implantation, whereas the volume of mMg-P+PLA is reduced by 7%, 8%, and 19%, and that of mMg-F+PLA is reduced by 2%, 4%, and 11% after 4, 8, and 12 weeks of implantation, as shown in Figure 9c. Moreover, direct contact of the Mg materials with the progressively enhanced surrounding new bone has been observed in Figure 9b, whereas bone resorption has been observed around the bare Mg, which might due to the fast corrosion rate. Micro-CT analysis in Figure 9d reveals that the new bone volume (BV) around mMg-P+PLA was 0.32 ± 0.04 mm3 and that around mMg-F+PLA was 0.53 ± 0.08 mm3, values which are greater than those of 0.23 ± 0.01 mm3 surrounding bare Mg at 8 weeks after surgery. Similar with the tendency of BV, the growths of the bone around the material H


Article


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Mg materials, such as in the case of protected Mg, provides a beneficial effect on the bone growth and osteopenia.

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CONCLUSIONS To overcome the drawbacks of fast degradation of the biomedical Mg implants in vivo, we have provided an available solution for retarded and controlled degradation of Mg by a novel multilayered surface coating on the Mg metal, which comprises a chemical conversion layer, an adhesive layer, and the biodegradable polymer film of PLA. The modified surface of Mg via chemical conversion demonstrates enhanced adhesion with PLA as compared to that of untreated Mg, and the multilayer coating provides the Mg substrate with adequate biodegradability and suitable mechanical properties. The corrosion rate of the protected Mg can be adjusted with either the chemical conversion type or the thickness of the PLA layer. In the in vivo study of rats’ tibia defect repair, the protected Mg possesses an improved anticorrosion property. The X-ray, micro-CT observations and histological analysis show that the protected Mg not only biodegrades slowly but also significantly promotes the new bone formation at the surroundings, so the healing time of the bone defect can be shortened. The innovated coating shows great potential as a good partner for biomedical implants or surgery devices of Mg or Mg alloy, making them more advanced in biodegradability and bioactivity in clinical applications than the traditional metal implants such as Ti and stainless steel.

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected] (D. Z.). *E-mail: [email protected] (Y. Z.). ORCID

Jian Shen: 0000-0002-9009-6151 Chi Wu: 0000-0002-5606-4789 Dan Zhu: 0000-0002-9834-7752 Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The financial support of the National Natural Science Foundation of China (51273091, 21204037, 50773077, 81772402, and 20934005), the National Key Technology Support Program (2015BAI08B02), the Hong Kong Special Administration Region Earmarked Projects (CUHK4042/09P, 2160396), Excellent Young Scholars NSFC (81622033), the Central University Basic Scientific Research Funding (021414380339), Science and Technology Projects of Jiangsu Province (BK20160633), and the Scientific Research Foundation of Graduate School of Nanjing University (2016CL12) is gratefully acknowledged.

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Biodegradable and Bioactive Orthopedic Magnesium Implants with

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