Abstract
Background/Aim: Hydroxyapatite (HA) coating can improve the degradation rate and biological activity of metallic implants. This study aimed to fabricate a hydroxyapatite-coated ultrafine-grained biodegradable WE43 magnesium (HA/UFG–WE43 Mg) implant for repairing bone fractures. Materials and Methods: A hybrid approach, including parallel tubular-channel angular pressing (PTCAP) and physical vapour deposition (PVD) magnetron sputtering, was employed. The HA/UFG–WE43 Mg samples were tested in terms of their physicochemical and biological properties. Results: The processed tubes exhibited ultrafine structures and the uniformity of microstructures improved following the two-pass PTCAP. The phase composition of the coating formed on UFG–WE43 Mg implant at 250 W for 90 min after heat treatment at 500°C for 60 min confirmed the presence of the HA characteristic peaks. Rat skeletal muscle cells were inoculated on the specimens and cultured for 1, 2, 6, 12, and 24 h, followed by evaluation of cell adhesion and morphology. The growth rates of cells were examined by the Cell Counting Kit8 (CCK-8) and cell survival was observed after 3 days of culture by fluorescence microscopy. The concentration of Mg ions in the blood of rats on 1, 3, 5, 7, and 15 days showed a reduction in Mg concentration after deposition of HA. Conclusion: Combination of PTCAP processing followed by surface modification led to tibial fracture healing, and histological analysis of implanted areas demonstrated an efficient biodegradation of the implanted material and a moderate inflammatory reaction.
- Hydroxyapatite
- ultra-fine grain WE43
- surface modification
- biomineralization
- in vitro and in vivo analysis
Out of all different biodegradable materials, Mg-based alloys have attracted much attention for biomedical applications, such as in the musculoskeletal field due to their excellent biodegradability, biocompatibility, and mechanical properties that are similar to the human bone (1, 2). Among them, the WE43 Mg alloy represents a promising material for fabrication of biodegradable cardiovascular stents (3). The WE43 Mg alloy also showed good biocompatibility in vivo and adequate mechanical strength in a canine fracture model (4). Hence, developing a more detailed understanding of the physicochemical properties and biological behaviour of the WE43 alloy may help extend its clinical applications (5).
In general, the application of Mg-based alloys is constrained in metal forming procedures as a consequence of the inferior plasticity at ambient temperature attributable to the limited number of slip planes in a hexagonal close-packed (hcp) lattice, and the lack of comprehensive information on the forming mechanisms at elevated temperatures (6, 7). Accordingly, extensive efforts are devoted to improving the microstructural properties and mechanical performance of Mg-based alloys (8-10). In particular, comprehensive investigations are conducted to overcome some structural weaknesses and simultaneously improve the mechanical strength and ductility behaviour of Mg-based alloys under various severe plastic deformation (SPD) methods, such as high-pressure torsion (HPT) (11), accumulative roll bonding (ARB) (12), multiaxial forging (MAF) (13), and equal channel angular pressing (ECAP) (14), as well as continuous SPD techniques (15). These methods are commonly utilized for bulk shape manufacturing. In addition to bulk components, the tubular and parallel tubular-channel angular pressing (TCAP/PTCAP) processes are at the forefront of an effective production strategy for UFG and nanostructured tubular components (16, 17). The PTCAP method as an emerging SPD method, in which the deformation takes place in two-half cycles, is affected by some parameters, such as channel and curvature angles, as well as deformation ratio (18, 19). Based on the literature, the lowest possible indispensable process load and the best strain consistency can be obtained at an angle of curvature equivalent to zero. It is also found that lower deformation ratio gives rise to the best strain consistency and lowest possible process load (20). To avoid cracking, the PTCAP process is usually performed at greater than or equal to 423 K due to the limited ductility of Mg at low temperatures (21, 22). However, coarsening and low plasticity may dominate at high (23) and low (24) temperatures, respectively, which are the major factors of defect formation during SPD processes. So, choosing the right temperature for the PTCAP process is of particular importance. On the other hand, Mg-based alloys are potential materials for degradable implant applications owing to their biocompatibility and biodegradability; nonetheless, the main concern is their high degradation rate in body fluid, which can be overcome using bioceramic coatings, e.g., HA (25, 26). HA-based coatings can diminish and improve the degradation rate and biological activity of Mg alloys, thereby possessing the capability to induce and conduct bone formation (27-29). In this context, the commonly used fabrication methods predominantly include electrochemical processes, spraying methods, micro-arc oxidation, hydrothermal process, and the sol-gel method (26).
Although extensive studies are available in the literature on the plasticity behaviour, mechanical properties, and in vitro/in vivo performance of the processed Mg-based alloys (27-29), there is no study on PTCAP processing of WE43 Mg alloy, as well as its surface modifications by bioceramic coatings. The present study described the fabrication of UFG–WE43 Mg implant by PTCAP, followed by its surface modification through HA PVD coating. The microstructural characteristics and biological behaviour of the system were investigated in vitro and in vivo in rats to reach equilibrium between UFG–WE43 Mg implant degradation and bone healing.
Materials and Methods
Materials. Cylinder samples with the thickness, outer diameter, and length of 2.5, 20, and 40 mm, respectively, were cut from an extruded WE43 Mg alloy bar.
PTCAP process. For the PTCAP operation, the die was fabricated from tool steel and hardened to 55 Hardness Rockwell C (HRC). The features of the PTCAP die are listed in Table I. One-, two-, and three-pass PTCAP was performed using a press machine at 300°C with a ram speed of 5 mm/min. To diminish the friction between the workpiece and the tool steel die, molybdenum disulfide (MoS2) was employed as an inorganic lubricant. The chemical composition and specifications of the WE43 extruded bar are presented in Table II.
Figure 1A displays the schematic diagram of the PTCAP procedure including preliminary status, first half-cycle, and second half-cycle along with the die parameters. The total cumulated plastic strain () caused by the shear and normal strains following N-pass PTCAP can be determined using the die geometry and the following equation (18):
I
where φi, Ψi and are channel angle, curvature angle, and , respectively. εθ is equal to ln R2/R1. Consistent with this equation, PTCAP processing is governed by an equivalent plastic strain of 1.8 following each pass. Thus, the cumulated strain following two- and three-pass PTCAP is 3.6 and 5.4, respectively. As shown in Figure 1A, in this two-part process, the workpiece is placed between the mandrel and die at the start of the process. Following, through pressing the first punch, the specimen moves into the deformation region consisting of a tubular groove with two shear sections. The sample is then pushed back into the same shear region through the second punch to regain its initial size. In such metal forming operation, high equivalent plastic strain is generated in the shear area thanks to the cumulative shear strain in each pass. It should be mentioned that tensile and compression circumferential strains are obtained all over the first and second half-cycles of PTCAP, respectively. Figure 1B shows the experimental output of UFG–WE43 Mg produced by three passes of PTCAP.
HA PVD coating. After PTCAP processing, the HA coating was sputtered on UFG–WE43 Mg by a PVD magnetron sputtering machine (SG Control Engineering Pte Ltd) equipped with a pure HA target (99.99%). Radio-frequency (RF) power magnetron sputtering was utilized to deposit the HA coating, wherein the HA target was connected to the RF terminal (Figure 1C). The distance between the HA target and sample was kept constant during the surface modification. Before deposition, UFG–WE43 Mg was rinsed with acetone and preheated at 300°C for 30 min to improve the adhesion of the HA coating. The chamber of deposition was evacuated to 1.9×10−6 Torr before applying the Ar gas for sputtering, whereas a working pressure of 2.75×10−3 to 3.15×10−3 Torr was preserved throughout sputtering with a constant 20 sccm Ar gas flow (Table III). To improve the crystallinity of the HA coating, heat treatment was executed at 500°C for 60 min.
The schematic of the HA/UFG–WE43 Mg implant with length and width of 10 and 1 mm, respectively, and the tibia fracture model in the rat, as well as the relevant characterization techniques are illustrated in Figure 1D. The proximal and distal parts of the tibia along with the fracture line and the proposed biodegradable implant are shown in Figure 1.
Microstructural characterization. The PTCAPed samples were cut in a perpendicular direction to their extrusion direction, followed by grinding with SiC paper until 2400-grit and polishing with diamond paste. For the final polish, an oxide polishing suspension was used to achieve a stress-free surface. Here, to observe the grain boundaries, 4.2 g picric acid moistened with water (C6H3N3O7, ≥98%, Sigma-Aldrich, St. Louis, MO, USA), 50 ml ethanol (C2H5OH, 95%, Sigma-Aldrich), 20 ml of distilled water, and 10 ml of acetic acid (CH3COOH, ≥99.8%, Sigma-Aldrich) were used as an etchant. For the microscopic examination, a Phenom ProX desktop scanning electron microscope (SEM) coupled with energy-dispersive X-ray spectroscopy (EDS) was utilized. Here, grazing incidence X-ray diffraction (GIXRD) helps distinguish the HA layer signal from UFG–WE43 Mg substrate. The GIXRD profiles were evaluated using “HighScore Plus” software using JCPDS card number #24-0033 for HA.
In vitro bioactivity. PTCAPed specimens before and after HA surface modification (10×2.5×1 mm) were cut and ultrasonicated for 30 min and dried for 1 h at 90°C. Then, the in vitro bioactivity was examined by immersing the samples in simulated body fluid (SBF), in which the SBF volume ratio to the sample surface area was designed as 1:10. It should be noted that the SBF concentration resembled that of human extracellular fluid based on the preparation approach proposed by Kokubo et al. (30). The immersed samples were placed in a constant temperature water bath at 37°C for different intervals of time, then the specimens were taken out and rinsed with de-ionized water and dried for 1 h at 90 °C. They were then exposed to Phenom ProX desktop Scanning electron microscopy/energy dispersive X-ray spectrometry (SEM/EDS) analysis.
Cell culture and cell adhesion. The L6 cells (derived from rat skeletal muscle) were purchased from the American Type Culture Collection (Rockville, MD, USA) and preserved in complete Dulbecco’s modified Eagle’s medium (DMEM) with 10% fetal bovine serum (FBS), 100 U/ml antibiotic (penicillin), as well as 100 μg/ml streptomycin. The cells were incubated in a humidified atmosphere comprising 95% air and 5% CO2 at 37°C. L6 cells were inoculated on UFG–WE43 and HA/UFG–WE43 Mg implants and cultured for 1, 2, 6, 12, and 24 h. The adhesive strength of rat skeletal muscle cells attached to the surfaces of untreated (WE43 Mg) and treated samples (i.e., UFG–WE43 and HA/UFG–WE43 Mg implants) was appraised through the percentage of residual cells following mechanical separation (31). For this purpose, cells incubated on the untreated and treated samples for 1, 2, 6, 12, and 24 h were washed with phosphate-buffered saline (PBS) to eliminate non-adherent cells and subsequently separated from the surface through vibrating a culture dish at 37°C for 5 min. The remaining cells were measured using an enzyme-linked immunosorbent assay reader and the cell adhesion morphology was observed by SEM.
Assessment of cell proliferation. The proliferative activity of the leaching solution from UFG–WE43 Mg and HA/UFG–WE43 Mg implants on L6 cells was quantitatively measured by the CCK 8 kit (Dojindo, Kumamoto, Japan). The rat skeletal muscle cells were cultured in 96 well plates at 5×103 cells/well in triplicate. Following 1, 3, 5, and 7 days of culture, 90 and 10 μl of culture medium and CCK 8 solution, respectively, were added to each well at each time point and incubated at 37°C for another 4 h. At the same time, the cells without alloy extract were used as controls. The optical density (OD) was measured using a Tecan’s Sunrise absorbance microplate reader at 450 nm.
Cell morphology. The morphology of cells was observed through fluorescence microscopy executed after 3 days of the co-incubation of the rat skeletal muscle cells with the untreated and treated specimens under the aforementioned circumstances. To this aim, the cells formerly rinsed with Sigma-Aldrich Hanks’ Balanced Salt Solution (HBSS) without red phenol were treated with 200 μl/well of a 1:1,000 Sigma-Aldrich Calcein AM solution in HBSS. Then, cells were incubated in a dark room at 37°C for 30 min to ensure the intracellular alteration of the cell-permeant dye to a green-fluorescent calcein inside of the viable cells. After that, the cellular features, i.e., morphology and distribution, were acquired using a Leica DMI4000B inverted microscope (Leica, Wetzlar, Germany).
Testing in an animal setting. The cylindrical UFG–WE43 Mg and HA/UFG–WE43 Mg implants with a diameter of 1 mm and a length of 10 mm were used in this study, where the relevant dimensions were calculated based on a certain orthopaedic algorithm. Since the animal weight is directly proportional to the long bone size, the length of the UFG–WE43 Mg and HA/UFG–WE43 Mg implants was set at one-third of the tibia length, and the thickness of the implant was designed concerning the fact that it must not exceed 60% of the bone diameter. All animal experiments were conducted under animal ethical respect and were authorized by Zhongke Inno International Medical Research Institute (Selangor, Malaysia). Forty male Sprague-Dawley (SD)-rats aged 3 months were randomized into four groups with 10 rats in each group as follows:
– G1: Normal control group;
– G2: Tibia fracture model group;
– G3: Tibia fracture model group + UFG–WE43 Mg implant;
– G4: Tibia fracture model group + HA/UFG–WE43 Mg implant
Preoperatively, the rats were sedated by inhalation of 1.5% isoflurane and oxygen. The surgical site was shaved and then disinfected with betadine solution. By locating the rats on the operating table, the UFG–WE43 Mg and HA/UFG–WE43 Mg implants were placed using an incision, and then blunt dissection was performed to reach the bone surfaces. Subsequently, the muscular layer and skin were sutured and closed. Postoperatively, the SD rats were kept in individual cages, and conventional X-ray imaging was performed 15 and 30 days after surgery to scrutinize the position of the implants and the healing of the fracture. Besides, the tissue response was assessed histologically using Masson’s trichrome staining. Blood biochemical testing was also performed on 1, 3, 5, 7, and 15 days after the operation to estimate concentrations of Mg and Ca, as the main elements of the substrate and coating, respectively.
Statistical analysis. Data are expressed as mean±standard deviation. Statistical analysis was performed with the SPSS 10.0 software (SPSS Inc., Chicago, IL, USA) and differences between groups were analysed using a one-way analysis of variance, followed by the Tukey test. p<0.05 was considered to indicate a statistically significant difference.
Results
Surface modification and microstructural features. Bioceramic coatings, such as HA and HA-based composite coatings, can increase the osseointegration degree between bone tissue and implant (32). Herein, to develop an HA coating on the UFG–WE43 Mg implant, the specimen with the optimum structural and mechanical properties (3 passes PTCAP) was chosen for further processing using PVD magnetron sputtering. The reason for this choice was that the microstructure and mechanical behaviour of the substrate can affect the biological performance of the system (33). Figure 2 shows the microstructure, chemical composition, and phase compositions of the HA layer formed on UFG–WE43 Mg implant at 250 W for 90 min after heat treatment at 500 °C for 1 h. From the top-view SEM images in Figure 2A and B, a stabilized homogeneous porous structure was formed after HA surface modification. It revealed that the coating has a uniform thickness.
From the EDS analysis (Figure 2C), the HA coating deposited at 250 W for 90 min had a Ca/P ratio of around 1.68, which is close to the stoichiometric ratio (1.67) that exists in natural bone tissue (34). Since the faster dissolution of HA creates a supersaturated medium and enables the precipitation of physiological HA on the coating, enhances the bone ingrowth; however, it causes faster resorption or degradation of the coating layer (35). The GIXRD profile of the HA coating formed on UFG–WE43 Mg implant at 250 W for 90 min after heat treatment at 500°C for 60 min is illustrated in Figure 2D. This profile confirms the presence of the HA characteristic peaks (JCPDS#24-0033), including (0 0 2), (2 1 0), (2 1 1), (1 1 2), (3 0 0), (2 0 2), (1 3 0), (2 2 2), (2 1 3), (0 0 4), and (3 2 2) planes.
Biomineralization. Figure 3 displays the SEM micrographs of the specimens immersed in SBF solution for 7-14 days. A coarse equiaxed grains structure with an average size of ~120 μm along with several irregular surface cracks and some white heterogeneous deposits were observed on the unprocessed alloy after 7 days of immersion (Figure 3A). This image exhibits the α-Mg matrix, the precipitation of some intermetallic compounds (IMCs), like Mg24Nd5 and Mg41Nd5 phases in both grains’ interior and grain boundaries, and some HA heterogeneous deposits (36). Following 14 days of soaking, the fraction of the white deposits increased, but the entire specimen surface was not covered with an apatite-like coating (Figure 3B). In the case of the PTCAPed processed alloy (UFG–WE43 Mg) with an average grain size of ~1 μm, thicker white precipitates were formed after 7 days of immersion, which spread further after 14 days of soaking in SBF compared to the unprocessed sample (Figure 3C and D). The morphological features of the precipitates on the HA/UFG–WE43 Mg implant were significantly different from the other specimens after 7 and 14 days of soaking in SBF (Figure 3E and F). As can be seen, a thick apatite layer has developed throughout the sample, which has materialized a bimodal flake-like structure, comprising HA PVD coating and a bone-like apatite layer (30).
Degradation mechanism. The degradation mechanism of WE43 Mg alloy with regard to dissimilar grain sizes immersed in SBF is schematically displayed in Figure 4. In this context, similar degradation mechanisms were proposed for coarse-grained and UFG Mg-based alloys immersed in SBF and PBS (37, 38).
Muscle cell adhesion and morphology. The cell adhesion strength was examined through assessment of the cell resistance against mechanical detachment, as shown in Figure 5A. For this purpose, the cells incubated for 1, 2, 6, 12, and 24 h were exposed to a vibrating force as explained in Materials and Methods and the obtained data are shown as the mean±standard deviation (n=3) and p<0.05 was considered to indicate a statistically significant difference. In Figure 5, the number of cells remaining after mechanical detachment of the 1, 2, 6, 12, and 24 h incubated samples were respectively 7.1, 11.1, 6.3, 5.9, and 4.4% greater on UFG–WE43 Mg surfaces compared to the untreated alloy. After the HA surface modification (HA/UFG–WE43 Mg), the number of cells remaining after mechanical detachment increased to 10, 13.9, 12.5, 10.6, and 5.6% following the same incubation periods.
Figure 5B shows the SEM micrographs of L6 cells (derived from rat skeletal muscle) cultured for 1, 2, 6, 12, and 24 h on the specimens. A group of cells appeared on HA/UFG–WE43 Mg specimens following 1 to 24 h incubation, unlike individual cells that materialized on the control. By increasing the incubation time from 1 to 24 h, it is obvious that the cells began to stretch on the HA/UFG–WE43 Mg specimen, and the interactions of cell–to–cell and cell–material increased substantially compared to the control. Following 24 h of culture, good stretching of the cells to cover the sample surface was seen implying that the beneficial surface properties of HA/UFG–WE43 Mg enhanced cell adhesion and spreading.
The effect of HA/UFG–WE43 Mg on cell proliferation. Herein, cell proliferation was measured via CCK 8 analysis following incubating UFG–WE43 Mg and HA/UFG–WE43 Mg implants for specified times with L6 cells, as shown in Figure 6A. Cell proliferation (relative growth rates) can be determined by (39):
X
The values of CCK 8 were estimated using mean±standard deviations from 5 wells (SD, n=5), where the disparities between the groups were considered statistically significant at p<0.05. From Figure 6A, an increase in cell proliferation was observed in both treatment groups, i.e., UFG–WE43 Mg and HA/UFG–WE43 Mg implants, compared to the control group. This shows that the formation of the UFG structure promoted cell growth and proliferation. Nevertheless, relative growth rates for HA/UFG–WE43 Mg implant were higher than that of UFG–WE43 Mg following 1, 3, 5, and 7 days, signifying that HA/UFG–WE43 Mg implant possesses higher bioactivity than UFG–WE43 Mg.
Figure 6B displays the morphology and density of cells, co-incubated with the WE43 Mg, UFG–WE43 Mg, and HA/UFG–WE43 Mg implants for 3 days, as observed by fluorescence microscopy. Contingent on their site with respect to the insert’s membrane, diverse cell morphologies were detected. Accordingly, cells with a lengthened form are seen in the zones with extreme cell density, whereas cells towards the centre of the well exhibit a polygonal shape with lower cell density.
Animal blood biochemical characteristics. The results of animal blood biochemical testing of UFG–WE43 Mg and HA/UFG–WE43 Mg implants performed on 1, 3, 5, 7, and 15 days after the operation are illustrated in Figure 7. As can be seen, there are significant differences in the serum concentrations of Mg, as the main element of the implant in both groups (Figure 7A and B).
X-ray photographs and histological analysis. The X-ray photographs of the fracture region in the HA/UFG–WE43 Mg group and the results of the histological analysis by the Masson’s trichrome staining of rat tissues at 15- and 30-days post-operation are shown in Figure 8. X-rays of the surgery zones in the tibia of rats at 15- and 30-days after surgery in each group showed that the HA/UFG–WE43 Mg implants were well located during bone healing and the fracture healed progressively.
Discussion
Compared to the single pass PTCAP, by increasing passes, the recrystallized grain volume fraction increased and inversely the mean grain size was reduced. This behaviour can be attributed to the alloying elements in the alloy composition. WE43 Mg alloy contains some rare elements, e.g., yttrium (Y), neodymium (Nd), and gadolinium (Gd), which inhibit the movement of grain boundaries through the Zener pinning effect, whereby has a strong influence on recovery, recrystallization, and grain growth (40). Besides, new grains are preferably nucleated near the initial grain boundaries that can be ascribed to the strain-induced boundary migration (41). In this case, grain refinement begins after the first pass of PTCAP, resulting in a better distribution of IMCs compared to the untreated sample. By executing the second pass of PTCAP, the structure became more homogenous, the fraction of new grains and grain boundaries increased and the coarse grains slowly disappeared. Finally, at the end of the third pass, a refined and homogeneous grain structure was obtained.
Herein, dislocations accumulation (so-called dislocations pill-up) at secondary phase particles creates a significant amount of stress that may put in force the dislocations to cut through the second phase particles. It seems that the level of requisite stress to shear the second phase can be attained during the PTCAP process taking into account the structural properties of the eutectic second phase. As the dislocations pass through the secondary particles, an absolute shearing occurs and the Cottrell atmosphere is formed by impurity atoms around dislocations. It should be noted that cutting and dissolving the second phases may end to a more active surface for diffusion that in turn would result in the dissolution process speeding up in a shorter period (42). On the other hand, Frank-Read sources of dislocations, which can be strongly activated during the PTCAP method, may also act as a non-contact mechanism for the dissolution of the second phase (43). In this approach, high equivalent strain during the PTCAP may activate Frank-Read sources, where the newly created dislocations holding no segregation can alter the equilibrium state of dislocation segregation to the non-equilibrium state. Under such a high strain rate, the diffusion process may be accelerated and, given enough time for diffusion, the Cottrell atmosphere around the newly created dislocations’ cores attracts impurity atoms through the matrix. On the basis of this mechanism, diffusion fluxes are directed toward the dislocation’s core atmospheres and make an effort to keep steady the impurity concentration in the overall matrix. This destroys the equilibrium state between the dissolving phase and the matrix, leading to an elevated thermodynamic motivation for the dissolution of the dissolving phase. Accordingly, the mechanical strength is expected to decrease; however, the substantial strengthening effects of grain refinement supported by the Hall-Petch theory can prevail over the negative effect of dynamic dissolution.
After surface modification, due to the flake-like structure of the HA coating, the thickness of the HA layer could not be clearly determined. However, the mean flake thickness was 12±5 nm. Furthermore, it can be seen that all the substrate surface defects, such as pits and cracks, were covered with HA coating deposited at 250 W for 90 min, compared to the bare substrate. We also found that the formed HA layer on the UFG–WE43 Mg implant was more dispersed at an RF power of 250 W than that of HA coating deposited with lower RF power levels. These results are in good agreement with the relevant literature (32, 44). XRD results indicate that a crystalline HA coating was successfully developed on UFG–WE43 Mg implant. So, based on the present observations, we expect to observe improved bioactivity in such a HA/UFG–WE43 Mg implant after immersion in SBF.
From the elemental analysis, Mg, O, C, P, Ca, and Cl were detected on the surface of the unprocessed sample, which originated from the bare substrate and SBF solution. The manifestation of Cl on the unprocessed sample is ascribed to the formation of magnesium chloride (MgCl2), as described in the literature (45). Thicker white precipitates on UFG–WE43 Mg show its higher bioactivity compared to the unprocessed sample. The results indicate that the combined effect of the UFG structure and HA surface modification of WE43 Mg led to rapid biomineralization and the formation of a bone-like apatite layer through the heterogeneous nucleation in SBF and thus enabling it to bond to the living bone through a calcium phosphate layer in vivo (26). Rapidly induced apatite formation is also reported in the case of bamboo leaf-like HA coating on the Mg alloy, which exhibited superior corrosion resistance after the 90-day immersion (46).
As Mg and its alloys are immersed in an aqueous solution, magnesium hydroxide (Mg(OH)2) is formed as a byproduct thanks to Mg corrosion. Mg(OH)2 deposits serve as a barrier layer for further corrosion in the initial stage of degradation. On the other hand, in the presence of Cl−, Mg(OH)2 layer is unstable. This results in the formation of MgCl2 in the physiological medium, in which Cl− concentration is high and Mg(OH)2 layer weakens, finally leading to increased pit size and number of pits. Due to the high solubility of MgCl2 in the aqueous solution, the rate of WE43 Mg alloy degradation increases (47). Suspending the development of MgCl2 can be an effective way to protect the Mg(OH)2 layer and achieve controlled Mg dissolution. This can be attained by increasing mineralization on surfaces with protective phases that further delay the dissolution rate. According to the biomineralization behaviour of the HA/UFG–WE43 Mg, the fraction of precipitates, which are a combination of different phases, increases as the immersion time increased to 14 days. In this case, the Mg(OH)2 layer becomes stable, and localized pitting reduces due to the enhanced biomineralization on it. This contributes to attaining a higher level of protection against the aggressive chloride ions on the HA/UFG–WE43 Mg sample. Although HA/UFG-WE43 Mg shows a favourable biodegradability rate, the optimal surface area and porosity must be achieved for chemical incorporation or physical encapsulation of diagnostic and therapeutic compounds. Since the present study uses a line-of-sight PVD coating process, it is very difficult to form a uniform layer on the UFG alloy. Due to the increased cohesion forces that act among finer grains, this crisis can be exacerbated when smaller grains are achieved, whereby the access to the overall surface of a grain bed is more restricted.
Cells derived from rat skeletal muscles showed increased adhesion strength at the initial stage of culture on HA/UFG–WE43 Mg surfaces compared to the unprocessed WE43 Mg surfaces. The main reasons behind the enhanced cell adhesion for HA/UFG–WE43 Mg are high surface energy and low degradation rate that favours the cell adhesion. Moreover, surface roughness takes the main part in cell-material interactions (45, 48) so an implant with high surface energy and optimum roughness can encourage protein adsorption that further helps for cell attachment and proliferation. Based on the SEM micrographs of L6 cells, the cellular response to produce and mineralize their extracellular matrix proteins, as well as cytoskeleton proteins and adhesion molecules to assist in better adhesion and proliferation (49) on the HA/UFG–WE43 Mg implant was noticeable owing to the improved biomineralization.
In fluorescence microscopy images, the cell density recorded for the UFG–WE43 Mg alloy was lower than that of control wells (comprising only cells without alloy), but higher than the unprocessed alloy (WE43 Mg) after 3 days. In the case of the HA/UFG–WE43 Mg implant, after 3 days of co-incubation, the cell density is similar to the control. These alterations in the morphology and density of cells can be ascribed to the disparities in the concentrations of Mg (50) and Ca (51) in the micro-medium of the well.
Alterations in blood biochemical characteristics clearly originated from the surface modification of the UFG–WE43 Mg alloy by HA PVD coating. Referring to the biomineralization behaviour of the HA/UFG–WE43 Mg, the Mg(OH)2 layer becomes stable owing to the HA coating, which contributes to attaining a higher level of protection against aggressive biodegradation. There is no significant difference in Ca levels at 15 days after implantation, showing a fully crystallized structure of the HA coating with a controllable biodegradation rate. In addition, the results of the histological assessments of the implanted zones at 15- and 30-days post-operation corroborate sufficient biodegradation of the implanted HA/UFG–WE43 Mg and a restrained inflammatory reaction. Also, the tissue biocompatibility of the implanted HA/UFG–WE43 Mg showed less systemic and local inflammation, as well as adequate resorption proportional to the speed of bone remodelling.
Conclusion
In this work, a hybrid approach was developed by integrating PTCAP and PVD coating techniques to form a single framework of a biodegradable implant. For this purpose, UFG–WE43 Mg alloy was first processed by PTCAP, and then PVD magnetron sputtering was used to modify its surface properties through a homogeneous flake-like HA coating (12±5 nm). The results showed that the combination of PTCAP processing followed by surface modification resulted in rapid biomineralisation and the development of a bone-like apatite layer after 14 days of immersion in SBF. Compared to the untreated alloy, cells derived from rat skeletal muscles showed increased adhesion strength in the early stage of cultivation on the processed alloy. HA coating helped achieve a higher level of protection against invasive biodegradation. Cells began to stretch on the HA/UFG–WE43 Mg specimen by increasing the incubation time, and the biological interactions increased substantially compared to controls. The relative growth rates for the HA/UFG–WE43 Mg implant were higher than that of UFG–WE43 Mg and the cell density was similar to the control after 3 days of co-incubation. X-ray photographs and the results of the histological analysis showed that HA/UFG–WE43 Mg implant had less inflammation and adequate resorption proportional to the speed of bone remodelling.
Acknowledgements
The Authors would like to acknowledge the Department of Orthopaedics, The First Affiliated Hospital of Xi’an Jiaotong University for providing the necessary resources and facilities for the present study.
Footnotes
Authors’ Contributions
Ma D: Data analysis; data curation; methodology; formal analysis; writing – original draft. Zhang K: investigation; software; critical revision. Dong B: critical revision; approval of the article. She J: concept; critical revision; approval of the article. Zhang Y: conceptualization; validation; writing – review & editing; supervision.
Conflicts of Interest
The Authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.
- Received May 13, 2022.
- Revision received July 23, 2022.
- Accepted July 26, 2022.
- Copyright © 2023, International Institute of Anticancer Research (Dr. George J. Delinasios), All rights reserved
This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY-NC-ND) 4.0 international license (https://creativecommons.org/licenses/by-nc-nd/4.0).