Bioabsorbable magnesium alloys are widely studied for various implant applications, as they reduce the risks such as severe inflammatory response existing in permanent metallic implants. However, the over-fast corrosion rate of magnesium alloy is usually an obstacle in biomedical applications. Here we report a simple two-step reaction to introduce anticorrosive silane pre-treatment on MgZnYNd alloys before coating with poly (glycolide-co-lactide) (PLGA). The first step is to immerse the NaOH-activated MgZnYNd with bistriethoxysilylethane (BTSE) to form a cross-linked silane coating layer with enhanced corrosion resistance; the second step involves immobilizing amine functional groups for forming hydrogen bond with outer PLGA coating by treating the BTSE-modified MgZnYNd with 3-amino-propyltrimethoxysilane (APTES). We characterized the BTSE-APTES pre-treated PLGA coating on MgZnYNd by scanning electron microscopy (SEM), Fourier transform infrared spectroscopy (FT-IR), X-ray photoelectron spectroscopy (XPS), static contact angle and Acid Orange 7 measurement. Nano-scratch test was to verify that the scratch resistance of the PLGA coating with BTSE-APTES pre-treatment was superior to direct PLGA coating. Standard electrochemical measurements along with the long-term immersion results indicated that the BTSE-APTES pre-treatment rendered better in vitro degradation behavior. Cell adhesion and cell viability tests with both vascular smooth muscle cells (VSMC) and human umbilical vein endothelial cells (EA. hy926) demonstrated that BTSE-APTES pre-treated MgZnYNd substrate had significantly more beneficial effects. The favorable anti-corrosion behavior and biocompatibility of BTSE-APTES pre-treated PLGA coatings on MgZnYNd alloy suggest that the novel two-step silanization procedure may have the great potential to enhance the performance of the magnesium-based biomaterials and provide a valid solution for the conversion modification of cardiovascular implants, taking the magnesium-based bioabsorbable materials closer to clinical application.
In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7].
Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19].
Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25].
Bistriethoxysilylethane (BTSE) and 3-amino-propyltriethoxysilane (APTES) are extensively researched bissilane and monosilane, respectively, providing functional groups that could further attach drug-loading layer or bioactive molecules as well as enhance the interfacial reactions with metallic surfaces[27]. However, the single silane layer will preferentially bond to surfaces of metal materials and result in occasional flaws in the silane layers, permitting solvent to permeate into the silane-metal interface, while much denser cross-linked three-dimensional network and further stronger interfacial adhesion with metal surfaces can be achieved with ‘bis-silanes’, i.e., metal-silane A-silane-B[28]. To make use of both types of silanes, a facile two-step 3-amino-propyltrimethoxysilane (γ-APS) pre-treatment was designed to provide fine anti-corrosion protection for steel and aluminum[29] and AZ31[18]. Further, as direct coating of PLGA layer onto metal surface could only form invalid bond, stronger hydrogen bond between exposed amino group of APTES and PLGA coating is another reason we adopted this two-step silane coating to improve the anti-corrosion ability and biocompatibility of PLGA protective layer on MgZnYNd alloy for cardiovascular stent. To our knowledge, the thorough study of the biocompatibility of the bilayer system has not been reported before.
We designed a novel two-step BTSE-APTES pretreating method to produce a functional corrosion resistant and biocompatible cross-linked silane coating on MgZnYNd substrate before spinning with PLGA solution. The silane layer was analyzed by scanning electron microscopy (SEM), Fourier transform infrared spectroscopy (FT-IR), X-ray photoelectron spectroscopy (XPS), static contact angle, and Acid Orange 7 (AO) measurement; the anti-corrosion properties of the coating were assessed by electrochemical measurements and the long-term immersion test. We confirmed that PLGA coating could be more tightly bonded onto the double silane layer modified MgZnYNd alloy surface via nano-scratch test. Cytocompatibility was verified with cell adhesion and cell viability tests via VSMC and EA. hy926. Using this technology, we have verified significantly improved anti-corrosion ability and biocompatibility of the prescribed Mg alloy.
2. Experimental
2.1. Materials and surface treatment
Every reagent was used as received. BTSE, APTES, phosphate-buffered saline (PBS), glutaraldehyde, Triton X-100, Acid Orange 7, FITC-phalloidin, and tetrazolium-8 (WST-8) were from Sigma-Aldrich, China; DAPI was purchased from CST, China; Griess reagent Kit was from Beyotime, China. PLGA (Mw = 100,000, GA:LA ratio = 25:75) was ordered from Daigang Biology, China. As-extruded MgZnYNd alloy bars (Mg-2Zn-0.5Y-0.5Nd)[8] were provided by Guan's lab at Zhengzhou University, China. Hank's solution (containing 8.00 g NaCl, 1.00 g glucose, 0.40 g KCl, 0.35 g NaHCO3, 0.14 g CaCl2, 0.06 g KH2PO4, 0.06 g MgSO4, 0.06 g Na2HPO4, 0.01 g MgCl2 per 1000 mL deionized water) was freshly self-prepared and pH value was modulated to 7.40 (37.0 ± 0.5 °C) prior to the tests.
Round sample discs (D = 10 mm, H = 1 mm) were cut from as-extruded Mg-Zn-Y-Nd alloy round bars by electrical discharge machining. All the samples were polished using up to 2000 grit SiC papers to ensure consistency across all tests. Samples were then ultrasonic cleaned for 20 min in ethanol, and then immersed in 3.0 mol/L NaOH solution for 2 h to generate a homogeneous hydroxide layer on the surface of samples before storing in a vacuum desiccator.
BTSE and APTES solutions were similarly prepared using 90% ethanol, 5% silane and 5% deionized water. Both of the solutions were mildly stirred at 25 °C for 1 h to make BTSE or APTES hydrolyze. The MgZnYNd substrates were first immersed in the hydrolyzed 5% BTSE solution at 25 °C for 1 h before being dried by N2 flowing and being cured and solidified for 1 h at 120 °C. After that, the BTSE pre-treated substrates (denoted as MgZnYNd-B) were immersed in 5% APTES solution at 25 °C for 30 min before being cured and solidified for 1 h at 120 °C. The resultant samples were referred as MgZnYNd-B-A.
The corresponding 4% wt/vol PLGA solution was dissolved with dichloromethane at room temperature. PLGA films on MgZnYNd substrates were prepared by spin coating[30]: 100 µL PLGA solution was dropped onto MgZnYNd-B-A sample, the spin process was set to be 300 rpm for 6 s then 7000 rpm for 20 s, both sides of substrates were prepared similarly before drying for 48 h at 37 °C. Samples with only APTES treatment before PLGA coating and samples with direct PLGA layer were likewise prepared for comparison and donated as MgZnYNd-A-P and MgZnYNd-P, respectively.
2.2. Surface microstructure characterization
Microstructural characterizations of both surfaces and cross-sections were observed for the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P samples by SEM (Hitachi s-4800) at 10 kV. The functional group of the coatings was further confirmed by attenuated total reflectance FT-IR (NICOLET 750, USA), in which spectra were obtained in a range of 675-4000 cm-1.
The AO test was carried out to confirm the amine concentration on the surface of MgZnYNd-B-A. MgZnYNd-B-A was immersed in 500 µmol/L AO-HCl solution (pH = 3,) and shaked for 5 h at 37 °C before rinsing with HCl solution (pH = 3). NaOH solution (pH = 12) was used to unload the adsorbed AO. Finally, the optical density (OD) of desorbed AO supernatant was recorded using a microplate reader (Model 680, Bio-Rad, CA) at 485 nm. The resulting OD was positively related to the amine concentration on the surface of the samples.
The structure of the silane film was further confirmed by X-ray photoelectronic spectroscopy (XPS) measurement (ESCALAB 250) with the Al Kα line (1486 eV) as the excitation source. Casa XPS software was used to fit the narrow spectra. The static contact angle (CA) was measured using a Contact Angle Meter (SL200B, USA) at 25 °C by dropping 2 µL deionized water on the surface. Each sample was tested in triplicate.
To evaluate the adhesion enhancement of the PLGA layer on MgZnYNd substrates by BTSE-APTES pre-treatment, nano-scratch resistance tests were conducted using a Nano Indenter system (Hysitron, USA) with a spherical diamond Rockwell indenter. The tip approached and loaded into the bi-layer coating with increasing rate of 30 µN/s up to 1000 µN, making a 10 µm scratching at the prescribed direction. The Nano Indenter system could detect the first cracks of the coating corresponding to the critical load[31]. The tests were performed in clean environment at 25 °C. For statistical purpose, at least three parallel results were deemed valid and average values were calculated and reported.
2.3. In vitro degradation behavior
2.3.1. Electrochemical test
Standard methods were applied in electrochemical measurement[32] and [33]. CHENHUA CHI650C system (Shanghai, China) with a three-electrode cell was used, in which MgZnYNd samples acted as working electrode, saturated calomel electrode (SCE) as the reference electrode, and the platinum as the counter electrode. Both potentiodynamic polarization curves analysis (Tafel plot) and electrochemical impedance spectroscopy (EIS) were performed in 125 mL Hank's solution at 37.0 ± 0.5 °C with working area of 0.332 cm2. For EIS measurements, the frequency varied from 105 Hz to 10-2 Hz after reaching a steady open-circuit potential value, and spectra were analyzed through Nyquist plots. The Tafel plot was tested at a rate of 0.5 mV ⋅ s-1 from -600 mV cathodically to +800 mV anodically relative to the open circuit potential (OCP), and the Ecorr and Icorr were obtained by fitting Tafel plot with the Corrview software according to the criterion ASTM-G102-89. For statistical purpose, at least three parallel results were deemed valid and average values were calculated and reported.
2.3.2. Long-term corrosion behavior
Long-term (30 days) corrosion behavior were proceeded according to ASTM G31-72 standard, each sample was immersed with a ratio of 20 mL/cm2 in Hank's solution at 37 °C with 5% CO2 in a humid and sterile condition. After prescribed immersion times (15 and 30 days), the samples were rinsed with deionized water to wipe out the corrosion products. SEM equipped with Energy Dispersive Spectrometry (EDS) was used to confirm post-corrosion morphology as well as composition of the surface. Every 3 days until the 30th day, the instant pH value was recorded, and the Mg2+ concentration was tested by analyzing the extracted incubation media after 15th and 30th day immersion with inductively coupled plasma optical emission spectroscopy (ICP-OES, Thermo Fisher, USA), the original Mg ion concentration in Hank's solution was excluded when comparing between different groups.
2.4. Cellular behavior
Human umbilical vein endothelial cells (EA. hy926) and rodent vascular smooth muscle cells (VSMC) were adopted for observing cellular behavior in vitro. Cells were cultured in 5% CO2, 37 °C, and 95% relative humidity atmosphere in Dulbecco's Modified Eagle Medium (DMEM, Corning, China) with 10% fetal bovine serum (FBS, Corning, China) and 1% penicillin/streptomycin antibiotics (P/S, Corning, China).
2.4.1. Cellular adhesion and morphology
Cellular morphology and cytoskeleton image were observed via actin and nucleus staining. VSMC cells were reseeded at a concentration of 50,000 mL-1 on sterilized MgZnYNd-P, MgZnYNd-A-P, MgZnYNd-B-A-P, bare MgZnYNd substrates and cultured for 24 h and 48 h. Polystyrene slice worked as blank control. Afterwards, cells were washed twice with PBS to remove unattached cells, and then fixed with 4% (w/v) paraformaldehyde for 8 min followed by being permeabilized with 0.1% (v/v) Triton X-100 for 10 min. After this, cellular actin and nuclei was stained with 1.0% (v/v) FITC-phalloidin and 1 µg/mL DAPI for 5 min, respectively. Cells morphologies were imaged by confocal microscopy (Nikon, Japan).
2.4.2. In vitro cytotoxicity
Both VSMC and EA. hy926 were adopted to assess the in vitro cytotoxicity of MgZnYNd-P, MgZnYNd-A-P, MgZnYNd-B-A-P, and bare MgZnYNd substrates. Cell viabilities were measured following the direct assays according to the criteria in ISO 10993-12.
UV radiation (254 nm, 40 min each side) was used to sterilize samples before use. Cells were reseeded at a concentration of 60,000 cells ⋅ mL-1 in 24-well plates (Costar, USA). After seeding for 24 h, MgZnYNd-B-A-P and MgZnYNd-P samples as well as uncoated MgZnYNd substrates were exposed to cells, DMEM solution with serum worked as blank control, DMEM with 10% dimethylsulfoxide (DMSO) as positive control. After culturing for 1, 3, 5 days, the media were extracted and unattached cells were removed by rinsing samples with PBS solution followed by adding 500 µL/well culture medium containing 10% (v/v) WST-8 before 3 h incubation. Optical density (OD) of supernatant was tested by microplate reader (Bio-Rad, CA) at 450 nm (630 nm as reference wavelength). Cell viabilities were calculated according to the following formula:
2.4.3. NO release
Both EA. hy926 cells and VSMC cells were collected and adjusted to a density of 5 × 104 cells/well before reseeding onto sterilized MgZnYNd-P, MgZnYNd-A-P, MgZnYNd-B-A-P and bare MgZnYNd substrates in a 24-well plate (Costar, USA). DMEM solution with serum acted as blank control. The cell culture supernatant was collected and tested by Griess reagent Kit following the instructions after 24 h culture.
2.5. Statistical analysis
All of the tests were conducted at least in triplicate independently for each replicates, and the data were expressed as means ± standard deviation. One-way ANOVA analysis was used for the statistical analysis, a p value of less than 0.05 was considered statistically significant.
3. Results and Discussion
3.1. Coating characterization
3.1.1. Surface microstructure characterization
Fig. 1(c) exhibits the procedure of surface modification for the MgZnYNd alloy. The MgZnYNd alloy was pre-treated with alkaline solution to produce hydroxide groups for bonding the silane coating. MgZnYNd-OH was first treated with the hydrolyzed BTSE (Fig. 1(a)) then APTES (Fig. 1(b)) to fabricate a bilayer cross-linked silane coating via eliminating water and forming oxane bond, i.e., polysiloxane networks with surface amine functionality formed through in-situ polycondensation of the hydrolyzed BTSE and APTES. Besides, PLGA was commonly employed as a functional drug-loaded coating model, forming hydrogen bonding with the inner aminized silane coating to improve the biocompatibility of MgZnYNd substrate.
Fig. 1.
Chemical structure of BTSE (a) and APTES (b); (c) surface modification procedure for the MgZnYNd alloy.
The SEM results (Fig. 2(a)) show that BTSE and BTSE-APTES pre-treated MgZnYNd substrate (MgZnYNd-B and MgZnYNd-B-A) exhibited a flat surface with only dark fringe, slight remnants of the polishing grooves from the final 2000# SiC abrasive paper, while PLGA coated MgZnYNd-B-A samples (MgZnYNd-B-A-P) exhibited a non-porous, continuous, dense and smooth coating (Fig. 2(a)). Generally, groove surface on the naked MgZnYNd substrate could not be totally covered by the nano dimension of silane layer thickness; however, the outer layer of PLGA coating masked the grooves very well. Besides, the direct PLGA-coated samples (MgZnYNd-P) showed acceptable smooth surface with a few polishing grooves. It has been reported that corrosion behavior of polymer coated Mg-based materials and cell behavior on it could be affected by film morphologies (porous or dense structure)[34], [35] and [36], while the smooth coating of MgZnYNd-B-A-P in the present study could well fit the anti-corrosive and biocompatible requirements. Moreover, for the cross-section of MgZnYNd-B-A-P, an obvious double layer could be distinguished (Fig. 2(a)), with the coating thickness of 6.4 ± 0.4 µm and 0.5 ± 0.4 µm for PLGA layer and silane coating, respectively. Besides, no significant difference was observed referring to the depth of PLGA layers between MgZnYNd-B-A-P and MgZnYNd-P groups (6.4 ± 0.4 µm vs 6.3 ± 0.8 µm as labeled in red in Fig. 2(a)). While the total coating thickness of MgZnYNd-B-A-P (about 6.9 µm) compared to MgZnYNd-P (about 6.3 µm) could explain the surface morphology difference discussed above.
Fig. 2.
Surface and cross-section morphology and composition analysis. (a) Surface images of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P, MgZnYNd-P and cross-section images of MgZnYNd-P, MgZnYNd-B-A-P; (b) FT-IR spectrum with characteristic absorption bands; (c) surface amine concentration of MgZnYNd-B-A. Scale bar is 20 µm for surface images.
FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39].
AO test was used to quantify amine concentrations on the surface of MgZnYNd-B-A, shown in Fig. 2(c). MgZnYNd-B-A showed a markedly higher (p > 0.05) amino concentration of 97.17 ± 6.52 nmol/cm2 when compared with the blank control (bare MgZnYNd alloys substrates), indicating that the APTES was successfully coated onto MgZnYNd-B with appropriate amount of amino groups. The data were comparable with the TiOH (pristine Ti activated by NaOH) silanized by γ-amino-propyltriethoxysilane (γ-APS), which is about 127 nmol/cm2[40]. However, few amino groups were detected on the surface of MgZnYNd-B-A-P samples because of the coverage effect of PLGA layer.
In order to understand the type the interface bonding exists in the cross-linked silane film on the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P surfaces, XPS spectra were scanned. The whole and narrow scan spectra of O 1s and Si 2p were analyzed using Casa XPS software and shown in Fig. 3(a, b), elemental compositions were calculated from XPS intensities and denoted as atomic percentage (at.%) in Table 1. The O 1s binding energy values in the silane film ( Fig. 3(b)) peaked at 533.3, 532.4, and 531.4 eV and were attributed to C-O-C or Si-O-Si [41], [42] and [43], Si-O-Mg [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample.
Fig. 3.
(a) Whole XPS spectra of MgZnYNd-B, MgZnYNd-B-A and MgZnYNd-B-A-P; (b) the fitted O 1s and Si 2p XPS spectra of cross-linked silane films on MgZnYNd-B, MgZnYNd-B-A and MgZnYNd-B-A-P, respectively.
Table 1
Table 1
Characterization of polysiloxane coatings by their elemental composition calculated from XPS intensities and denoted as atomic percentage (at.%)
Sample
O 1s (%)
C 1s (%)
Si 2p (%)
Y 3p (%)
N 1s (%)
MgZnYNd-B
38.36
33.77
26.07
1.86
—
MgZnYNd-B-A
29.37
41.34
23.23
1.4
4.66
MgZnYNd-B-A-P
36.31
61.56
2.13
—
—
Table 1
Characterization of polysiloxane coatings by their elemental composition calculated from XPS intensities and denoted as atomic percentage (at.%)
3.1.2. Contact angles
The static contact angle (CA) measurements could reflect changes of surface wettability on the modified MgZnYNd alloys, thus confirming successful bonding of silane with the presence of characteristic groups as shown in Fig. 4. The CA of the original bare MgZnYNd sample was 64.02° ± 3.42°, and increased to 101.09° ± 8.14° with the hydrophobic BTSE coating, which was due to the cross-linked Si-O-Si structure, and was consistent with the reported data[18]. While with a lower CA of 55.18° ± 0.87°, MgZnYNd-B-A was more hydrophilic, attributing to the superficial amine groups. However, subsequent coating of PLGA layer reproduced the surface hydrophobicity, with the CA value increasing to 66.39° ± 2.90°, which was closer to a pure PLGA with less influence of the underneath MgZnYNd substrate[38].
Fig. 4.
Water contact angle (θ) for bare MgZnYNd, MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and MgZnYNd-P.
3.1.3. Nano-scratch resistance
Nano-indentation is a popular technique to detect micro-scale mechanical properties of surfaces; particularly it offers an efficient and simple means to estimate the scratch-resistance of the coated films[31] and [44]. The initial coating failure could indicate the critical lateral force in the ramp-load scratch steps and is in positive correlation with the adhesion strength of the coating, and the data are shown in Fig. 5. Given this, the critical lateral forces for MgZnYNd-B-A-P was significantly higher than the MgZnYNd-P and MgZnYNd-A-P group, i.e., 435 µN at 27 s vs. 300 µN at as early as 25 s and 320 µN at as early as 22 s for MgZnYNd-P and MgZnYNd-A-P, respectively, a 45% and 36% higher critical lateral force of MgZnYNd-B-A-P than MgZnYNd-P and MgZnYNd-A-P. While both the force and time was necessary to consider when analyzing the nano-scratch resistance data. These lateral force/time data indicate overall the much stronger adhesion between PLGA and MgZnYNd substrates can be influenced by two-step BTSE-APTES pre-treatment than the one-step APTES pre-treated coating or direct PLGA coating. The stronger adhesion between coating and substrate existed in MgZnYNd-B-A-P system can be attributed to strong covalent bond between substrate and silane layer, besides, the intermolecular hydrogen bond and Van der Waals force between silane layer and PLGA could be deduced from the chemical formula of APTES. The critical lateral force for MgZnYNd-B-A-P group (435 µN at 27 s) was a little higher than the published research by Liu et al., phenylalanine-based poly (ester amide)s (8-Phe-4) films coated magnesium substrate[38], which is 380 µN at 25 s - 403 µN at 26 s, respectively, for 4% and 2% 8-Phe-4 coatings.
Fig. 5.
(a) Steps for the nano-scratch resistance test; (b) lateral forces changed with time were recorded for MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples during the tests, green dots signifies the critical point of the lateral forces.
3.2. In vitro degradation behavior
3.2.1. Electrochemical test
Electrochemical measurement is one of the most universal tests to estimate the in vitro corrosion and degradation behavior of metal-based materials, which mainly includes OCP plots, Tafel plots and EIS plots.
OCP plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate are shown in Fig. 6(a). The steady OCP after 3600 s for MgZnYNd-B-A-P was -0.48 V, which was much less negative than the MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd groups (with a similar OCP of -1.55 V), indicating the lessened corrosion possibility of the MgZnYNd-B-A-P compared to other groups.
Fig. 6.
Electrochemical measurement for bare MgZnYNd, MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples: (a) open circuit potential; (b) potentiodynamic polarization plots; (c) Nyquist plots.
Tafel plots for various MgZnYNd samples are revealed in Fig. 6(b). From the extrapolation of the Tafel plots, the corrosion potential (Ecorr) and corrosion current density (Icorr) were calculated and shown in Table 2. The BTSE-APTES pre-treatment on MgZnYNd substrates led to Ecorr values (-0.495 V) less negative than the MgZnYNd-P (-1.568 V), MgZnYNd-A-P (-1.550 V), and uncoated substrates (-1.556 V), implying retard of corrosion. Moreover, the reduction in Icorr values by 77.60%, 48.70%, and 31.81% in the MgZnYNd-B-A-P than the bare MgZnYNd sample, MgZnYNd-P, and MgZnYNd-A-P, respectively, was also companied by the higher Ecorr values of the BTSE-APTES pre-treatment on MgZnYNd. Conversely, MgZnYNd substrates with direct PLGA layer led to a more negative Ecorr, implying a less effective barrier function against corrosion provided by the PLGA without silane pre-treatment. The similar trend of Ecorr increases from -1.614 V to -1.490 V was also observed in Mg-Li alloy with assembled ZSM-5 layer adopting silane coupling agent as linkage[45].
Table 2
Table 2
Open circuit potential (OCP), corrosion potential (Ecorr), and current density (Icorr) values for MgZnYNd samples using different treatment methods
Sample
OCP (V)
Ecorr (V)
Icorr (µA cm-2)
Zre (Ω)
MgZnYNd
-1.556
-1.396
31.910
1860
MgZnYNd -P
-1.568
-1.475
13.935
3410
MgZnYNd -A-P
-1.550
-1.148
10.482
27800
MgZnYNd -B-A-P
-0.495
-0.405
7.148
118750
Table 2
Open circuit potential (OCP), corrosion potential (Ecorr), and current density (Icorr) values for MgZnYNd samples using different treatment methods
The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50].
After NaOH passivation, MgZnYNd surface was covered by Mg(OH)2 film, and reactions existing in anode and cathode are Mg-2e-→Mg2+Mg-2e-→Mg2+and Mg2++2H2O+2e-→Mg(OH)2+H2Mg2++2H2O+2e-→Mg(OH)2+H2, respectively. As anodic Mg dissolves beneath the Mg(OH)2 layer on the MgZnYNd surface, Mg2+ transport determines the anodic kinetics; while charge transferring relates the cathodic reaction can occur both underlying and over the silane film, leading to a higher cathodic current density of bare MgZnYNd than the bilayer silane groups. After one-layer pretreating with APTES, only a limited retardant silane molecule layer formed on the surface; thus the Icorr of MgZnYNd-A-P decreased by only 3.45 µA/cm2 compared with MgZnYNd-P. However, after BTSE coating, the exchange current density results from the cathodic reaction is greatly decreased because the hydrophobic silane film could retard electron transport and water penetration, thus forming a physical barrier. The anodic reaction is also slowed as the bilayer silane retards the Mg2+ transport[28]. Meanwhile, the Si-O-Mg bonding at the MgZnYNd-BTSE interface could also block anodic reactions[39], which lowers the MgZnYNd-B-A-P corrosion current value by an order of magnitude than MgZnYNd-A-P.
EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53].
Consequently, the significantly more positive OCP and Ecorr values with lower Icorr values in the Tafel curves of the MgZnYNd-B-A-P (Fig. 6(a, b)) combined with capacitive loops with larger diameter in the EIS plots (Fig. 6(c)) indicated that the cross-linked bilayer BTSE-APTES pre-treatment had the most effective effect in enhancing corrosion resistance ability among all the testing samples.
3.2.2. Long-term corrosion behavior
Complying with the current recommended criterion in ASTM G31-72[54] and [55], various MgZnYNd alloy substrates were immersed in the Hank's solution for 30 days at 37 °C to quantify the silane coatings effect on the corrosion rate of the MgZnYNd substrate.
SEM results (Fig. 7(a)) with EDS spectra (Fig. 7(b)) of the samples after immersing for 15 and 30 days are shown in Fig. 7. The MgZnYNd-P group showed severe delamination and cracks (labeled by red dotted circle) at both the 15th and 30th day data. Similar images of delamination were also reported on poly (l-lactic acid) (PLLA) and poly (ε-caprolactone) (PCL) modified Mg[56] in DMEM solution, and some bubbles effect formed on PLGA-coated Mg4Y and AZ31 alloy after a 3-day incubation[57]. Contrarily, MgZnYNd-B-A-P samples showed uniform and compact surface with only a few superficial narrow cracks and delamination but no gas bubbles, as it escaped from the narrow cracks before it could gather, indicating better anti-corrosion ability of PLGA layer after pre-treatment with BTSE-APTES.
Fig. 7.
(a) SEM images of the MgZnYNd-P, MgZnYNd-B-A-P after immersing in Hank's solution for 15 and 30 days. (b) The EDS spectra after 30th day immersion show Mg, Ca and P peaks. Scale bar is 20 µm and 200 µm, respectively.
In addition, as shown on the 30th day EDS graph and the inserted table (Fig. 7(b)), the MgZnYNd-P group had more precipitation and crystallization containing O, Ca and P elements during the 30-day immersion than the MgZnYNd-B-A-P group, which were also reflected in SEM images with a few spherical and cylindrical crystals. Furthermore, the calculated total Ca and P elements percentage (wt%) was 40.03% from the 30th day EDS spectrum, indicating that apatite crystals probably formed on the PLGA-direct coated samples. Semblable pattern of hydroxyapatite (HA) crystals also formed on the upper surface of PLGA-coated Mg[38]. When it was used as vascular stents, the delamination, cracks and deposited apatite on surface of Mg material would destroy blood homeostasis as reported by Witte et al.[58] that aggregation of erythrocytes would be aroused by apatite particles, further leading to an over-high blood viscosity and clotting[59]. Contrary to the MgZnYNd-P sample, less apatite existed on surface of the MgZnYNd-B-A-P with smaller Ca and P peak value (with a total of 26.4 wt% of Ca and P elements). This different degradation morphology and composition between MgZnYNd-B-A-P and MgZnYNd-P was attributed to the compactness of cross-linking layer and strong chemical bond between PLGA and MgZnYNd substrates, thus forming effective physical barrier to improve the anti-corrosion ability of MgZnYNd alloy.
The pH data of the extracts from Hank's solution was tested every 3 days until the 30th day as degradation behavior of Mg substrates can also be reflected by the pH value (Fig. 8(a)). The in vitro corrosion of magnesium alloy could lead to an alkaline environment with higher pH value.
Fig. 8.
(a,b) Mg2+ concentrations and pH values of the medium for bare MgZnYNd, MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples at different time points (15th day and 30th day). Hank's solution without sample worked as blank control. Data were statistically analyzed using a one-way ANOVA, *** Represents p <0.001, ** Represents p <0.01.
Each of MgZnYNd sample exhibited an increase from 7.4 to over 8.0 initially over the first 3 immersion days. The bare MgZnYNd and MgZnYNd-P samples showed an over-high value of 10.4 by the end of testing period. However, both the MgZnYNd-A-P and MgZnYNd-B-A-P groups could restrain the Hank's solution from becoming over alkaline to avoid overpassing local tissue receptivity, which was much lower than the MgZnYNd-P group with a slight steady increase over the whole period, and reached a steady pH value at about 10.0. The data were lower than the surface-modified magnesium alloys with other methods. For examples, after immersing for 15 days in Hank's solution, AZ91D modified by stabilization-hydrothermal process showed the pH of over 10.0[60]; pure magnesium with alkali-heat treating showed pH value of 10.1 after only 14-day immersion in SBF[61]; after immersing for 30 days in SBF, fluoride-treated AZ31 exhibited pH value of 10.2[16]; at the end of 30-day immersion in Hank's, MAO-modified Mg-Ca showed an over-high pH of 10.5[33].
As for the MgZnYNd-P, the pH value reached 10.4 at the 30th day with a steady increase, similar to the bare MgZnYNd control. Water could penetrate through cracks of PLGA coating and contact with MgZnYNd substrate, resulting from the bulk degradation mode of PLGA polymer and led magnesium alloy to a faster corrosion. Hence, the bilayer BTSE-APTES silane coating could help to keep a much steadier and reasonable in vitro acid-base atmosphere compared with the direct-coated PLGA group. Generally, samples with silane pre-treatment showed more effective barrier performance in significantly keeping pH value under physiological acceptance.
The released Mg2+ concentration from various MgZnYNd samples after immersion over 30 days in Hank's solution was measured by the ICP-OES, which is also an essential indicator in judging the anti-corrosion ability of the Mg-based biomaterials. Data were shown in Fig. 8(b). With the same trend shown in pH value, both MgZnYNd-A-P and MgZnYNd-B-A-P coatings reduced the Mg2+ released more effectively than the MgZnYNd and MgZnYNd-P, indicating an effective barrier functions from the silane coatings. Most importantly, over the whole 30-day immersion period, the MgZnYNd-B-A-P (about 7.77 ppm at 15th day, 83.67 ppm at 30th day) showed much less released Mg2+ compared with the MgZnYNd-P (35.97 ppm at 15th day, 105.27 ppm at 30th day), i.e., a 18.0%-55.8% reduction in the MgZnYNd-B-A-P sample. Besides, when compared with the reported studies, the MgZnYNd-B-A-P released less Mg2+: the stabilization-hydrothermal treated AZ91D shows 80 µmol mL-1 cm-2 Mg2+ after 15-day immersion[60], the accumulated Mg2+ after a 7-day incubation for the LBL PEI-PCL-PAH treated AZ31 is 13 mmol[62], the cumulative Mg2+ value is 80-100 ppm for PLLA/PCL-coated magnesium after 10-day immersion[56]. These tested Mg2+ concentration value trend is in accordance with the recorded degradation surface morphology images and pH data shown above (Fig. 7 and Fig. 8(a)).
3.3. Cellular behavior
3.3.1. Cellular adhesion and morphology
To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend.
Fig. 9.
Human umbilical vein endothelial (EA. hy926) and rodent vascular smooth muscle cell (VSMC) morphology imagined at 24 h ((a)-1, (b)-1) and 48 h ((a)-2, (b)-2) after seeding through actin-nucleus co-staining on bare MgZnYNd, MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples, tissue culture plate acted as blank control. Scale bar is 200 µm and 10 µm, respectively.
On the whole, while contacting with biomaterials, cells will adhere, spread, proliferate, and migrate as the first ordinal reactions. Thus, cytocompatibility of materials is the first priority when considered as biomaterials. The recorded sufficient attachment and spreading of EA. hy926 and VSMC onto the MgZnYNd-B-A-P illustrate that this novel BTSE-APTES pre-treatment should have a beneficial effect on acquiring better biological property of PLGA-coated MgZnYNd than the direct-coated PLGA which was also verified in the cell cytotoxicity data later (Fig. 10).
Fig. 10.
Cell viabilities cultured on bare MgZnYNd, MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples by the CCK-8 assay over 5 days incubation: (a) EA. hy926 and (b) VSMC; DMEM with serum worked as blank control, DMEM with 10% DMSO as positive control; data presented were statistically analyzed using a one-way ANOVA, *** Represents p <0.001.
3.3.2. In vitro cytotoxicity
The cell viabilities of both EA. hy926 (Fig. 10(a)) and VSMC (Fig. 10(b)) grown on the surface of the various MgZnYNd samples for 1, 3, 5 days were tested by the CCK-8 assay, and results are shown in Fig. 10. Viabilities of both EA. hy926 and VSMC seeded on the surface of MgZnYNd-B-A-P, MgZnYNd-A-P, and MgZnYNd-P samples were significantly improved with the cell viabilities of 96.79%-86.83%, 90.20%-62.66%, and 73.85%-51.36%, respectively, for the 5-day incubation period, in contrast with the bare MgZnYNd substrates (ranging 46.11%-20.27%). Therefore, MgZnYNd with silane and/or PLGA coating could all significantly increase the viabilities compared to the bare MgZnYNd. However, for EA. hy926, MgZnYNd-A-P, and MgZnYNd-P groups showed obvious continuous trend of decline with the incubation time, and reached viabilities of 63.00% and 57.45%, respectively. According to the standard ISO 10993-5 2009, viability between 50.00%-79.00% is considered to be Grade II toxicity, which is not acceptable for clinical use. On the contrary, MgZnYNd-B-A-P group showed much higher viabilities than the MgZnYNd-A-P and MgZnYNd-P, which increased slightly during the whole incubation period with the viability of about 96.79% at the 5th day, indicating the good cytocompatibility of BTSE-APTES bilayer, confirming the potential of more remarkable effect in promoting endothelialization. While for VSMC, the data showed significant advantages of MgZnYNd-B-A-P over both MgZnYNd-A-P and MgZnYNd-P, i.e., 6.59%-15.30% higher than MgZnYNd-A-P and 30.23%-35.51% higher than MgZnYNd-P. Besides, for all the groups with silane or PLGA coatings, the 3rd day viability reached to a peak but declined since then. The gradual reduction in cell viability after 3rd day culture in MgZnYNd-B-A-P was probably due to its potential in prevention of restenosis. When referring to other published studies, the cell viability data of the MgZnYNd-B-A-P group for both ECV304 and VSMC were much higher, for example, the L929 cell viabilities cultured for 4 days in the extracts of phytic acid coated WE43 are 28%-70%[68], the HEK293 cell viabilities seeded on the surfaces of BMS-Br and BMS-g-HPBBEA are 68% and 76%, respectively[69], the 7-day HUVECs proliferation data cultured with calcium silicate extracts ranged between 30% and 70%[70].
In the direct assay, cells were in contact with the surface of Mg samples directly, thus the nature and feature of the superficial coating material might be reflected more markedly. Furthermore, the degradation and corrosion behavior of coatings and MgZnYNd substrate in the DMEM could also influence the results of the cell cytotoxicity data. When combining the two characteristics of the direct cell viability assay summarized above, a more notable advantage of BTSE-APTES bilayer in biocompatibility when adopted as Mg alloy conversion coating could be concluded.
3.3.3. NO release
NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P).
Fig. 11.
NO released of (a) EA. hy926 and (b) VSMC in the culture media for bare MgZnYNd, MgZnYNd-P, MgZnYNd-A-P and MgZnYNd-B-A-P samples. DMEM with serum worked as blank control. Data presented were statistically analyzed using a one-way ANOVA, *** Represents p <0.001.
Coupled the data of cell adhesion, cell cytotoxicity and NO release showed above, it is obvious that the bilayer BTSE-APTES pre-treatment benefits a lot when coating MgZnYNd with PLGA, which is reflected in the ability of MgZnYNd-B-A-P to significantly enhance ECV304 and VSMC viability over MgZnYNd-A-P and MgZnYNd-P. Further, the difference in the chemical composition plus the further physical barrier functionality of cross-linked silane coating between MgZnYNd-B-A-P and MgZnYNd-A-P should result to the total different responses for both types of cells. It was reported by Zheng et al. that MC3T3-E1 attachment, stretching, and proliferation were effectively enhanced on GRGD-modified PEEK surface by tailored silanization technique[75]. Besides, silane-coupling agent aminopropyl triethoxysilane was used as a very effective system for collagen immobilizing onto stainless steel to improve human mesenchymal stem cells adhesion and proliferation[27].
4. Conclusion
Using a two-step conversion coating process, we prepared bio-functionalized corrosion-resistant cross-linked silane coatings for bioabsorbable MgZnYNd alloys before spinning with PLGA polymer. Compared with naked MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P, MgZnYNd-B-A-P shows both enhanced anti-corrosion capacity and biocompatibility for cardiovascular stent implants. Furthermore, this research initiates a new method to functionalize magnesium alloys on account of chemical conversion with in situ condensation of two kinds of silane coupling agents, during which process a foundation may be laid for barrier function establishment and further immobilization of functionalized drug-release compositions to facilitate the ongoing evolution of bioabsorbable cardiovascular stent implants.
Acknowledgments
We are thankful for the financial support from the National Basic Research Program of China (Grant No. 2012CB619102), the Shenzhen Special Funds for Development of Strategic Emerging Industries (Project No. JCYJ20130402172114948), the Shenzhen Science and Technology Funding Projects (No. CXZZ20140730091722497), and the Natural Science Foundation of Guangdong Province, China (No. 2016A030310245, No. 2016A030310244).
The authors have declared that no competing interests exist.
B.Heublein, R.Rohde, V.Kaese, M.Niemeyer, W.Hartung, A.Haverich.Heart, 89(2003), pp. 651-656
Abstract OBJECTIVES: To develop and test a new concept of the degradation kinetics of newly developed coronary stents consisting of magnesium alloys. METHODS: Design of a coronary stent prototype consisting of the non-commercial magnesium based alloy AE21 (containing 2% aluminium and 1% rare earths) with an expected 50% loss of mass within six months. Eleven domestic pigs underwent coronary implantation of 20 stents (overstretch injury). RESULTS: No stent caused major problems during implantation or showed signs of initial breakage in the histological evaluation. There were no thromboembolic events. Quantitative angiography at follow up showed a significant (p < 0.01) 40% loss of perfused lumen diameter between days 10 and 35, corresponding to neointima formation seen on histological analysis, and a 25% re-enlargement (p < 0.05) between days 35 and 56 caused by vascular remodelling (based on intravascular ultrasound) resulting from the loss of mechanical integrity of the stent. Inflammation (p < 0.001) and neointimal plaque area (p < 0.05) depended significantly on injury score. Planimetric degradation correlated with time (r = 0.67, p < 0.01). CONCLUSION: Vascular implants consisting of magnesium alloy degradable by biocorrosion seem to be a realistic alternative to permanent implants.
R.Erbel, C. DiMario, J.Bartunek, J.Bonnier, B. de Bruyne, F.R. Eberli, P. Erne, M. Haude, B. Heublein, M. Horrigan, C. Ilsley, D. Böse, J. Koolen, T.F. Lüscher, N. Weissman, R. Waksman. Lancet, 369(2007), pp. 1869-1875
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">The in vivo corrosion of magnesium alloys might provide a new mechanism which would allow degradable metal implants to be applied in musculo-skeletal surgery. This would particularly be true if magnesium alloys with controlled in vivo corrosion rates could be developed. Since the magnesium corrosion process depends on its corrosive environment, the corrosion rates of magnesium alloys under standard in vitro environmental conditions were compared to corrosion rates in an in vivo animal model. Two gravity-cast magnesium alloys (AZ91D, LAE442) were used in these investigations. Standardized immersion and electrochemical tests according to ASTM norms were performed. The in vivo corrosion tests were carried out by intramedullar implantation of sample rods of the magnesium alloys in guinea pig femura. The reduction in implant volume was determined by synchrotron-radiation-based microtomography. We found that in vivo corrosion was about four orders of magnitude lower than in vitro corrosion of the tested alloys. Furthermore, the tendency of the corrosion rates obtained from in vitro corrosion tests were in the opposite direction as those obtained from the in vivo study. The results of this study suggest, that the conclusions drawn from current ASTM standard in vitro corrosion tests cannot be used to predict in vivo corrosion rates of magnesium alloys.</p>
M.P.Staiger, A.M.Pietak, J.Huadmai, G.Dias.Biomaterials, 27(2006), pp. 1728-1734
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">As a lightweight metal with mechanical properties similar to natural bone, a natural ionic presence with significant functional roles in biological systems, and in vivo degradation <em>via</em> corrosion in the electrolytic environment of the body, magnesium-based implants have the potential to serve as biocompatible, osteoconductive, degradable implants for load-bearing applications. This review explores the properties, biological performance, challenges and future directions of magnesium-based biomaterials.</p>
M.B.Kannan, R.K.Raman.Biomaterials, 29(2008), pp. 2306-2314
The successful applications of magnesium-based alloys as degradable orthopaedic implants are mainly inhibited due to their high degradation rates in physiological environment and consequent loss in the mechanical integrity. This study examines the degradation behaviour and the mechanical integrity of calcium-containing magnesium alloys using electrochemical techniques and slow strain rate test (SSRT) method, respectively, in modified-simulated body fluid (m-SBF). Potentiodynamic polarisation and electrochemical impedance spectroscopy (EIS) results showed that calcium addition enhances the general and pitting corrosion resistances of magnesium alloys significantly. The corrosion current was significantly lower in AZ91Ca alloy than that in AZ91 alloy. Furthermore, AZ91Ca alloy exhibited a five-fold increase in the surface film resistance than AZ91 alloy. The SSRT results showed that the ultimate tensile strength and elongation to fracture of AZ91Ca alloy in m-SBF decreased only marginally (similar to 15% and 20%, respectively) in comparison with these properties in air. The fracture morphologies of the failed samples are discussed in the paper. The in vitro study suggests that calcium-containing magnesium alloys to be a promising candidate for their applications in degradable orthopaedic implants, and it is worthwhile to further investigate the in vivo corrosion behaviour of these alloys. (c) 2008 Elsevier Ltd. All rights reserved.
L.Choudhary, R.K.Raman.Acta Biomater, 8(2012), pp. 916-923
It is essential that a metallic implant material possesses adequate resistance to cracking/fracture under the synergistic action of a corrosive physiological environment and mechanical loading (i.e. stress corrosion cracking (SCC)), before the implant can be put to actual use. This paper presents a critique of the fundamental issues with an assessment of SCC of a rapidly corroding material such as magnesium alloys, and describes an investigation into the mechanism of SCC of a magnesium alloy in a physiological environment. The SCC susceptibility of the alloy in a simulated human body fluid was established by slow strain rate tensile (SSRT) testing using smooth specimens under different electrochemical conditions for understanding the mechanism of SCC. However, to assess the life of the implant devices that often possess fine micro-cracks, SCC susceptibility of notched specimens was investigated by circumferential notch tensile (CNT) testing. CNT tests also produced important design data, i.e. threshold stress intensity for SCC (K-ISCC) and SCC crack growth rate. Fractographic features of SCC were examined using scanning electron microscopy. The SSRT and CNT results, together with fractographic evidence, confirmed the SCC susceptibility of both smooth and notched specimens of a magnesium alloy in the physiological environment. (C) 2011 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">Today, more than 200 years after the first production of metallic magnesium by Sir Humphry Davy in 1808, biodegradable magnesium-based metal implants are currently breaking the paradigm in biomaterial science to develop only highly corrosion resistant metals. This groundbreaking approach to temporary metallic implants is one of the latest developments in biomaterials science that is being rediscovered. It is a challenging topic, and several secrets still remain that might revolutionize various biomedical implants currently in clinical use. Magnesium alloys were investigated as implant materials long ago. A very early clinical report was given in 1878 by the physician Edward C. Huse. He used magnesium wires as ligature for bleeding vessels. Magnesium alloys for clinical use were explored during the last two centuries mainly by surgeons with various clinical backgrounds, such as cardiovascular, musculoskeletal and general surgery. Nearly all patients benefited from the treatment with magnesium implants. Although most patients experienced subcutaneous gas cavities caused by rapid implant corrosion, most patients had no pain and almost no infections were observed during the postoperative follow-up. This review critically summarizes the in vitro and in vivo knowledge and experience that has been reported on the use of magnesium and its alloys to advance the field of biodegradable metals.</p>
B.Wang, S.Guan, J.Wang, L.Wang, S. Zhu.Mater.Sci. Eng. B, 176(2011), pp. 1673-1678
The microstructures, mechanical and corrosion properties of three extruded Mg–2Zn–0.46Y– x Nd alloys ( x 02=020.0, 0.5, 1.002wt%) were studied by optical microscopy, scanning electronic microscopy (SEM), electrochemical measurements and tensile tests. Microstructural observations indicated that Nd led to the uniformity and the variation of morphology of major second phase; tensile tests showed that Nd can improve the ductility at moderate amount (0.502wt%) and will be detrimental up to 1.0%; Mg–2Zn–0.46Y–0.5Nd alloy exhibited excellent mechanical properties ( σ b , 269.002MPa, σ 0.2 , 165.602MPa and elongation, 24%); electrochemical tests revealed that Nd can enhance the corrosion resistance and Mg–2Zn–0.46Y–1.0Nd alloy had lowest corrosion current density, which was reasoned that the line-shape and rodlike NdZn 2 phase might serve as corrosion barriers and the dissolved Nd can raise the electrode potential of the matrix.
L.Mao, G.Yuan, S.Wang, J.Niu, G.Wu, W. Ding.Mater.Lett, 88(2012), pp. 1-4
Magnesium alloys have been currently investigated as potential biodegradable implant materials. In this study, a patent Mg alloy Mg–Nd–Zn–Zr (Jiao Da BioMg, hereafter, denoted as JDBM) was researched as a potential biodegradable stent material in comparison with the clinical trial Mg alloy WE43. The corrosion behaviors of the as-extruded JDBM and WE43 were investigated in artificial plasma by in vitro degradation measurements and cyclic polarization tests. The results showed that the corrosion rate of JDBM was much lower than that of WE43 alloy. Most importantly, the JDBM alloy showed a uniform corrosion behavior in artificial plasma, which could avoid stress concentration as well as a rapid reduction in the mechanical integrity. The investigations indicated that the as-extruded JDBM alloy may be a promising implant material suitable for stent applications.
Zinc (Zn) and silver (Ag) are co-implanted into titanium by plasma immersion ion implantation. A Zn containing film with Ag nanoparticles (Ag NPs) possessing a wide size distribution is formed on the surface and the corrosion resistance is improved due to the micro-galvanic couples formed by the implanted Zn and Ag. Not only are the initial adhesion, spreading, proliferation and osteogenic differentiation of rBMSCs observed from the Zn/Ag implanted Ti in vitro, but also bacteria killing is achieved both in vitro and in vivo. Electrochemical polarization and ion release measurements suggest that the excellent osteogenic activity and antibacterial ability of the Zn/Ag co-implanted titanium are related to the synergistic effect resulting from the long-range interactions of the released Zn ions and short-range interactions of the embedded Ag NPs. The Zn/Ag co-implanted titanium offers both excellent osteogenic activity and antibacterial ability and has large potential in orthopedic and dental implants. (C) 2014 Elsevier Ltd. All rights reserved.
C.Blawert, W.Dietzel, E.Ghali, G. Song.Adv.Eng. Mater, 8(2006), pp. 511-533
[Cited within:1]
[12]
M.Abulsain, A.Berkani, F.A.Bonilla, Y.Liu, M.A.Arenas, P.Skeldon, G.E.Thompson, P.Bailey, T.C.Q.Noakes, K. Shimizu, H. Habazaki. Electrochim. Acta, 49(2004), pp. 899-904
Metastable, solid–solution Mg-0.802at.% Cu and Mg-1.402at.% Zn alloys have been anodized up to 25002V at 1002mA02cm 612 in an alkaline phosphate electrolyte at 29302K in order to investigate the enriching of alloying elements beneath the anodic films. Rutherford backscattering spectroscopy (RBS) revealed enrichments to about 4.1×10 15 Cu atoms cm 612 and 5.2×10 15 Zn atoms cm 612 , which correlate with the higher standard Gibbs free energies per equivalent for formation of copper and zinc oxides relative to that for formation of MgO. The enriched layers were of thickness about 1.5–4.002nm by medium energy ion scattering (MEIS). The anodic films, composed mainly of magnesium hydroxide, contained copper and zinc species throughout their thicknesses; the Cu:Mg and Zn:Mg atomic ratios were about 18 and 25% of those of the alloys, respectively. Phosphorus species were present in most of the film regions, with a P:Mg atomic ratio of about 0.16. The magnesium ions in the film account for about 30% of the charge passed during anodizing.
Y.Chen, G.Wan, J.Wang, S.Zhao, Y.Zhao, N. Huang.Corros.Sci, 75(2013), pp. 280-286
Magnesium-based materials are promising in biodegradable implants, but the proper corrosion control is a challenge. Phytic acid has potential as a coating on Mg to control the degradation rate. An alkaline pre-treatment process is implemented to form the hydroxyl group on the Mg surface and covalently immobilize phytic acid. The surface-immobilized phytic acid molecules chelate with Mg to form a dense and homogenous protective coating. The coated Mg exhibits a smaller corrosion current density and degradation rate as compared to bare Mg in the phosphate buffered saline.
M.D.Pereda, C.Alonso, L.Burgos-Asperilla, J.A. del Valle, O.A. Ruano, P. Perez, M.A. Fernandez Lorenzo de Mele. Acta Biomater, 6(2010), pp. 1772-1782
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">Pure Mg has been proposed as a potential degradable biomaterial to avoid both the disadvantages of non-degradable internal fixation implants and the use of alloying elements that may be toxic. However, it shows excessively high corrosion rate and insufficient yield strength. The effects of reinforcing Mg by a powder metallurgy (PM) route and the application of biocompatible corrosion inhibitors (immersion in 0.1 and 1 M KF solution treatments, 0.1 M FST and 1 M FST, respectively) were analyzed in order to improve Mg mechanical and corrosion resistance, respectively. Open circuit potential measurements, polarization techniques (PT), scanning electrochemical microscopy (SECM) and electrochemical impedance spectroscopy (EIS) were performed to evaluate its corrosion behavior. SECM showed that the local current of attacked areas decreased during the F<sup>−</sup> treatments. The corrosion inhibitory action of 0.1 M FST and 1 M FST in phosphate buffered solution was assessed by PT and EIS. Under the experimental conditions assayed, 0.1 M FST revealed better performance. X-ray photoelectron spectroscopy, energy dispersive X-ray and X-ray diffraction analyses of Mg(PM) with 0.1 M FST showed the presence of KMgF<sub>3</sub> crystals on the surface while a MgF<sub>2</sub> film was detected for 1 M FST. After fluoride inhibition treatments, promising results were observed for Mg(PM) as degradable metallic biomaterial due to its higher yield strength and lower initial corrosion rate than untreated Mg, as well as a progressive loss of the protective characteristics of the F<sup>−</sup>-containing film which ensures the gradual degradation process.</p>
X.L.Wang, N.Haraikawa, S. Suda.J.Alloy. Compd, 231(1995), pp. 397-402
It has been observed that hydriding alloys may provide a highly reactive surface for hydrogen uptake and a protective nature against impurities by treatment with a F-containing aqueous solution (also containing K). In this study, the surface composition and structure of fluorinated Mg, MgNi alloy, MgMgNi eutectic alloy and amorphous Mg(LaNi)were studied by means of X-ray photoemission spectroscopy, electron probe microanalysis and X-ray diffractometry. The samples were treated under various conditions by changing the treatment time and the ratio of sample weight to solution volume (ratio). The surface morphology, composition and chemical valence state after F-treatment were analyzed and compared with those of untreated samples. It was observed that a rose petal-shaped structure was formed and both MgFand KMgFwere formed on the surfaces of the samples after F-treatment. The formation of these compounds was independent of the initial surface state of particles, i.e. regardless of oxidized or hydroxidized or clean surface states, but dependent on treatment conditions. When the treatment time is longer or theratio is smaller, KMgFforms on the surface and co-exists with MgF. Contrary to this, only MgFforms on the surface when using a shorter treatment time or largerratio.
T.Yan, L.Tan, D.Xiong, X.Liu, B.Zhang, K. Yang.Mater.Sci. Eng. C, 30(2010), pp. 740-748
As a new class of biodegradable material, magnesium alloys have attracted much attention in recent years. In order to improve the corrosion resistance, a fluoride coating was prepared on the surface of AZ31B magnesium alloy. The surface characterization analysis showed a dense coating with some irregular pores was formed. The TF-XRD analysis indicated that the coating was mainly composed of MgO and MgF(2). Electrochemical and immersion tests proved that the fluoride conversion coating significantly improved the corrosion resistance of AZ31B. Three-point bending test revealed that the degradation behavior of the fluoride treated AZ31B could meet the requirement as a biodegradable material. (c) 2010 Elsevier B.V. All rights reserved.
Biodegradable magnesium alloys are advantageous in various implant applications, as they reduce the risks associated with permanent metallic implants. However, a rapid corrosion rate is usually a hindrance in biomedical applications. Here we report a facile two step procedure to introduce multifunctional, anti-corrosive coatings on Mg alloys, such as AZ31. The first step involves treating the NaOH-activated Mg with bistriethoxysilylethane to immobilize a layer of densely crosslinked silane coating with good corrosion resistance; the second step is to impart amine functionality to the surface by treating the modified Mg with 3-amino-propyltrimethoxysilane. We characterized the two-layer anticorrosive coating of Mg alloy AZ31 by Fourier transform infrared spectroscopy, static contact angle measurement and optical profilometry, potentiodynamic polarization and AC impedance measurements. Furthermore, heparin was covalently conjugated onto the silane-treated AZ31 to render the coating haemocompatible, as demonstrated by reduced platelet adhesion on the heparinized surface. The method reported here is also applicable to the preparation of other types of biofunctional, anti-corrosive coatings and thus of significant interest in biodegradable implant applications.
R.Pinto, M.J.Carmezim, M.G.S.Ferreira, M.F. Montemor. Prog. Org. Coat, 69(2010), pp. 143-149
The present work reports an innovative approach for surface functionalisation of magnesium alloys (WE54) containing yttrium and rare-earth (RE) elements. A two-step surface modification consisting on the formation of an anodic film followed by the deposition of a hybrid silane layer obtained by dipping in a bis-[triethoxysilylpropyl]tetrasulfide silane (BTESPT) solution was performed on the WE54 alloy. The anodic films were formed by two differentways: (i) a passive film was formed by anodic polarisation of the alloy at 0.5 V in NaOH solutions of pH 13 for 24 h and (ii) an anodic layer was formed by anodising the substrate for 1 h at a potential of 50V in the same solution. The corrosion behaviour was evaluated by electrochemical impedance spectroscopy in 0.005 M NaCl solutions for one week. The results show that the two-step treatment improves the corrosion resistance comparatively to the untreated blank substrate. The corrosion resistance of the anodised sample treated with silane is significantly better than the treatment combining passive film and silane. For this system, the overall impedance increases by three orders of magnitude.
F.Zucchi, V.Grassi, A.Frignani, C.Monticelli, G. Trabanelli.Surf.Coat. Tech, 200(2006), pp. 4136-4143
The possibility that various chemical treatments may enhance the corrosion resistance of a magnesium alloy (containing rare earths, WE43 alloy) has been investigated. Cerium(III) conversion coatings and silanic coatings, with and without acid pickling, were tested in 0.1 N Na
F.Zucchi, A.Frignani, V.Grassi, A.Balbo, G. Trabanelli.Mater.Chem. Phys, 110(2008), pp. 263-268
The coatings formed at all the pH by the lower homologue were always porous and scarcely protective. Those built by the higher homologue were more corrosion resistant and the best results were observed when octadecyl-trimethoxy-silane was hydrolyzed at pH 5: in this case a thick, scarcely defective, layer was formed and no evident corrosion attack was observed after 100002h immersion in 0.0502M Na 2 SO 4 solution.
A high school teacher invited four men from the community to meet with his class on creativity. He showed them an assortment of rocks on a table and asked for their comments.
W.J. van Ooij, D.Q. Zhu, G. Prasad, S. Jayaseelan, Y. Fu, N. Teredesai. Surf. Eng, 16(2000), pp. 386-396
[Cited within:1]
[24]
C.R.Jenney, K.M.DeFife, E. Colton, J.M. Anderson. In: Proceedings of 24th Annual Meeting of the Society for Biomaterials, San Diego, CA, (1998)
A cytokine-based, in vitro model of foreign body giant (FBGC) was utilized to examine the effect of biomaterial surface chemistry on the adhesion, motility, and fusion of monocytes and macrophages. monocytes were cultured for 10 days on 14 different -modified glass surfaces, during which time the assumed the macrophage phenotype. The adhesion of monocytes and macrophages during the culture period decreased by an average of approximately 50%, with the majority of loss observed during days 1-3. Most important, the adhesion of monocytes and macrophages was surface independent except for two surfaces containing terminal methyl groups, which decreased adhesion levels. () and () were added to the medium to induce FBGC and enhance macrophage adhesion, respectively. Surprisingly, decreased long-term monocyte/macrophage adhesion. -induced FBGC density was strongly influenced by the surface content, as determined by X-ray photoelectron spectroscopy (XPS). In contrast, contact angle and surface energy displayed no correlation with FBGC . The motility of adherent macrophages, as measured by time-lapse confocal microscopy, was not affected significantly by differences in surface chemistry or the addition of cytokines. The surface dependence of FBGC is hypothesized to be the result of varying levels of -derived surface .
M. Pagliaro. Silica-based materials for advanced chemical applications, CNR, Instituto per lo Studiodei Materiali Nanostrutturati and Institute for Scientific Methodology, Palermo, Italy, (2009)
A new class of nanosized metal–organic alloys (MORALs) has been synthesized for the first time. Silver nanoparticles doped with Cu(II) and Fe(III) phthalocyanines were thus synthesized in the presence of sodium dodecyl sulfate (SDS). The resulting materials were characterized by means of XRD, SEM, TEM coupled to energy dispersive X-ray analysis, and thermogravimetric analysis. No leaching of the photoactive dopant species was observed in water or in common organic solvents.On a synthétisé pour la première fois une nouvelle classe d’alliages métallo-organiques 00 MORALs 03 à l’échelle nano. On a ainsi synthétisé des nanoparticules d’argent dopées avec des phtalocyanines de Cu(II) et de Fe(III) en présence de sulfate de dodécyle et de sodium. Les produits obtenus ont été caractérisés par les techniques de diffraction des rayons X (DRX), de microscopie électronique à balayage (MEB) et de microscopie électronique à transmission (MET) couplées à des analyses de rayons X à dispersion d'énergie et d’analyse thermogravimétrique. On n’a observé aucune lixiviation des espèces dopantes photoréactives, que ce soit dans l’eau ou dans les solvants organiques communs.
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">It was shown recently that the deposition of thin films of tantalum and tantalum oxide enhanced the long-term biocompatibility of stainless steel biomaterials due to an increase in their corrosion resistance. In this study, we used this tantalum oxide coating as a basis for covalent immobilization of a collagen layer, which should result in a further improvement of implant tissue integration. Because of the high degradation rate of natural collagen in vivo, covalent immobilization as well as carbodiimide induced cross-linking of the protein was performed. It was found that the combination of the silane-coupling agent aminopropyl triethoxysilane and the linker molecule <em>N</em>,<em>N′</em>-disulphosuccinimidyl suberate was a very effective system for collagen immobilizing. Mechanical and enzymatic stability testing revealed a higher stability of covalent bound collagen layers compared to physically adsorbed collagen layers. The biological response induced by the surface modifications was evaluated by in vitro cell culture with human mesenchymal stem cells as well as by in vivo subcutaneous implantation into nude mice. The presence of collagen clearly improved the cytocompatibility of the stainless steel implants which, nevertheless, significantly depended on the cross-linking degree of the collagen layer.</p>
J.Song, W.J. van Ooij. J. Adhes. Sci. Technol, 17(2003), pp. 2191-2221
[Cited within:2]
[30]
L.Xu, A.Yamamoto.Colloids Surf B Biointerfaces, 93(2012), pp. 67-74
Low molecular weight (LMW) nanoparticles have attracted considerable attention as colloidal drug carriers, but when applied to intravascular drug delivery, they are easy to be removed from by the reticuloendothelial system, which limits their applications as long-circulating or target-specific carriers. Erythrocytes have a long time in the blood, but they are sometimes not suitable for loading and releasing of drug directly. The combination of LMW nanoparticles and erythrocytes that complement each other is a desirable strategy to develop a multifunctional drug carrier. In this study, monodisperse, LMW nanoparticles were prepared by ionic gelation technique and these nanoparticles were investigated with regard to their erythrocyte compatibility. Then the interactions between erythrocytes and fluorescence-labeled LMW nanoparticles were studied by confocal microscopy. The results of this study indicate that LMW nanoparticles show good compatibility with erythrocytes and they can be easily attached to the surface of erythrocyte , suggesting that erythrocytes load of LMW nanoparticles can be served as a potential vascular drug delivery system.
A.Mandelli, M.Bestetti, A. DaForno, N.Lecis, S.P.Trasatti, M. Trueba.Surf.Coat. Technol, 205(2011), pp. 4459-4465
Oxide films have been produced on AM60B magnesium alloy by micro-arc anodic oxidation in an environmentally friendly alkaline solution, with and without addition of nanoparticles (TiO
J.N.Li, P.Cao, X.N.Zhang, S.X.Zhang, Y.H. He. J.Mater.Sci, 45(2010), pp. 6038-6045
Currently available engineering magnesium alloys have several critical concerns if they are about to be used as biomaterials, particularly the concern about the toxicity of the common alloying elements such as aluminum and rare earth (RE). There is an increasing demand to develop new magnesium alloys that do not contain any toxic elements. It is also desirable, yet challenging, to develop such a material that has a controllable degradation rate in the human fluid environment. This paper presents mechanical properties, degradation, and in vitro cell attachment of a newly developed Mg-6Zn magnesium alloy. The alloy demonstrated comparable mechanical properties with typical engineering magnesium alloys. However, the bare alloy did not show an acceptable corrosion (degradation) rate. Application of a polymeric PLGA or poly(lactide- co-glycolide) coating significantly decreased the degradation rate. The results obtained from cell attachment experiments indicated that the mouse osteoblast-like MC3T3 cells could develop enhanced confluence on and interactions with the coated samples.
The Mg-Ca alloy system has been proposed as a potential new kind of degradable biomaterial with possible application within bone. Here microarc oxidation (MAO) coatings were fabricated on top of a Mg-Ca alloy using different applied voltages and the effect of applied voltage on the surface morphology and phase constitution, hydrogen evolution, pH variation in the immersion solution and in vitro biocompatibility of the MAO coating on the Mg-Ca alloy were extensively studied. It was found that the thickness and pore size of the MAO coating increased with the increasing applied voltage, whereas some micropores could be seen inside the 400 V treated MAO coating. The 360 V treated MAO coating gave the best long-term corrosion resistance during a 50 days immersion test. All the MAO coatings could promote MG63 cell adhesion, proliferation and differentiation in comparison with the uncoated Mg-Ca alloy sample, due to significantly reduced Mg ion release and pH value variations in the culture medium. After 5 days culture well-spread and elongated MG63 cells could be seen on the surface of the 360 V and 400V MAO coatings, in contrast to no cells on the uncoated Mg-Ca alloy sample. In summary, MAO showed beneficial effects on the corrosion resistance of, and thus improved cell adhesion to, the Mg-Ca alloy, and should be a good surface modification method for other biomedical magnesium alloys. (C) 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
A. Abdal-hay, M. Dewidar, J.K. Lim. Appl. Surf. Sci, 261(2012), pp. 536-546
The present study was ultimately aimed to design novel adhesive biodegradable polymer, poly(vinyl acetate) (PVAc), coatings onto Mg based alloys by the dip-coating technique in order to control the degradation rate and enhance the biocompatibility of magnesium alloys. The influence of various solvents on PVAc surface topography and their protection of Mg alloys were dramatically studied in vitro . Electrochemical polarization, degradation, and PVAc film cytocompatibility were also tested. Our results showed that the solvent had a significant effect on coating quality. PVAc/dichloromethane solution showed a porous structure and solution concentration could control the porous size. The coatings prepared using tetrahydrofuran and dimethylformamide solvents are exceptional in their ability to generate porous morphology even at low polymer concentration. In general, the corrosion performance appears to be different on different PVAc鈥搒olvent system. Immersion tests illustrated that the porous morphology on PVAc stabilized corrosion rates. A uniform corrosion attack in artificial simulation body fluid was also exhibited. The cytocompatibility of osteoblast cells (MC3T3) revealed high adherence, proliferation, and survival on the porous structure of PVAc coated Mg alloy, which was not observed for the uncoated samples. This novel PVAc coating is a promising candidate for biodegradable implant materials, which might widen the use of Mg based implants.
B.S.Kopf, A.Schipanski, M.Rottmar, S.Berner, K.Maniura-Weber.Acta Biomater, 19(2015), pp. 180-190
Early and effective integration of a metal implant into bone tissue is of crucial importance for its long-term stability. While different material properties including surface roughness and wettability but also initial blood-implant surface interaction are known to influence this osseointegration, implications of the latter process are still poorly understood. In this study, early interaction between blood and the implant surface and how this affects the mechanism of osseointegration were investigated. For this, blood coagulation on a micro-roughened hydrophobic titanium (Ti) surface (SLA-H phob ) and on a hydrophilic micro-roughened Ti surface with nanostructures (SLActive-H phil NS), as well as the effects of whole human blood pre-incubation of these two surfaces on the differentiation potential of primary human bone cells (HBC) was assessed. Interestingly, pre-incubation with blood resulted in a dense fibrin network over the entire surface on SLActive-H phil NS but only in single patches of fibrin and small isolated fibre complexes on SLA-H phob . On SLActive-H phil NS, the number of HBCs attaching to the fibrin network was greatly increased and the cells displayed enhanced cell contact to the fibrin network. Notably, HBCs displayed increased expression of the osteogenic marker proteins alkaline phosphatase and collagen-I when cultivated on both surfaces upon blood pre-incubation. Additionally, blood pre-treatment promoted an earlier and enhanced mineralization of HBCs cultivated on SLActive-H phil NS compared to SLA-H phob . The results presented in this study therefore suggest that blood pre-incubation of implant surfaces mimics a more physiological situation, eventually providing a more predictive in vitro model for the evaluation of novel bone implant surfaces.
Q.H.Zhou, J.Xie, M.Bao, H.H.Yuan, Z.Y.Ye, X.X.Lou, Y.Z. Zhang. J. Mater. Chem.B, 3(2015), pp. 4439-4450
[Cited within:1]
[37]
D.Susac, X.Sun, K.A.R.Mitchell. Appl. Surf. Sci, 207(2003), pp. 40-50
The adsorption of bis-1,2-(triethoxysilyl)ethane (BTSE) and γ-aminopropyltriethoxysilane (γ-APS) has been studied on 2024-T3 aluminum alloy samples using scanning electron microscopy (SEM) and energy dispersive X-ray (EDX) spectroscopy. The thickness and coverage of the BTSE film formed on a mechanically polished sample was found to be strongly affected by the distribution of second-phase particles on the alloy surface. Specifically, although the adsorption occurs on the alloy matrix away from the particles, and on the particles themselves, the regions immediately surrounding the particles had less BTSE. The amounts of this adsorption on an air-oxidized sample is increased on average, although the coating is non-uniform with regions (6510002μm) of higher and lower coverages. There appears to be a competition between the silane deposition and etching of the underlying oxide film for the initially oxidized sample. The adsorption of γ-APS is different with films of relatively even thickness formed over large areas of the polished sample; hydrogen bonding through the amino groups probably helps the distribution in this case.
J.Liu, X.L.Liu, T.F.Xi, C.C. Chu. J. Mater. Chem.B, 3(2015), pp. 878-893
[Cited within:4]
[39]
D.Zhu, W.J. van Ooij. Prog. Org. Coat, 49(2004), pp. 42-53
A silane surface treatment based on the water soluble mixtures of bis-[trimethoxysilylpropyl]amine and vinyltriacetoxysilane has been developed with the aim of replacing conventional chromating processes in metal-finishing industries. A variety of performance tests were employed to evaluate the anti-corrosive ability of these silane mixtures on different metals. The test results consistently demonstrated that corrosion protection performance of silane mixtures is comparable to that of chromates. In addition, a characterization study of these silane mixtures was also carried out with respect to solution stability, film thickness measurements and the effect of metal substrates on silane film structures, in order to obtain a better understanding of these silane mixtures.
G.Li, P.Yang, Y.Liao, N.Huang.Biomacromolecules, 12(2011), pp. 1155-1168
To improve the blood compatibility and endothelialization simultaneously and to ensure the long-term effectiveness of the cardiovascular implants, we developed a surface modification method, enabling the coimmobilization of biomolecules to metal surfaces. In the present study, a heparin and fibronectin mixture (Hep/Fn) covalently immobilized on a titanium (Ti) substrate for biocompatibility was investigated. Different systems [N-(3-dimethylaminopropyl)-N'-ethylcarbodiimide and N-hydroxysuccinimide, electrostatic] were used for the formation of Hep/Fn layers. Atomic force microscopy (AFM) showed that the roughness of the silanized Ti surface decreased after the immobilization of Hep/Fn. Fourier transform infrared spectroscopy (FTIR), Toluidine Blue O (TBO) test, and immunochemistry assay showed that Hep/Fn mixture was successfully immobilized on Ti surface. Blood compatibility tests (hemolysis rate, APTT, platelet adhesion, fibrinogen conformational change) showed that the coimmobilized films of Hep/Fn mixture reduced blood hemolysis rate, prolonged blood coagulation time, reduced platelets activation and aggregation, and induced less fibrinogen conformational change compared with a bare Ti surface. Endothelial cell (EC) seeding showed more EC with better morphology on pH 4 samples than on pH 7 and EDC/NHS samples, which showed rounded and aggregated cells. Systematic evaluation showed that the pH 4 samples also had much better blood compatibility. All results suggest that the coimmobilized films of Hep/Fn can confer excellent antithrombotic properties and with good endothelialization. We envisage that this method will provide a potential and effective solution for the surface modification of cardiovascular implant materials.
C.Y.Wu. University of Science and Technology Beijing. Ph.D. thesis (2011) in Chinese
[Cited within:2]
[42]
M.F.Montemor, M.G.S.Ferreira. Prog. Org. Coat, 60(2007), pp. 228-237
In this work solutions of bis-[triethoxysilylpropyl] tetrasulphide silane modified with cerium nitrate are used as pre-treatments for the AZ31 Mg alloy. The silane films formed after the pre-treatment of the metallic substrates were characterised by X-ray photoelectron spectroscopy (XPS), atomic force microscopy (AFM) and scanning electron microscopy (SEM). The films were studied before and after immersion in corrosive NaCl solutions in order to study the ageing of the silane film and to understand the role of the dopant in this process. The results demonstrate that the addition of cerium ions modifies the morphology of the surface film and leads to important changes in the film thickness and composition. Ageing in NaCl solutions decreases the film thickness and affects the chemistry of the surface film.
X.Lu, Y.Zuo, X.Zhao, Y. Tang.Corros.Sci, 60(2012), pp. 165-172
A silane film is prepared on AZ91D magnesium alloy and the effect of the silane pretreatment on the performance of a Mg-rich primer on AZ91D alloy are studied. After the silane treatment, Si–O–Mg covalent bonds form between the silane film and magnesium substrate and Si–O–Si structure forms in the silane film. As the result, the adhesion of the Mg-rich primer to AZ91D substrate increases obviously. Machu test and electrochemical measurements indicate that the silane pre-treatment significantly improves the performance of the Mg-rich primer on AZ91D alloy, which is attributed to strengthened barrier effect of the coating system.
W.Tang, X.Weng, L.Deng, K.Xu, J. Lu.Surf.Coat. Technol, 201(2007), pp. 5664-5666
Au/NiCr multi-layered metallic films used in microwave integrated circuits were deposited on Al 2 O 3 substrate by magnetron sputtering. The deformation behavior of films was investigated by nano-scratch measurement technique. It was found that deformation proceeds can be divided into three stages: at the low load regime, plastic deformation is dominant; then, the plastic/total deformation ratio decreases with the increase of load; finally, as the scratch load is increased to a critical value, film failure occurs. Experimental results indicate that plastic deformation always dominates over elastic deformation before the film failure. Most interesting is that as the load increases, the elastic and plastic curves come to intersect at the point where the elastic and plastic deformations have the same amount. It indicates that the film begins to delaminate and the critical load is determined.
D.Song, X.Jing, J.Wang, P.Yang, Y.Wang, M. Zhang.J.Alloy. Compd, 510(2012), pp. 66-70
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[46]
R.Arrabal, E.Matykina, F.Viejo, P.Skeldon, G.E.Thompson.Corros. Sci, 50(2008), pp. 1744-1752
[Cited within:1]
[47]
H.Ardelean, I.Frateur, P. Marcus.Corros.Sci, 50(2008), pp. 1907-1918
A new Ce, Zr and Nb-based conversion coating was designed for AZ91 and AM50 magnesium alloys. The corrosion protection provided by this coating was evaluated by electrochemical measurements (polarization curves, electrochemical impedance spectroscopy) in Na 2 SO 4 electrolyte, and accelerated atmospheric corrosion tests (humid, SO 2 polluted air, and salt spray). Its chemical composition was characterized by X-ray photoelectron spectroscopy (XPS). Electrochemical measurements showed that Mg alloys treated during 24h in the Ce–Zr–Nb conversion bath exhibit: (i) increased corrosion potential, (ii) decreased corrosion and anodic dissolution current densities, and (iii) increased polarization and charge transfer resistances. The accelerated corrosion tests revealed excellent atmospheric corrosion resistance for all Ce–Zr–Nb-treated samples, with or without an additional layer of epoxy–polyamide resin lacquer or paint. XPS analysis showed that the coating includes CeO 2 , Ce 2 O 3 , ZrO 2 , Nb 2 O 5 , MgO, and MgF 2 as main components. No significant modification of the chemical composition was observed after cathodic and anodic polarization in Na 2 SO 4 . This new coating provides improved corrosion resistance, and excellent paint adhesion. It offers an alternative to the chromate conversion coating for magnesium alloys.
E.M.Sherif, A.A.Almajid.Int. J. Electrochem. Sci, 6(2011), pp. 2131-2148
[Cited within:1]
[53]
Y.Cheng, T.Qin, H.Wang, Z. Zhang.Trans.Nonferrous Met. Soc. China, 19(2009), pp. 517-524
[Cited within:1]
[54]
X.Lin, X.Yang, L.Tan, M.Li, X.Wang, Y.Zhang, K.Yang, Z.Hu, J. Qiu.Appl.Surf. Sci, 288(2014), pp. 718-726
Magnesium alloys are promising biodegradable implant candidates for orthopedic application. In the present study, a phosphate-based micro-arc oxidation (MAO) coating was applied on the ZK60 alloy to decrease its initial degradation rate. Strontium (Sr) was incorporated into the coating in order to improve the bioactivity of the coating. The in vitro degradation studies showed that the MAO coating containing Sr owned a better initial corrosion resistance, which was mainly attributed to the superior inner barrier layer, and a better long-term protective ability, probably owning to its larger thickness, superior inner barrier layer and the superior apatite formation ability. The degradation of MAO coating was accompanied by the formation of degradation layer and Ca-P deposition layer. The in vitro cell tests demonstrated that the incorporation of Sr into the MAO coating enhanced both the proliferation of preosteoblast cells and the alkaline phosphatase activity of the murine bone marrow stromal cells. In conclusion, the MAO coating with Sr is a promising surface treatment for the biodegradable magnesium alloys.
L.Xu, A. Yamamoto.Appl.Surf. Sci, 258(2012), pp. 6353-6358
Magnesium (Mg) coated with four kinds of polymers, poly ( l -lactic acid) (PLLA)-high molecular weight (HMW), PLLA-low molecular weight (LMW), poly (蓻-caprolactone) (PCL)-HMW and PCL-LMW, and uncoated Mg were immersed under cell culture condition to study the degradation/corrosion behavior of the polymer-coated Mg. The releases of Mg 2+ are measured during the immersion. Surface morphology and chemical composition are observed and identified by SEM and EDX. The tomography is obtained by X-ray CT observation and degradation rate is calculated by image analysis after 10-day immersion. All kinds of polymer-coated Mg showed significantly low release of Mg 2+ ( p <0.05) in the whole immersion process comparing to that of uncoated Mg. In SEM and EDX results show, a corrosion layer can be observed on both polymer-coated and uncoated Mg after immersion. There is no obvious difference on the morphology and chemical composition of the corrosion layer between polymer-coated and uncoated Mg, indicating the corrosion/degradation process and corrosion product of Mg substrate are not changed by the polymer films under the present condition compared with uncoated Mg. Concerning the tomography and degradation rate of 10-day immersion, it can be found that the polymer-coated Mg shows a significantly low corrosion rate ( p <0.05) compared with that of uncoated Mg. PLLA coated Mg shows relatively uniform corrosion than PCL coated Mg and uncoated Mg. The largest pitting corrosion depth of PCL-LMW is about 3 times as large as the PLLA-LMW, which might be attributed to the difference of polymer microstructure. It is suggested that PLLA coating might be a suitable option for retarding the loss of mechanical properties of Mg substrate.
N.J.Ostrowski, B.Lee, A.Roy, M.Ramanathan, P.N. Kumta. J. Mater. Sci.Mater. Med, 24(2013), pp. 85-96
Polymeric film coatings were applied by dip coating on two alloy systems, AZ31 and Mg4Y, in an attempt to slow the of these alloys under in vitro conditions. Poly(lactic-co-) polymer in solution was explored at various concentrations, yielding coatings of varying thicknesses on the alloy substrates. Electrochemical corrosion studies indicate that the coatings initially provide some corrosion protection. studies showed reduced over 3 days, but beyond this time point however, do not maintain a reduction in corrosion rate. Scanning microscopy indicates inhomogeneous coating durability, with gas pocket in the polymer coating, resulting in eventual detachment from the alloy surface. In vitro studies of viability utilizing osteoblast showed improved biocompatibility of polymer coated substrates over the AZ31 and Mg4Y substrates. Results demonstrate that while challenges remain for long term control, the developed polymeric coatings nevertheless provide short term corrosion protection and improved biocompatibility of alloys for possible use in orthopedic applications.
F.Witte, V.Kaese, H.Haferkamp, E.Switzer, A.Meyer-Lindenberg, C.J.Wirth, H.Windhagen.Biomaterials, 26(2005), pp. 3557-3563
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">Degrading metal alloys are a new class of implant materials suitable for bone surgery. The aim of this study was to investigate the degradation mechanism at the bone–implant interface of different degrading magnesium alloys in bone and to determine their effect on the surrounding bone. Sample rods of four different magnesium alloys and a degradable polymer as a control were implanted intramedullary into the femora of guinea pigs. After 6 and 18 weeks, uncalcified sections were generated for histomorphologic analysis. The bone–implant interface was characterized in uncalcified sections by scanning electron microscopy (SEM), element mapping and X-ray diffraction. Results showed that metallic implants made of magnesium alloys degrade in vivo depending on the composition of the alloying elements. While the corrosion layer of all magnesium alloys accumulated with biological calcium phosphates, the corrosion layer was in direct contact with the surrounding bone. The results further showed high mineral apposition rates and an increased bone mass around the magnesium rods, while no bone was induced in the surrounding soft tissue. From the results of this study, there is a strong rationale that in this research model, high magnesium ion concentration could lead to bone cell activation.</p>
K.M.DeFife, K. Grako, G. Cruz-Aranda, S. Price, R. Chantung, K. MacPherson, R. Khoshabeh, S. Gopalan, W.G. Turnell. J. Biomater. Sci. Polym. Ed, 20(2009), pp. 1495-1511
[Cited within:1]
[60]
J.Zhou, X.Zhang, Q.Li, Y.Liu, F.Chen, L. Li.J.Mater. Chem. B, 1(2013), p. 6213
Abstract Hydrogels find widespread applications in biomedical engineering due to their hydrated environment and tunable properties (e.g., mechanical, chemical, biocompatible) similar to the native extracellular matrix (ECM). However, challenges still exist regarding utilizing hydrogels in applications such as engineering 3D tissue constructs and active targeting in drug delivery, due to the lack of controllability, actuation, and quick-response properties. Recently, magnetic hydrogels have emerged as a novel biocomposite for their active response properties and extended applications. In this review, the state-of-the-art methods for magnetic hydrogel preparation are presented and their advantages and drawbacks in applications are discussed. The applications of magnetic hydrogels in biomedical engineering are also reviewed, including tissue engineering, drug delivery and release, enzyme immobilization, cancer therapy, and soft actuators. Concluding remarks and perspectives for the future development of magnetic hydrogels are addressed.
L.Li, J.Gao, Y. Wang.Surf.Coat. Technol, 185(2004), pp. 92-98
Compared to other popular metallic biomaterials, magnesium has many advantages, which include high specific strength-to-mass ratio, non-toxicity and similar elastic modulus to that of human bone. However, the knowledge gap in corrosion resistance in physiological environment has prevented it from being a substitute for human hard tissues. In this paper, preliminary corrosion tests on magnesium specimens in Simulated Body Fluid (SBF) with and without Cl 鈭 ions have been investigated. Cytotoxicity tests were then carried out for developing a new biomaterial. The corrosion results showed that alkali and heat-treated magnesium has relatively high corrosion resistance in SBF, compared to untreated samples. Calcium-phosphate apatites were detected on the treated samples after they had been soaked in SBF for 14 days. In cytotoxicity tests, no signs of morphological changes on cells or inhibitory effect on cell growth were detected.
Composite coatings of electrostatically assembled layer-by-layer anionic and cationic polymers combined with an Mg(OH)2 surface treatment serve to provide a protective coating on AZ31 magnesium alloy substrates. These ceramic conversion coating and layer-by-layer polymeric coating combinations reduced the initial and long-term corrosion progression of the AZ31 alloy. X-ray diffraction and Fourier transform infrared spectroscopy confirmed the successful application of coatings. Potentiostatic polarization tests indicate improved initial corrosion resistance. Hydrogen evolution measurements over a 2 week period and magnesium ion levels over a 1 week period indicate longer range corrosion protection and retention of the Mg(OH)2 passivation layer in comparison to the uncoated substrates. Live/dead staining and DNA quantification were used as measures of biocompatibility and proliferation while actin staining and scanning electron microscopy were used to observe the cellular morphology and integration with the coated substrates. The coatings simultaneously provided improved biocompatibility, cellular adhesion and proliferation in comparison to the uncoated alloy surface utilizing both murine pre-osteoblast MC3T3 cells and human mesenchymal stem cells. The implementation of such coatings on magnesium alloy implants could serve to improve the corrosion resistance and cellular integration of these implants with the native tissue while delivering vital drugs or biological elements to the site of implantation.
K.Cai, A.Rechtenbach, J.Hao, J.Bossert, K.D.Jandt.Biomaterials, 26(2005), pp. 5960-5971
<h2 class="secHeading" id="section_abstract">Abstract</h2><p id="">To improve the surface biocompatibility of titanium films, a layer-by-layer (LBL) self-assembly technique, based on the polyelectrolyte-mediated electrostatic adsorption of chitosan (Chi) and gelatin (Gel), was used leading to the formation of multilayers on the titanium thin film surfaces. The film growth was initialized by deposition of one layer of positively charged poly(ethylene imine) (PEI). Then the thin film was formed by the alternate deposition of negatively charged Gel and positively charged Chi utilizing electrostatic interactions. The LBL film growth was monitored by several techniques. The chemical composition, surface topography as well as wettability were investigated by using X-ray photoelectron spectroscopy (XPS), atomic force microscopy (AFM), confocal laser scanning microscopy (CLSM) and water contact angle measurement, respectively. Quantitative XPS analysis showed the alternative change of C/N ratio after four sequential cycles coating of Ti/PEI/Gel/Chi/Gel, which indicated the discrete layer structure of coatings. Uncoated titanium (control sample) displayed a smooth surface morphology (root mean square (RMS) roughness was around 2.5 nm). A full coverage of coating with Gel/Chi layers was achieved on the titanium surface only after the deposition layers of PEI/(Gel/Chi)<sub>2</sub>. The PEI/Gel/(Chi/Gel)<sub>3</sub> layer displayed a rough surface morphology with a tree-like structure (RMS roughness is around 82 nm). These results showed that titanium films could be modified with Chi/Gel which may affect the biocompatibility of the modified titanium films. To confirm this hypothesis, cell proliferation and cell viability of osteoblasts on LBL-modified titanium films as well as control samples were investigated in vitro. The proliferation of osteoblasts on modified titanium films was found to be greater than that on control (<span id="mmlsi1" onclick="submitCitation('/science?_ob=MathURL&_method=retrieve&_eid=1-s2.0-S0142961205002541&_mathId=si1.gif&_pii=S0142961205002541&_issn=01429612&_acct=C000228598&_version=1&_userid=10&md5=0f1f9d59c6d2138343cbbd42c03ed0c3')" style="cursor:pointer;" alt="Click to view the MathML source" title="Click to view the MathML source"><span class="formulatext" title="click to view the MathML source"><em>p</em><0.05</span></span>) after 1 and 7 days culture, respectively. Cell viability measurement showed that the Chi/Gel-modified films have higher cell viability (<span id="mmlsi2" onclick="submitCitation('/science?_ob=MathURL&_method=retrieve&_eid=1-s2.0-S0142961205002541&_mathId=si2.gif&_pii=S0142961205002541&_issn=01429612&_acct=C000228598&_version=1&_userid=10&md5=d6a28f3d3c540dd43d10d91e19665d98')" style="cursor:pointer;" alt="Click to view the MathML source" title="Click to view the MathML source"><span class="formulatext" title="click to view the MathML source"><em>p</em><0.05</span></span>) than the control. These data suggest that Chi/Gel were successfully employed to surface engineer titanium via LBL technique, and enhanced its cell biocompatibility. The approach presented here may be exploited for fabrication of titanium-based implant surfaces.</p>
R.Rajaraman, D.E.Rounds, S.P.S.Yen, A. Rembaum. Exp. Cell Res, 88(1974), pp. 327-339
ABSTRACT Using normal diploid WI-38 cells, the changes in cell morphology during adhesion and flattening on to the glass surface in vitro were studied with a scanning electron microscope. The trypsinized cells were spherical to ovoid in shape with ‘bleb’-like vesicles on the surface. The process of cell adhesion and spreading consisted of four events: 1.1. attachment of cells at point of contact with the substratum;2.2. centrifugal growth of filopodia;3.3. cytoplasmic webbing;4.4. flattening of the central mass. The filopodial tips were spherical in shape with a diameter of about 800611 600 03 and were very close to the theoretical calculations of Pethica [18]. A probable mechanism responsible for the morphological events during cell adhesion is discussed.
W.Chen, T.Long, Y.J.Guo, Z.A.Zhu, Y.P. Guo. J. Mater. Chem.B, 2(2014), pp. 1653-1660
Ideal biocoatings for bone implants should be similar to the minerals of natural bones in chemical composition, crystallinity, and crystallographic texture. Herein, magnetic hydroxyapatite (HA) coatings (MHACs) with oriented nanorod arrays have been fabricated by using magnetic bioglass coatings (CaO–SiO2–P2O5–Fe3O4, MBGCs) as sacrificial templates. After the hydrothermal reaction for 24 h, the MBGCs are converted to MHACs in a simulated body fluid (SBF) via a dissolution–precipitation reaction. The formed HA nanorods with a preferential (002) orientation are perpendicular to the coating surfaces. The Fe3O4 nanoparticles in the coatings improve the nucleation rate of HA, so the elongated HA nanocrystals are retained even after hydrothermal reaction for 3 days. In contrast, if no magnetic nanoparticles are incorporated into the bioglass coatings (BGCs), the HA nanorods turn into blocky HA particles upon increasing the reaction time from 12 h to 24 h. Moreover, the MHACs possess much better hydrophilicity with a contact angle of 10.8° than the HA coatings because of the presence of Fe3O4 nanoparticles. The biocompatibility tests have been investigated by using human bone marrow stromal cells (hBMSCs) as cell models. The hBMSCs have better cell adhesion, spreading and proliferation on the MHACs than on the BGCs or MBGCs because of the HA phase, good hydrophilicity and oriented nanorod arrays. The excellent biocompatibility of the MHACs suggests that they have great potential for bone implants.
In this paper, synchronization of fuzzy modeling chaotic time delay memristor-based Chua’s circuits is presented. Based on T–S fuzzy models, not only fuzzy model of the time delay memristor-based Chua’ circuit is constructed but the fuzzy control vector can also be derived to synchronize two different time delay memristor-based Chua’s circuits. Due to the dynamical behavior with complex transient transitions of the memristor-based chaotic system which is heavily dependent on the initial state of the memristor except for the circuit parameters, the memristor-based chaotic system can generate more complex and unpredictable time domain signals. An application to chaos secure communication is used to demonstrate the effectiveness of the proposed chaotic synchronization scheme.
C.H.Ye, Y.F.Zheng, S.Q.Wang, T.F.Xi, Y.D.Li.Appl. Surf. Sci, 258(2012), pp. 3420-3427
Phytic acid (PA) conversion coating on WE43 magnesium alloy was prepared by the method of immersion. The influences of phytic acid solution with different pH on the microstructure, properties of the conversion coating and the corrosion resistance were investigated by SEM, FTIR and potentiodynamic polarization method. Furthermore, the biocompatibility of different pH phytic acid solution modified WE43 magnesium alloys was evaluated by MTT and hemolysis test. The results show that PA can enhance the corrosion resistance of WE43 magnesium especially when the pH value of modified solution is 5 and the cytotoxicity of the PA coated WE43 magnesium alloy is much better than that of the bare WE43 magnesium alloy. Moreover, all the hemolysis rates of the PA coated WE43 Mg alloy were lower than 5%, indicating that the modified Mg alloy met the hemolysis standard of biomaterials. Therefore, PA coating is a good candidate to improve the biocompatibility of WE43 magnesium alloy.
X.Wang, X.Chen, L.Xing, C.Mao, H.Yu, J. Shen.J.Mater. Chem. B, 1(2013), p. 5036
[Cited within:1]
[70]
N.Kong, K.Lin, H.Li, J. Chang.J.Mater. Chem. B, 2(2014), p. 1100
Copper (Cu) has been reported to be able to stimulate vascularization/angiogenesis, which is critical for regeneration of vascularized tissue in tissue engineering. Silicate bioceramics have also been reported to have stimulatory effects on vascularization due to the silicon (Si) ions released from silicate biomaterials. Therefore, we hypothesize that a combination of Cu and Si ions may show synergy effects on vascularization. Therefore, a copper-doped calcium silicate bioceramic (Cu-CaSiO3, Cu-CS) was designed and synthesized with the purpose to enhance the stimulatory effects of copper salts or pure silicate bioceramics on vascularization by combining the effects of Cu and Si ions. The cytocompatibility of Cu-CS was firstly assessed by testing the influence of Cu-CS ion extracts on proliferation of human umbilical vein endothelial cells (HUVECs). Thereafter, vascularization of HUVECs on ECMatrix鈩 gel or co-cultured with human dermal fibroblasts (HDFs) in Cu-CS extracts was evaluated and expression of angiogenic growth factors was analyzed. Results revealed that, as compared to CS extracts and media containing soluble CuSO4, Cu-CS extracts possessed stronger stimulatory effects on upregulation of angiogenic growth factors, which finally resulted in better stimulatory effects on vascularization. During the vascularization process, paracrine effects dominated in the co-culture system. In addition, lower concentrations of Cu and Si ions released from Cu-CS than those released from pure CS or CuSO4 were enough to stimulate vascularization, which indicated that there were synergy effects between Cu and Si ions during stimulation of vascularization by Cu-CS. Taken together, the designed Cu-CS may be suitable as a new biomaterial for regenerating blood vessels in tissue engineering.
T.Nakaki, M.Nakayama, R. Kato.Eur.J. Pharmacol, 189(1990), pp. 347-353
Effects of nitric oxide (NO) and NO-producing vasodilators such as glyceryl trinitrate and sodium nitroprusside were tested on DNA synthesis in the clonal rat aortic smooth muscle cells, RACS-1. DNA synthesis was estimated by [3H]thymidine incorporation to DNA. NO and NO-producing vasodilators inhibited the DNA synthesis that was induced by 10% fetal calf serum. NO and NO-producing vasodilators also inhibited the basal level of DNA synthesis that occurred possibly as a result of autocrine mechanisms. NO-producing vasodilators also inhibited the fetal calf serum-induced proliferation of cells. Sodium nitroprusside inhibited the endothelin-mediated DNA synthesis. In another mesenchymal cell line, Chinese hamster fibroblast V79 cells, NO and NO-producing vasodilators failed to inhibit DNA synthesis, excluding the possibility of general cell toxicity. An exposure to NO and NO-producing vasodilators resulted in an increase of cyclic GMP (cGMP) content in the RACS-1 cells. A cGMP analog, 8-bromo-cGMP, inhibited DNA synthesis in the RACS-1 cells. These results suggest that EDRF/nitric oxide released from endothelium possibly contributes to inhibition of the DNA synthesis in vascular smooth muscle cells.
Y.Y.Zheng, C.D.Xiong, X.Y.Li, L.F.Zhang.Appl. Surf. Sci, 320(2014), pp. 93-101
Preclinical studies suggest that dopamine D3 receptor (D3R) antagonists are promising for the treatment of drug abuse and addiction. However, few D3R antagonists have potential to be tested in humans due to short half-life, toxicity or limited preclinical research into pharmacotherapeutic efficacy. Here, we report on a novel D3R antagonist YQA14, which has improved half-life and pharmacokinetic profile and which displays potent pharmacotherapeutic efficacy in attenuating cocaine reward and relapse to drug-seeking behavior. Electrical brain-stimulation reward (BSR) in laboratory animals is a highly sensitive experimental approach to evaluate a drug's rewarding effects. We found that cocaine (2mg/kg) significantly enhanced electrical BSR in rats (i.e., decreased stimulation threshold for BSR), while YQA14 alone had no effect on BSR. Pretreatment with YQA14 significantly and dose-dependently attenuated cocaine-enhanced BSR. YQA14 also facilitated extinction from drug-seeking behavior in rats during early behavioral extinction, and attenuated cocaine- or contextual cue-induced relapse to drug-seeking behavior. YQA14 alone did not maintain self-administration in either na茂ve rats or in rats experienced at cocaine self-administration. YQA14 also inhibited expression of repeated cocaine-induced behavioral sensitization. These findings suggest that YQA14 may have pharmacotherapeutic potential in attenuating cocaine-taking and cocaine-seeking behavior. Thus, YQA14 deserves further investigation as a promising agent for treatment of cocaine addiction.
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
1
2007
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
1
2006
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
1
2006
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
1
2008
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
1
2012
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
2
2010
... In the recent years, magnesium alloy has been widely studied as a potential material in bioabsorbable implants, such as scaffolds for bone repair and fixation, novel cardiovascular stents, and very promising data were reported[1] and [2]. The magnesium-based cardiovascular stents can be absorbed in in vivo environment after the completion of the expectant functions, which are beneficial as they can significantly lower the risk of long-term complications compared with permanent implants. The benefits include avoiding painful secondary removal surgery, foreign-body induced inflammatory response, displacement, and restenosis of the stent, delayed type hypersensitivity and late lumen loss [3] and [4]. In light of this, magnesium-based alloy have become the most popular bioabsorbable implants materials because of their biodegradability, proper mechanical properties, biocompatibility and function in bone formation enhancement [5], [6] and [7]. ...
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
2
2011
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
... Every reagent was used as received. BTSE, APTES, phosphate-buffered saline (PBS), glutaraldehyde, Triton X-100, Acid Orange 7, FITC-phalloidin, and tetrazolium-8 (WST-8) were from Sigma-Aldrich, China; DAPI was purchased from CST, China; Griess reagent Kit was from Beyotime, China. PLGA (Mw = 100,000, GA:LA ratio = 25:75) was ordered from Daigang Biology, China. As-extruded MgZnYNd alloy bars (Mg-2Zn-0.5Y-0.5Nd)[8] were provided by Guan's lab at Zhengzhou University, China. Hank's solution (containing 8.00 g NaCl, 1.00 g glucose, 0.40 g KCl, 0.35 g NaHCO3, 0.14 g CaCl2, 0.06 g KH2PO4, 0.06 g MgSO4, 0.06 g Na2HPO4, 0.01 g MgCl2 per 1000 mL deionized water) was freshly self-prepared and pH value was modulated to 7.40 (37.0 ± 0.5 °C) prior to the tests. ...
1
2012
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
2014
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
2006
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
2004
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
2013
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
2010
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
1
1995
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
2
2010
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
... Each of MgZnYNd sample exhibited an increase from 7.4 to over 8.0 initially over the first 3 immersion days. The bare MgZnYNd and MgZnYNd-P samples showed an over-high value of 10.4 by the end of testing period. However, both the MgZnYNd-A-P and MgZnYNd-B-A-P groups could restrain the Hank's solution from becoming over alkaline to avoid overpassing local tissue receptivity, which was much lower than the MgZnYNd-P group with a slight steady increase over the whole period, and reached a steady pH value at about 10.0. The data were lower than the surface-modified magnesium alloys with other methods. For examples, after immersing for 15 days in Hank's solution, AZ91D modified by stabilization-hydrothermal process showed the pH of over 10.0[60]; pure magnesium with alkali-heat treating showed pH value of 10.1 after only 14-day immersion in SBF[61]; after immersing for 30 days in SBF, fluoride-treated AZ31 exhibited pH value of 10.2[16]; at the end of 30-day immersion in Hank's, MAO-modified Mg-Ca showed an over-high pH of 10.5[33]. ...
1
2010
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
5
2013
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
... Bistriethoxysilylethane (BTSE) and 3-amino-propyltriethoxysilane (APTES) are extensively researched bissilane and monosilane, respectively, providing functional groups that could further attach drug-loading layer or bioactive molecules as well as enhance the interfacial reactions with metallic surfaces[27]. However, the single silane layer will preferentially bond to surfaces of metal materials and result in occasional flaws in the silane layers, permitting solvent to permeate into the silane-metal interface, while much denser cross-linked three-dimensional network and further stronger interfacial adhesion with metal surfaces can be achieved with ‘bis-silanes’, i.e., metal-silane A-silane-B[28]. To make use of both types of silanes, a facile two-step 3-amino-propyltrimethoxysilane (γ-APS) pre-treatment was designed to provide fine anti-corrosion protection for steel and aluminum[29] and AZ31[18]. Further, as direct coating of PLGA layer onto metal surface could only form invalid bond, stronger hydrogen bond between exposed amino group of APTES and PLGA coating is another reason we adopted this two-step silane coating to improve the anti-corrosion ability and biocompatibility of PLGA protective layer on MgZnYNd alloy for cardiovascular stent. To our knowledge, the thorough study of the biocompatibility of the bilayer system has not been reported before. ...
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
... The static contact angle (CA) measurements could reflect changes of surface wettability on the modified MgZnYNd alloys, thus confirming successful bonding of silane with the presence of characteristic groups as shown in Fig. 4. The CA of the original bare MgZnYNd sample was 64.02° ± 3.42°, and increased to 101.09° ± 8.14° with the hydrophobic BTSE coating, which was due to the cross-linked Si-O-Si structure, and was consistent with the reported data[18]. While with a lower CA of 55.18° ± 0.87°, MgZnYNd-B-A was more hydrophilic, attributing to the superficial amine groups. However, subsequent coating of PLGA layer reproduced the surface hydrophobicity, with the CA value increasing to 66.39° ± 2.90°, which was closer to a pure PLGA with less influence of the underneath MgZnYNd substrate[38]. ...
... EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53]. ...
2
2010
... Magnesium alloys usually degrade fast in solutions containing chloride ion, such as physiological environments, thus the poor anti-corrosion ability has been the main hindrance for potential clinical applications of Mg alloys. Further, corrosion of magnesium alloys is inevitably accompanied by hydrogen evolution, and pH changes will also affect cell adhere onto the implant surface and damage surrounding tissues[7]. Therefore, an applicable corrosion rate should be obtained as essential requirement before taking Mg alloys as biodegradable metallic implants for clinical application. Various methods have been developed to improve the anti-corrosion ability of Mg alloys, including the manufacture of new Mg alloys, such as MgZnYNd alloy[8], MgNdZnZr alloy[9], etc., conversion surface coating by nitrogen ion implantation[10], anodization[11] and [12], alkaline treatment[13], fluoride treatments[14], [15] and [16], and silanization[17], [18] and [19]. ...
... In order to understand the type the interface bonding exists in the cross-linked silane film on the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P surfaces, XPS spectra were scanned. The whole and narrow scan spectra of O 1s and Si 2p were analyzed using Casa XPS software and shown in Fig. 3(a, b), elemental compositions were calculated from XPS intensities and denoted as atomic percentage (at.%) in Table 1. The O 1s binding energy values in the silane film ( Fig. 3(b)) peaked at 533.3, 532.4, and 531.4 eV and were attributed to C-O-C or Si-O-Si [41], [42] and [43], Si-O-Mg [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
1
2006
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
2
2008
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
1
1991
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
1
2000
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
1
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
2
2009
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
... [25]. ...
1
2004
... Among these technologies, silane-based conversion coatings for magnesium alloys have been confirmed to be valid, eco-friendly and economical[20] and [21]. Silanes refer to a series of silicon-based organic-inorganic liquids with general chemical formula of R1(CH2)nSi(OR2)3, in which R1 stands for an organo-functional group and R2 for a hydrolysable alkoxy group. Thus, Si(OR)3 could hydrolyze to yield silanol groups Si(OH)3 which could bond with hydrated metal surfaces metal-OH with the linking of Si-O-metal bonds[22] when react with water, while the silanol groups Si(OR)3 process self-cross-linking by forming siloxane bonds (Si-O-Si), resulting in a cross-linked anti-corrosion layer which bonds to the metal sample via chemical bonding[23]. Silane-based coatings have been certified to have excellent biocompatibility reflected in good cellular and bacterial adhesion behavior, proper biological interactions, and protein absorption [24] and [25]. It has been confirmed to degrade into single molecule of Si(OH)4 and would be removed by the urinary system[26] without adverse tissue reactions[25]. ...
2
2005
... Bistriethoxysilylethane (BTSE) and 3-amino-propyltriethoxysilane (APTES) are extensively researched bissilane and monosilane, respectively, providing functional groups that could further attach drug-loading layer or bioactive molecules as well as enhance the interfacial reactions with metallic surfaces[27]. However, the single silane layer will preferentially bond to surfaces of metal materials and result in occasional flaws in the silane layers, permitting solvent to permeate into the silane-metal interface, while much denser cross-linked three-dimensional network and further stronger interfacial adhesion with metal surfaces can be achieved with ‘bis-silanes’, i.e., metal-silane A-silane-B[28]. To make use of both types of silanes, a facile two-step 3-amino-propyltrimethoxysilane (γ-APS) pre-treatment was designed to provide fine anti-corrosion protection for steel and aluminum[29] and AZ31[18]. Further, as direct coating of PLGA layer onto metal surface could only form invalid bond, stronger hydrogen bond between exposed amino group of APTES and PLGA coating is another reason we adopted this two-step silane coating to improve the anti-corrosion ability and biocompatibility of PLGA protective layer on MgZnYNd alloy for cardiovascular stent. To our knowledge, the thorough study of the biocompatibility of the bilayer system has not been reported before. ...
... Coupled the data of cell adhesion, cell cytotoxicity and NO release showed above, it is obvious that the bilayer BTSE-APTES pre-treatment benefits a lot when coating MgZnYNd with PLGA, which is reflected in the ability of MgZnYNd-B-A-P to significantly enhance ECV304 and VSMC viability over MgZnYNd-A-P and MgZnYNd-P. Further, the difference in the chemical composition plus the further physical barrier functionality of cross-linked silane coating between MgZnYNd-B-A-P and MgZnYNd-A-P should result to the total different responses for both types of cells. It was reported by Zheng et al. that MC3T3-E1 attachment, stretching, and proliferation were effectively enhanced on GRGD-modified PEEK surface by tailored silanization technique[75]. Besides, silane-coupling agent aminopropyl triethoxysilane was used as a very effective system for collagen immobilizing onto stainless steel to improve human mesenchymal stem cells adhesion and proliferation[27]. ...
2
2003
... Bistriethoxysilylethane (BTSE) and 3-amino-propyltriethoxysilane (APTES) are extensively researched bissilane and monosilane, respectively, providing functional groups that could further attach drug-loading layer or bioactive molecules as well as enhance the interfacial reactions with metallic surfaces[27]. However, the single silane layer will preferentially bond to surfaces of metal materials and result in occasional flaws in the silane layers, permitting solvent to permeate into the silane-metal interface, while much denser cross-linked three-dimensional network and further stronger interfacial adhesion with metal surfaces can be achieved with ‘bis-silanes’, i.e., metal-silane A-silane-B[28]. To make use of both types of silanes, a facile two-step 3-amino-propyltrimethoxysilane (γ-APS) pre-treatment was designed to provide fine anti-corrosion protection for steel and aluminum[29] and AZ31[18]. Further, as direct coating of PLGA layer onto metal surface could only form invalid bond, stronger hydrogen bond between exposed amino group of APTES and PLGA coating is another reason we adopted this two-step silane coating to improve the anti-corrosion ability and biocompatibility of PLGA protective layer on MgZnYNd alloy for cardiovascular stent. To our knowledge, the thorough study of the biocompatibility of the bilayer system has not been reported before. ...
... After NaOH passivation, MgZnYNd surface was covered by Mg(OH)2 film, and reactions existing in anode and cathode are Mg-2e-→Mg2+Mg-2e-→Mg2+and Mg2++2H2O+2e-→Mg(OH)2+H2Mg2++2H2O+2e-→Mg(OH)2+H2, respectively. As anodic Mg dissolves beneath the Mg(OH)2 layer on the MgZnYNd surface, Mg2+ transport determines the anodic kinetics; while charge transferring relates the cathodic reaction can occur both underlying and over the silane film, leading to a higher cathodic current density of bare MgZnYNd than the bilayer silane groups. After one-layer pretreating with APTES, only a limited retardant silane molecule layer formed on the surface; thus the Icorr of MgZnYNd-A-P decreased by only 3.45 µA/cm2 compared with MgZnYNd-P. However, after BTSE coating, the exchange current density results from the cathodic reaction is greatly decreased because the hydrophobic silane film could retard electron transport and water penetration, thus forming a physical barrier. The anodic reaction is also slowed as the bilayer silane retards the Mg2+ transport[28]. Meanwhile, the Si-O-Mg bonding at the MgZnYNd-BTSE interface could also block anodic reactions[39], which lowers the MgZnYNd-B-A-P corrosion current value by an order of magnitude than MgZnYNd-A-P. ...
2
2003
... Bistriethoxysilylethane (BTSE) and 3-amino-propyltriethoxysilane (APTES) are extensively researched bissilane and monosilane, respectively, providing functional groups that could further attach drug-loading layer or bioactive molecules as well as enhance the interfacial reactions with metallic surfaces[27]. However, the single silane layer will preferentially bond to surfaces of metal materials and result in occasional flaws in the silane layers, permitting solvent to permeate into the silane-metal interface, while much denser cross-linked three-dimensional network and further stronger interfacial adhesion with metal surfaces can be achieved with ‘bis-silanes’, i.e., metal-silane A-silane-B[28]. To make use of both types of silanes, a facile two-step 3-amino-propyltrimethoxysilane (γ-APS) pre-treatment was designed to provide fine anti-corrosion protection for steel and aluminum[29] and AZ31[18]. Further, as direct coating of PLGA layer onto metal surface could only form invalid bond, stronger hydrogen bond between exposed amino group of APTES and PLGA coating is another reason we adopted this two-step silane coating to improve the anti-corrosion ability and biocompatibility of PLGA protective layer on MgZnYNd alloy for cardiovascular stent. To our knowledge, the thorough study of the biocompatibility of the bilayer system has not been reported before. ...
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
1
2012
... The corresponding 4% wt/vol PLGA solution was dissolved with dichloromethane at room temperature. PLGA films on MgZnYNd substrates were prepared by spin coating[30]: 100 µL PLGA solution was dropped onto MgZnYNd-B-A sample, the spin process was set to be 300 rpm for 6 s then 7000 rpm for 20 s, both sides of substrates were prepared similarly before drying for 48 h at 37 °C. Samples with only APTES treatment before PLGA coating and samples with direct PLGA layer were likewise prepared for comparison and donated as MgZnYNd-A-P and MgZnYNd-P, respectively. ...
2
2011
... To evaluate the adhesion enhancement of the PLGA layer on MgZnYNd substrates by BTSE-APTES pre-treatment, nano-scratch resistance tests were conducted using a Nano Indenter system (Hysitron, USA) with a spherical diamond Rockwell indenter. The tip approached and loaded into the bi-layer coating with increasing rate of 30 µN/s up to 1000 µN, making a 10 µm scratching at the prescribed direction. The Nano Indenter system could detect the first cracks of the coating corresponding to the critical load[31]. The tests were performed in clean environment at 25 °C. For statistical purpose, at least three parallel results were deemed valid and average values were calculated and reported. ...
... Nano-indentation is a popular technique to detect micro-scale mechanical properties of surfaces; particularly it offers an efficient and simple means to estimate the scratch-resistance of the coated films[31] and [44]. The initial coating failure could indicate the critical lateral force in the ramp-load scratch steps and is in positive correlation with the adhesion strength of the coating, and the data are shown in Fig. 5. Given this, the critical lateral forces for MgZnYNd-B-A-P was significantly higher than the MgZnYNd-P and MgZnYNd-A-P group, i.e., 435 µN at 27 s vs. 300 µN at as early as 25 s and 320 µN at as early as 22 s for MgZnYNd-P and MgZnYNd-A-P, respectively, a 45% and 36% higher critical lateral force of MgZnYNd-B-A-P than MgZnYNd-P and MgZnYNd-A-P. While both the force and time was necessary to consider when analyzing the nano-scratch resistance data. These lateral force/time data indicate overall the much stronger adhesion between PLGA and MgZnYNd substrates can be influenced by two-step BTSE-APTES pre-treatment than the one-step APTES pre-treated coating or direct PLGA coating. The stronger adhesion between coating and substrate existed in MgZnYNd-B-A-P system can be attributed to strong covalent bond between substrate and silane layer, besides, the intermolecular hydrogen bond and Van der Waals force between silane layer and PLGA could be deduced from the chemical formula of APTES. The critical lateral force for MgZnYNd-B-A-P group (435 µN at 27 s) was a little higher than the published research by Liu et al., phenylalanine-based poly (ester amide)s (8-Phe-4) films coated magnesium substrate[38], which is 380 µN at 25 s - 403 µN at 26 s, respectively, for 4% and 2% 8-Phe-4 coatings. ...
1
2010
... Standard methods were applied in electrochemical measurement[32] and [33]. CHENHUA CHI650C system (Shanghai, China) with a three-electrode cell was used, in which MgZnYNd samples acted as working electrode, saturated calomel electrode (SCE) as the reference electrode, and the platinum as the counter electrode. Both potentiodynamic polarization curves analysis (Tafel plot) and electrochemical impedance spectroscopy (EIS) were performed in 125 mL Hank's solution at 37.0 ± 0.5 °C with working area of 0.332 cm2. For EIS measurements, the frequency varied from 105 Hz to 10-2 Hz after reaching a steady open-circuit potential value, and spectra were analyzed through Nyquist plots. The Tafel plot was tested at a rate of 0.5 mV ⋅ s-1 from -600 mV cathodically to +800 mV anodically relative to the open circuit potential (OCP), and the Ecorr and Icorr were obtained by fitting Tafel plot with the Corrview software according to the criterion ASTM-G102-89. For statistical purpose, at least three parallel results were deemed valid and average values were calculated and reported. ...
2
2011
... Standard methods were applied in electrochemical measurement[32] and [33]. CHENHUA CHI650C system (Shanghai, China) with a three-electrode cell was used, in which MgZnYNd samples acted as working electrode, saturated calomel electrode (SCE) as the reference electrode, and the platinum as the counter electrode. Both potentiodynamic polarization curves analysis (Tafel plot) and electrochemical impedance spectroscopy (EIS) were performed in 125 mL Hank's solution at 37.0 ± 0.5 °C with working area of 0.332 cm2. For EIS measurements, the frequency varied from 105 Hz to 10-2 Hz after reaching a steady open-circuit potential value, and spectra were analyzed through Nyquist plots. The Tafel plot was tested at a rate of 0.5 mV ⋅ s-1 from -600 mV cathodically to +800 mV anodically relative to the open circuit potential (OCP), and the Ecorr and Icorr were obtained by fitting Tafel plot with the Corrview software according to the criterion ASTM-G102-89. For statistical purpose, at least three parallel results were deemed valid and average values were calculated and reported. ...
... Each of MgZnYNd sample exhibited an increase from 7.4 to over 8.0 initially over the first 3 immersion days. The bare MgZnYNd and MgZnYNd-P samples showed an over-high value of 10.4 by the end of testing period. However, both the MgZnYNd-A-P and MgZnYNd-B-A-P groups could restrain the Hank's solution from becoming over alkaline to avoid overpassing local tissue receptivity, which was much lower than the MgZnYNd-P group with a slight steady increase over the whole period, and reached a steady pH value at about 10.0. The data were lower than the surface-modified magnesium alloys with other methods. For examples, after immersing for 15 days in Hank's solution, AZ91D modified by stabilization-hydrothermal process showed the pH of over 10.0[60]; pure magnesium with alkali-heat treating showed pH value of 10.1 after only 14-day immersion in SBF[61]; after immersing for 30 days in SBF, fluoride-treated AZ31 exhibited pH value of 10.2[16]; at the end of 30-day immersion in Hank's, MAO-modified Mg-Ca showed an over-high pH of 10.5[33]. ...
1
2012
... The SEM results (Fig. 2(a)) show that BTSE and BTSE-APTES pre-treated MgZnYNd substrate (MgZnYNd-B and MgZnYNd-B-A) exhibited a flat surface with only dark fringe, slight remnants of the polishing grooves from the final 2000# SiC abrasive paper, while PLGA coated MgZnYNd-B-A samples (MgZnYNd-B-A-P) exhibited a non-porous, continuous, dense and smooth coating (Fig. 2(a)). Generally, groove surface on the naked MgZnYNd substrate could not be totally covered by the nano dimension of silane layer thickness; however, the outer layer of PLGA coating masked the grooves very well. Besides, the direct PLGA-coated samples (MgZnYNd-P) showed acceptable smooth surface with a few polishing grooves. It has been reported that corrosion behavior of polymer coated Mg-based materials and cell behavior on it could be affected by film morphologies (porous or dense structure)[34], [35] and [36], while the smooth coating of MgZnYNd-B-A-P in the present study could well fit the anti-corrosive and biocompatible requirements. Moreover, for the cross-section of MgZnYNd-B-A-P, an obvious double layer could be distinguished (Fig. 2(a)), with the coating thickness of 6.4 ± 0.4 µm and 0.5 ± 0.4 µm for PLGA layer and silane coating, respectively. Besides, no significant difference was observed referring to the depth of PLGA layers between MgZnYNd-B-A-P and MgZnYNd-P groups (6.4 ± 0.4 µm vs 6.3 ± 0.8 µm as labeled in red in Fig. 2(a)). While the total coating thickness of MgZnYNd-B-A-P (about 6.9 µm) compared to MgZnYNd-P (about 6.3 µm) could explain the surface morphology difference discussed above. ...
1
2015
... The SEM results (Fig. 2(a)) show that BTSE and BTSE-APTES pre-treated MgZnYNd substrate (MgZnYNd-B and MgZnYNd-B-A) exhibited a flat surface with only dark fringe, slight remnants of the polishing grooves from the final 2000# SiC abrasive paper, while PLGA coated MgZnYNd-B-A samples (MgZnYNd-B-A-P) exhibited a non-porous, continuous, dense and smooth coating (Fig. 2(a)). Generally, groove surface on the naked MgZnYNd substrate could not be totally covered by the nano dimension of silane layer thickness; however, the outer layer of PLGA coating masked the grooves very well. Besides, the direct PLGA-coated samples (MgZnYNd-P) showed acceptable smooth surface with a few polishing grooves. It has been reported that corrosion behavior of polymer coated Mg-based materials and cell behavior on it could be affected by film morphologies (porous or dense structure)[34], [35] and [36], while the smooth coating of MgZnYNd-B-A-P in the present study could well fit the anti-corrosive and biocompatible requirements. Moreover, for the cross-section of MgZnYNd-B-A-P, an obvious double layer could be distinguished (Fig. 2(a)), with the coating thickness of 6.4 ± 0.4 µm and 0.5 ± 0.4 µm for PLGA layer and silane coating, respectively. Besides, no significant difference was observed referring to the depth of PLGA layers between MgZnYNd-B-A-P and MgZnYNd-P groups (6.4 ± 0.4 µm vs 6.3 ± 0.8 µm as labeled in red in Fig. 2(a)). While the total coating thickness of MgZnYNd-B-A-P (about 6.9 µm) compared to MgZnYNd-P (about 6.3 µm) could explain the surface morphology difference discussed above. ...
1
2015
... The SEM results (Fig. 2(a)) show that BTSE and BTSE-APTES pre-treated MgZnYNd substrate (MgZnYNd-B and MgZnYNd-B-A) exhibited a flat surface with only dark fringe, slight remnants of the polishing grooves from the final 2000# SiC abrasive paper, while PLGA coated MgZnYNd-B-A samples (MgZnYNd-B-A-P) exhibited a non-porous, continuous, dense and smooth coating (Fig. 2(a)). Generally, groove surface on the naked MgZnYNd substrate could not be totally covered by the nano dimension of silane layer thickness; however, the outer layer of PLGA coating masked the grooves very well. Besides, the direct PLGA-coated samples (MgZnYNd-P) showed acceptable smooth surface with a few polishing grooves. It has been reported that corrosion behavior of polymer coated Mg-based materials and cell behavior on it could be affected by film morphologies (porous or dense structure)[34], [35] and [36], while the smooth coating of MgZnYNd-B-A-P in the present study could well fit the anti-corrosive and biocompatible requirements. Moreover, for the cross-section of MgZnYNd-B-A-P, an obvious double layer could be distinguished (Fig. 2(a)), with the coating thickness of 6.4 ± 0.4 µm and 0.5 ± 0.4 µm for PLGA layer and silane coating, respectively. Besides, no significant difference was observed referring to the depth of PLGA layers between MgZnYNd-B-A-P and MgZnYNd-P groups (6.4 ± 0.4 µm vs 6.3 ± 0.8 µm as labeled in red in Fig. 2(a)). While the total coating thickness of MgZnYNd-B-A-P (about 6.9 µm) compared to MgZnYNd-P (about 6.3 µm) could explain the surface morphology difference discussed above. ...
3
2003
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
... [37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
... ], [37] and [39]. ...
4
2015
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
... The static contact angle (CA) measurements could reflect changes of surface wettability on the modified MgZnYNd alloys, thus confirming successful bonding of silane with the presence of characteristic groups as shown in Fig. 4. The CA of the original bare MgZnYNd sample was 64.02° ± 3.42°, and increased to 101.09° ± 8.14° with the hydrophobic BTSE coating, which was due to the cross-linked Si-O-Si structure, and was consistent with the reported data[18]. While with a lower CA of 55.18° ± 0.87°, MgZnYNd-B-A was more hydrophilic, attributing to the superficial amine groups. However, subsequent coating of PLGA layer reproduced the surface hydrophobicity, with the CA value increasing to 66.39° ± 2.90°, which was closer to a pure PLGA with less influence of the underneath MgZnYNd substrate[38]. ...
... Nano-indentation is a popular technique to detect micro-scale mechanical properties of surfaces; particularly it offers an efficient and simple means to estimate the scratch-resistance of the coated films[31] and [44]. The initial coating failure could indicate the critical lateral force in the ramp-load scratch steps and is in positive correlation with the adhesion strength of the coating, and the data are shown in Fig. 5. Given this, the critical lateral forces for MgZnYNd-B-A-P was significantly higher than the MgZnYNd-P and MgZnYNd-A-P group, i.e., 435 µN at 27 s vs. 300 µN at as early as 25 s and 320 µN at as early as 22 s for MgZnYNd-P and MgZnYNd-A-P, respectively, a 45% and 36% higher critical lateral force of MgZnYNd-B-A-P than MgZnYNd-P and MgZnYNd-A-P. While both the force and time was necessary to consider when analyzing the nano-scratch resistance data. These lateral force/time data indicate overall the much stronger adhesion between PLGA and MgZnYNd substrates can be influenced by two-step BTSE-APTES pre-treatment than the one-step APTES pre-treated coating or direct PLGA coating. The stronger adhesion between coating and substrate existed in MgZnYNd-B-A-P system can be attributed to strong covalent bond between substrate and silane layer, besides, the intermolecular hydrogen bond and Van der Waals force between silane layer and PLGA could be deduced from the chemical formula of APTES. The critical lateral force for MgZnYNd-B-A-P group (435 µN at 27 s) was a little higher than the published research by Liu et al., phenylalanine-based poly (ester amide)s (8-Phe-4) films coated magnesium substrate[38], which is 380 µN at 25 s - 403 µN at 26 s, respectively, for 4% and 2% 8-Phe-4 coatings. ...
... In addition, as shown on the 30th day EDS graph and the inserted table (Fig. 7(b)), the MgZnYNd-P group had more precipitation and crystallization containing O, Ca and P elements during the 30-day immersion than the MgZnYNd-B-A-P group, which were also reflected in SEM images with a few spherical and cylindrical crystals. Furthermore, the calculated total Ca and P elements percentage (wt%) was 40.03% from the 30th day EDS spectrum, indicating that apatite crystals probably formed on the PLGA-direct coated samples. Semblable pattern of hydroxyapatite (HA) crystals also formed on the upper surface of PLGA-coated Mg[38]. When it was used as vascular stents, the delamination, cracks and deposited apatite on surface of Mg material would destroy blood homeostasis as reported by Witte et al.[58] that aggregation of erythrocytes would be aroused by apatite particles, further leading to an over-high blood viscosity and clotting[59]. Contrary to the MgZnYNd-P sample, less apatite existed on surface of the MgZnYNd-B-A-P with smaller Ca and P peak value (with a total of 26.4 wt% of Ca and P elements). This different degradation morphology and composition between MgZnYNd-B-A-P and MgZnYNd-P was attributed to the compactness of cross-linking layer and strong chemical bond between PLGA and MgZnYNd substrates, thus forming effective physical barrier to improve the anti-corrosion ability of MgZnYNd alloy. ...
2
2004
... FT-IR was employed to confirm the successful grafting of the polysiloxane coatings on metal surface. Fig. 2(b) shows the full spectra of MgZnYNd-B, MgZnYNd-B-A, MgZnYNd-B-A-P and bare MgZnYNd samples. Peaks around 1045 cm-1 indicate asymmetric stretching of -Si-O- exists in -Si-O-Si-[29] and [37]. The following APTES cross-linked coating markedly increased the -Si-O- asymmetric stretching intensities, showing a large peak around 1590 cm-1 attributed to the protonated amino groups[37]. The C = O stretching peaks at 1750 cm-1, C-H bending peaks at 1500 cm-1 and C-O-C stretching peaks at 1080 cm-1 were all shown on MgZnYNd-B-A-P samples, while it didn't show the characteristic peaks of -Si-O- asymmetric stretching and amino bands observed in the MgZnYNd-B or MgZnYNd-B-A, and was consistent with previous spectra[38]. Moreover, no such characteristic bands showed on naked MgZnYNd substrates as expected. Besides, our FT-IR results are similar with previously reported bistriethoxysilylethane (BTSE) and 3-amino-propyltrimethoxysilane (γ-APS, with the same -NH2 group as APTES) treated Mg, Al, Fe, and Mg based materials[18], [21], [37] and [39]. ...
... After NaOH passivation, MgZnYNd surface was covered by Mg(OH)2 film, and reactions existing in anode and cathode are Mg-2e-→Mg2+Mg-2e-→Mg2+and Mg2++2H2O+2e-→Mg(OH)2+H2Mg2++2H2O+2e-→Mg(OH)2+H2, respectively. As anodic Mg dissolves beneath the Mg(OH)2 layer on the MgZnYNd surface, Mg2+ transport determines the anodic kinetics; while charge transferring relates the cathodic reaction can occur both underlying and over the silane film, leading to a higher cathodic current density of bare MgZnYNd than the bilayer silane groups. After one-layer pretreating with APTES, only a limited retardant silane molecule layer formed on the surface; thus the Icorr of MgZnYNd-A-P decreased by only 3.45 µA/cm2 compared with MgZnYNd-P. However, after BTSE coating, the exchange current density results from the cathodic reaction is greatly decreased because the hydrophobic silane film could retard electron transport and water penetration, thus forming a physical barrier. The anodic reaction is also slowed as the bilayer silane retards the Mg2+ transport[28]. Meanwhile, the Si-O-Mg bonding at the MgZnYNd-BTSE interface could also block anodic reactions[39], which lowers the MgZnYNd-B-A-P corrosion current value by an order of magnitude than MgZnYNd-A-P. ...
1
2011
... AO test was used to quantify amine concentrations on the surface of MgZnYNd-B-A, shown in Fig. 2(c). MgZnYNd-B-A showed a markedly higher (p > 0.05) amino concentration of 97.17 ± 6.52 nmol/cm2 when compared with the blank control (bare MgZnYNd alloys substrates), indicating that the APTES was successfully coated onto MgZnYNd-B with appropriate amount of amino groups. The data were comparable with the TiOH (pristine Ti activated by NaOH) silanized by γ-amino-propyltriethoxysilane (γ-APS), which is about 127 nmol/cm2[40]. However, few amino groups were detected on the surface of MgZnYNd-B-A-P samples because of the coverage effect of PLGA layer. ...
2
2011
... In order to understand the type the interface bonding exists in the cross-linked silane film on the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P surfaces, XPS spectra were scanned. The whole and narrow scan spectra of O 1s and Si 2p were analyzed using Casa XPS software and shown in Fig. 3(a, b), elemental compositions were calculated from XPS intensities and denoted as atomic percentage (at.%) in Table 1. The O 1s binding energy values in the silane film ( Fig. 3(b)) peaked at 533.3, 532.4, and 531.4 eV and were attributed to C-O-C or Si-O-Si [41], [42] and [43], Si-O-Mg [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
... [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
4
2007
... In order to understand the type the interface bonding exists in the cross-linked silane film on the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P surfaces, XPS spectra were scanned. The whole and narrow scan spectra of O 1s and Si 2p were analyzed using Casa XPS software and shown in Fig. 3(a, b), elemental compositions were calculated from XPS intensities and denoted as atomic percentage (at.%) in Table 1. The O 1s binding energy values in the silane film ( Fig. 3(b)) peaked at 533.3, 532.4, and 531.4 eV and were attributed to C-O-C or Si-O-Si [41], [42] and [43], Si-O-Mg [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
... ] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
... [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
... [42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
2
2012
... In order to understand the type the interface bonding exists in the cross-linked silane film on the MgZnYNd-B, MgZnYNd-B-A, and MgZnYNd-B-A-P surfaces, XPS spectra were scanned. The whole and narrow scan spectra of O 1s and Si 2p were analyzed using Casa XPS software and shown in Fig. 3(a, b), elemental compositions were calculated from XPS intensities and denoted as atomic percentage (at.%) in Table 1. The O 1s binding energy values in the silane film ( Fig. 3(b)) peaked at 533.3, 532.4, and 531.4 eV and were attributed to C-O-C or Si-O-Si [41], [42] and [43], Si-O-Mg [41] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
... ] and [43] and Mg(OH)2[19] and [42], respectively. The Si 2p could be observed at 102.06 eV, corresponding to the Si-O-Mg group [42], peaked at 102.8 eV indicated the bonding energy value for Si-O-Si[42]. Therefore, the binding energy values shown in XPS spectra for both O 1s and Si 2p verify the existence of the Si-O-Mg structure, indicating that the effective covalent bonds formed between Mg substrate and the silane layer. Moreover, Si-O-Si is the main constitute of the silane layer, resulting from the cross-linking reaction of SiOH. On the contrary, few above characteristic peaks were detected on the surface of MgZnYNd-B-A-P sample. ...
1
2007
... Nano-indentation is a popular technique to detect micro-scale mechanical properties of surfaces; particularly it offers an efficient and simple means to estimate the scratch-resistance of the coated films[31] and [44]. The initial coating failure could indicate the critical lateral force in the ramp-load scratch steps and is in positive correlation with the adhesion strength of the coating, and the data are shown in Fig. 5. Given this, the critical lateral forces for MgZnYNd-B-A-P was significantly higher than the MgZnYNd-P and MgZnYNd-A-P group, i.e., 435 µN at 27 s vs. 300 µN at as early as 25 s and 320 µN at as early as 22 s for MgZnYNd-P and MgZnYNd-A-P, respectively, a 45% and 36% higher critical lateral force of MgZnYNd-B-A-P than MgZnYNd-P and MgZnYNd-A-P. While both the force and time was necessary to consider when analyzing the nano-scratch resistance data. These lateral force/time data indicate overall the much stronger adhesion between PLGA and MgZnYNd substrates can be influenced by two-step BTSE-APTES pre-treatment than the one-step APTES pre-treated coating or direct PLGA coating. The stronger adhesion between coating and substrate existed in MgZnYNd-B-A-P system can be attributed to strong covalent bond between substrate and silane layer, besides, the intermolecular hydrogen bond and Van der Waals force between silane layer and PLGA could be deduced from the chemical formula of APTES. The critical lateral force for MgZnYNd-B-A-P group (435 µN at 27 s) was a little higher than the published research by Liu et al., phenylalanine-based poly (ester amide)s (8-Phe-4) films coated magnesium substrate[38], which is 380 µN at 25 s - 403 µN at 26 s, respectively, for 4% and 2% 8-Phe-4 coatings. ...
1
2012
... Tafel plots for various MgZnYNd samples are revealed in Fig. 6(b). From the extrapolation of the Tafel plots, the corrosion potential (Ecorr) and corrosion current density (Icorr) were calculated and shown in Table 2. The BTSE-APTES pre-treatment on MgZnYNd substrates led to Ecorr values (-0.495 V) less negative than the MgZnYNd-P (-1.568 V), MgZnYNd-A-P (-1.550 V), and uncoated substrates (-1.556 V), implying retard of corrosion. Moreover, the reduction in Icorr values by 77.60%, 48.70%, and 31.81% in the MgZnYNd-B-A-P than the bare MgZnYNd sample, MgZnYNd-P, and MgZnYNd-A-P, respectively, was also companied by the higher Ecorr values of the BTSE-APTES pre-treatment on MgZnYNd. Conversely, MgZnYNd substrates with direct PLGA layer led to a more negative Ecorr, implying a less effective barrier function against corrosion provided by the PLGA without silane pre-treatment. The similar trend of Ecorr increases from -1.614 V to -1.490 V was also observed in Mg-Li alloy with assembled ZSM-5 layer adopting silane coupling agent as linkage[45]. ...
1
2008
... The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50]. ...
1
2008
... The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50]. ...
1
2008
... The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50]. ...
2
2001
... The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50]. ...
... EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53]. ...
2
2015
... The Icorr values of the MgZnYNd-B-A-P (7.148 µA/cm2) were, however, only a small fraction (51.30%) of the MgZnYNd-P samples (13.935 µA/cm2), and were much lower than Mg alloys AZ91 and WE43 with phosphate PEO coating (40 µA/cm2 and 30 µA/cm2, respectively)[46] and conversion coatings composed of niobium, zirconium and cerium on AZ91 and AM50 alloys (13-54 µA/cm2)[47]. As reported by Cao[48], both Icorr and Ecorr are the decisive factors that characterize corrosion rate. Hence, we may predict that PLGA coating after BTSE-APTES pre-treatment owns a much better potential to protect Mg from over-fast corrosion than direct PLGA coating. The conclusion that more negative Ecorr corresponds to a faster corrosion rate was also reported in other published studies, for example, corrosion rate of pure Mg in a sulphate medium by Baril et al.[49] and MgF2/polydopamine-coated MgZnYNd alloy corrosion rate in DMEM solution by Liu et al.[50]. ...
... NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P). ...
1
2010
... EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53]. ...
1
2011
... EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53]. ...
1
2009
... EIS analysis was analyzed by Nyquist plots for MgZnYNd-B-A-P, MgZnYNd-A-P, MgZnYNd-P, and bare MgZnYNd substrate shown in Fig. 6(c). The Zre values, equals Z' when Z” is 0, which is an essential indicator in electrochemical impedance studies, higher Zre is in positive correlation to anti-corrosion ability and data are listed in Table 2. Both MgZnYNd-B-A-P (118,750 Ω ⋅ cm2) and MgZnYNd-A-P (27, 800 Ω ⋅ cm2) values were significantly higher than the naked MgZnYNd (1860 Ω ⋅ cm2), verifying the larger electrochemical impedance of PLGA-coated MgZnYNd after pre-treated with BTSE-APTES silane, thus providing much super ability in retarding Mg alloy from over-quick corrosion. The Zre value of MgZnYNd-B-A-P was much larger than that reported for BTSE-γ-APS-coated AZ31 magnesium alloy (nearly 16,000 Ω ⋅ cm2)[18] and the value of hydrofluoric acid treated AZ31 (about 3000 Ω ⋅ cm2) in 3.5% NaCl solution for various time[51]. Generally, the MgZnYNd-B-A-P group (pink triangle shown in Fig. 6(c)) exhibited a much larger capacitive loop than the MgZnYNd-A-P (blue triangle shown in Fig. 6(c)), while both MgZnYNd-P (red dots shown in Fig. 6(c)) and MgZnYNd (black squares shown in Fig. 6(c)) showed capacitive loops with even smaller diameter. As the anti-corrosion ability is correlated to the enlargement of capacitive loops, we may conclude a super ability of MgZnYNd-B-A-P in corrosion resistance. The above analysis of EIS plot is reported to be the most widely adopted interpretation [49] and [52], even the explanation for EIS curve of Mg alloys has never reached an agreement, with only a few other editions reported[53]. ...
1
2014
... Complying with the current recommended criterion in ASTM G31-72[54] and [55], various MgZnYNd alloy substrates were immersed in the Hank's solution for 30 days at 37 °C to quantify the silane coatings effect on the corrosion rate of the MgZnYNd substrate. ...
1
2013
... Complying with the current recommended criterion in ASTM G31-72[54] and [55], various MgZnYNd alloy substrates were immersed in the Hank's solution for 30 days at 37 °C to quantify the silane coatings effect on the corrosion rate of the MgZnYNd substrate. ...
2
2012
... SEM results (Fig. 7(a)) with EDS spectra (Fig. 7(b)) of the samples after immersing for 15 and 30 days are shown in Fig. 7. The MgZnYNd-P group showed severe delamination and cracks (labeled by red dotted circle) at both the 15th and 30th day data. Similar images of delamination were also reported on poly (l-lactic acid) (PLLA) and poly (ε-caprolactone) (PCL) modified Mg[56] in DMEM solution, and some bubbles effect formed on PLGA-coated Mg4Y and AZ31 alloy after a 3-day incubation[57]. Contrarily, MgZnYNd-B-A-P samples showed uniform and compact surface with only a few superficial narrow cracks and delamination but no gas bubbles, as it escaped from the narrow cracks before it could gather, indicating better anti-corrosion ability of PLGA layer after pre-treatment with BTSE-APTES. ...
... The released Mg2+ concentration from various MgZnYNd samples after immersion over 30 days in Hank's solution was measured by the ICP-OES, which is also an essential indicator in judging the anti-corrosion ability of the Mg-based biomaterials. Data were shown in Fig. 8(b). With the same trend shown in pH value, both MgZnYNd-A-P and MgZnYNd-B-A-P coatings reduced the Mg2+ released more effectively than the MgZnYNd and MgZnYNd-P, indicating an effective barrier functions from the silane coatings. Most importantly, over the whole 30-day immersion period, the MgZnYNd-B-A-P (about 7.77 ppm at 15th day, 83.67 ppm at 30th day) showed much less released Mg2+ compared with the MgZnYNd-P (35.97 ppm at 15th day, 105.27 ppm at 30th day), i.e., a 18.0%-55.8% reduction in the MgZnYNd-B-A-P sample. Besides, when compared with the reported studies, the MgZnYNd-B-A-P released less Mg2+: the stabilization-hydrothermal treated AZ91D shows 80 µmol mL-1 cm-2 Mg2+ after 15-day immersion[60], the accumulated Mg2+ after a 7-day incubation for the LBL PEI-PCL-PAH treated AZ31 is 13 mmol[62], the cumulative Mg2+ value is 80-100 ppm for PLLA/PCL-coated magnesium after 10-day immersion[56]. These tested Mg2+ concentration value trend is in accordance with the recorded degradation surface morphology images and pH data shown above (Fig. 7 and Fig. 8(a)). ...
1
2013
... SEM results (Fig. 7(a)) with EDS spectra (Fig. 7(b)) of the samples after immersing for 15 and 30 days are shown in Fig. 7. The MgZnYNd-P group showed severe delamination and cracks (labeled by red dotted circle) at both the 15th and 30th day data. Similar images of delamination were also reported on poly (l-lactic acid) (PLLA) and poly (ε-caprolactone) (PCL) modified Mg[56] in DMEM solution, and some bubbles effect formed on PLGA-coated Mg4Y and AZ31 alloy after a 3-day incubation[57]. Contrarily, MgZnYNd-B-A-P samples showed uniform and compact surface with only a few superficial narrow cracks and delamination but no gas bubbles, as it escaped from the narrow cracks before it could gather, indicating better anti-corrosion ability of PLGA layer after pre-treatment with BTSE-APTES. ...
1
2005
... In addition, as shown on the 30th day EDS graph and the inserted table (Fig. 7(b)), the MgZnYNd-P group had more precipitation and crystallization containing O, Ca and P elements during the 30-day immersion than the MgZnYNd-B-A-P group, which were also reflected in SEM images with a few spherical and cylindrical crystals. Furthermore, the calculated total Ca and P elements percentage (wt%) was 40.03% from the 30th day EDS spectrum, indicating that apatite crystals probably formed on the PLGA-direct coated samples. Semblable pattern of hydroxyapatite (HA) crystals also formed on the upper surface of PLGA-coated Mg[38]. When it was used as vascular stents, the delamination, cracks and deposited apatite on surface of Mg material would destroy blood homeostasis as reported by Witte et al.[58] that aggregation of erythrocytes would be aroused by apatite particles, further leading to an over-high blood viscosity and clotting[59]. Contrary to the MgZnYNd-P sample, less apatite existed on surface of the MgZnYNd-B-A-P with smaller Ca and P peak value (with a total of 26.4 wt% of Ca and P elements). This different degradation morphology and composition between MgZnYNd-B-A-P and MgZnYNd-P was attributed to the compactness of cross-linking layer and strong chemical bond between PLGA and MgZnYNd substrates, thus forming effective physical barrier to improve the anti-corrosion ability of MgZnYNd alloy. ...
1
2009
... In addition, as shown on the 30th day EDS graph and the inserted table (Fig. 7(b)), the MgZnYNd-P group had more precipitation and crystallization containing O, Ca and P elements during the 30-day immersion than the MgZnYNd-B-A-P group, which were also reflected in SEM images with a few spherical and cylindrical crystals. Furthermore, the calculated total Ca and P elements percentage (wt%) was 40.03% from the 30th day EDS spectrum, indicating that apatite crystals probably formed on the PLGA-direct coated samples. Semblable pattern of hydroxyapatite (HA) crystals also formed on the upper surface of PLGA-coated Mg[38]. When it was used as vascular stents, the delamination, cracks and deposited apatite on surface of Mg material would destroy blood homeostasis as reported by Witte et al.[58] that aggregation of erythrocytes would be aroused by apatite particles, further leading to an over-high blood viscosity and clotting[59]. Contrary to the MgZnYNd-P sample, less apatite existed on surface of the MgZnYNd-B-A-P with smaller Ca and P peak value (with a total of 26.4 wt% of Ca and P elements). This different degradation morphology and composition between MgZnYNd-B-A-P and MgZnYNd-P was attributed to the compactness of cross-linking layer and strong chemical bond between PLGA and MgZnYNd substrates, thus forming effective physical barrier to improve the anti-corrosion ability of MgZnYNd alloy. ...
2
2013
... Each of MgZnYNd sample exhibited an increase from 7.4 to over 8.0 initially over the first 3 immersion days. The bare MgZnYNd and MgZnYNd-P samples showed an over-high value of 10.4 by the end of testing period. However, both the MgZnYNd-A-P and MgZnYNd-B-A-P groups could restrain the Hank's solution from becoming over alkaline to avoid overpassing local tissue receptivity, which was much lower than the MgZnYNd-P group with a slight steady increase over the whole period, and reached a steady pH value at about 10.0. The data were lower than the surface-modified magnesium alloys with other methods. For examples, after immersing for 15 days in Hank's solution, AZ91D modified by stabilization-hydrothermal process showed the pH of over 10.0[60]; pure magnesium with alkali-heat treating showed pH value of 10.1 after only 14-day immersion in SBF[61]; after immersing for 30 days in SBF, fluoride-treated AZ31 exhibited pH value of 10.2[16]; at the end of 30-day immersion in Hank's, MAO-modified Mg-Ca showed an over-high pH of 10.5[33]. ...
... The released Mg2+ concentration from various MgZnYNd samples after immersion over 30 days in Hank's solution was measured by the ICP-OES, which is also an essential indicator in judging the anti-corrosion ability of the Mg-based biomaterials. Data were shown in Fig. 8(b). With the same trend shown in pH value, both MgZnYNd-A-P and MgZnYNd-B-A-P coatings reduced the Mg2+ released more effectively than the MgZnYNd and MgZnYNd-P, indicating an effective barrier functions from the silane coatings. Most importantly, over the whole 30-day immersion period, the MgZnYNd-B-A-P (about 7.77 ppm at 15th day, 83.67 ppm at 30th day) showed much less released Mg2+ compared with the MgZnYNd-P (35.97 ppm at 15th day, 105.27 ppm at 30th day), i.e., a 18.0%-55.8% reduction in the MgZnYNd-B-A-P sample. Besides, when compared with the reported studies, the MgZnYNd-B-A-P released less Mg2+: the stabilization-hydrothermal treated AZ91D shows 80 µmol mL-1 cm-2 Mg2+ after 15-day immersion[60], the accumulated Mg2+ after a 7-day incubation for the LBL PEI-PCL-PAH treated AZ31 is 13 mmol[62], the cumulative Mg2+ value is 80-100 ppm for PLLA/PCL-coated magnesium after 10-day immersion[56]. These tested Mg2+ concentration value trend is in accordance with the recorded degradation surface morphology images and pH data shown above (Fig. 7 and Fig. 8(a)). ...
1
2004
... Each of MgZnYNd sample exhibited an increase from 7.4 to over 8.0 initially over the first 3 immersion days. The bare MgZnYNd and MgZnYNd-P samples showed an over-high value of 10.4 by the end of testing period. However, both the MgZnYNd-A-P and MgZnYNd-B-A-P groups could restrain the Hank's solution from becoming over alkaline to avoid overpassing local tissue receptivity, which was much lower than the MgZnYNd-P group with a slight steady increase over the whole period, and reached a steady pH value at about 10.0. The data were lower than the surface-modified magnesium alloys with other methods. For examples, after immersing for 15 days in Hank's solution, AZ91D modified by stabilization-hydrothermal process showed the pH of over 10.0[60]; pure magnesium with alkali-heat treating showed pH value of 10.1 after only 14-day immersion in SBF[61]; after immersing for 30 days in SBF, fluoride-treated AZ31 exhibited pH value of 10.2[16]; at the end of 30-day immersion in Hank's, MAO-modified Mg-Ca showed an over-high pH of 10.5[33]. ...
1
2013
... The released Mg2+ concentration from various MgZnYNd samples after immersion over 30 days in Hank's solution was measured by the ICP-OES, which is also an essential indicator in judging the anti-corrosion ability of the Mg-based biomaterials. Data were shown in Fig. 8(b). With the same trend shown in pH value, both MgZnYNd-A-P and MgZnYNd-B-A-P coatings reduced the Mg2+ released more effectively than the MgZnYNd and MgZnYNd-P, indicating an effective barrier functions from the silane coatings. Most importantly, over the whole 30-day immersion period, the MgZnYNd-B-A-P (about 7.77 ppm at 15th day, 83.67 ppm at 30th day) showed much less released Mg2+ compared with the MgZnYNd-P (35.97 ppm at 15th day, 105.27 ppm at 30th day), i.e., a 18.0%-55.8% reduction in the MgZnYNd-B-A-P sample. Besides, when compared with the reported studies, the MgZnYNd-B-A-P released less Mg2+: the stabilization-hydrothermal treated AZ91D shows 80 µmol mL-1 cm-2 Mg2+ after 15-day immersion[60], the accumulated Mg2+ after a 7-day incubation for the LBL PEI-PCL-PAH treated AZ31 is 13 mmol[62], the cumulative Mg2+ value is 80-100 ppm for PLLA/PCL-coated magnesium after 10-day immersion[56]. These tested Mg2+ concentration value trend is in accordance with the recorded degradation surface morphology images and pH data shown above (Fig. 7 and Fig. 8(a)). ...
1
2005
... To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend. ...
1
1974
... To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend. ...
1
2014
... To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend. ...
1
1962
... To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend. ...
1
2015
... To assess biocompatibility of biomaterials, cellular response and behavior are essential considerations[63]. The process of cellular adhesion followed by spreading happened in the following procedure: cellular attachment, filopodial stretching, cytoplasmic webbing, cell mass flattening, and peripheral cytoplasm ruffling[64]. Actin staining of EA. hy926 and VSMC after reseeding for 24 h and 48 h is shown in Fig. 9. It is seen from the figure that the long bundles of green fibers consisting of actin filaments certified the favorable cytoskeleton morphology of the adhered EA. hy926 and VSMC cells. Generally, for both EA. hy926 and VSMC, obviously more cells tended to adhere on MgZnYNd-B-A-P surface than the bare MgZnYNd, MgZnYNd-P, and MgZnYNd-A-P groups, indicating that MgZnYNd-B-A-P was more favorable in terms of biocompatibility. Cell morphology of EA. hy926 adhered on the bare MgZnYNd substrate was poorly defined without strong stained actin fiber: elongated, triangular and irregular for both 24 h and 48 h, verifying that EA. hy926 were weakly adhered to the bare MgZnYNd, and adhered cellular number even declined after cultivation for 48 h than that for 24 h. Although cells on MgZnYNd-P and MgZnYNd-A-P were of more normal morphology with acceptable stained actin fiber but the attached cell numbers were no more than the bare MgZnYNd group, with the same trend of cell number declining from 24 h to 48 h. However, EA. hy926 on MgZnYNd-B-A-P fully spread out, showing intense actin filaments stretching into various orientations and reset cytoskeleton with strong actin fibers[65], indicating that the MgZnYNd-B-A-P had more super cellular spreading and adhesion capability. Without sufficient passivation layer of cross-linked silane coating, it was deduced that the irregular cell morphology and decreased cell numbers on the MgZnYNd-P and MgZnYNd-A-P group resulted from the over-fast magnesium corrosion: the bulk degradation mode of PLGA was a self-acceleration process, leading to the acid micro-environment; further, PLGA degradation permitted acidic degradation products to contact and cause corrosion of magnesium substrate, resulting in alkali atmosphere, and the local over-alkaline micro-environment caused by magnesium corrosion could be the main hazard contributing to the decreased cell compatibility[66] and [67]. Moreover, the single layer of APTES silane with occasional flaws could not establish effective physical barrier. On the contrary, the cross-linked silane coating formed solid isolation zone against water, thus preventing corrosion products contacting and injuring cells. While the data for VSMC cells showed the same trend. ...
1
2012
... The cell viabilities of both EA. hy926 (Fig. 10(a)) and VSMC (Fig. 10(b)) grown on the surface of the various MgZnYNd samples for 1, 3, 5 days were tested by the CCK-8 assay, and results are shown in Fig. 10. Viabilities of both EA. hy926 and VSMC seeded on the surface of MgZnYNd-B-A-P, MgZnYNd-A-P, and MgZnYNd-P samples were significantly improved with the cell viabilities of 96.79%-86.83%, 90.20%-62.66%, and 73.85%-51.36%, respectively, for the 5-day incubation period, in contrast with the bare MgZnYNd substrates (ranging 46.11%-20.27%). Therefore, MgZnYNd with silane and/or PLGA coating could all significantly increase the viabilities compared to the bare MgZnYNd. However, for EA. hy926, MgZnYNd-A-P, and MgZnYNd-P groups showed obvious continuous trend of decline with the incubation time, and reached viabilities of 63.00% and 57.45%, respectively. According to the standard ISO 10993-5 2009, viability between 50.00%-79.00% is considered to be Grade II toxicity, which is not acceptable for clinical use. On the contrary, MgZnYNd-B-A-P group showed much higher viabilities than the MgZnYNd-A-P and MgZnYNd-P, which increased slightly during the whole incubation period with the viability of about 96.79% at the 5th day, indicating the good cytocompatibility of BTSE-APTES bilayer, confirming the potential of more remarkable effect in promoting endothelialization. While for VSMC, the data showed significant advantages of MgZnYNd-B-A-P over both MgZnYNd-A-P and MgZnYNd-P, i.e., 6.59%-15.30% higher than MgZnYNd-A-P and 30.23%-35.51% higher than MgZnYNd-P. Besides, for all the groups with silane or PLGA coatings, the 3rd day viability reached to a peak but declined since then. The gradual reduction in cell viability after 3rd day culture in MgZnYNd-B-A-P was probably due to its potential in prevention of restenosis. When referring to other published studies, the cell viability data of the MgZnYNd-B-A-P group for both ECV304 and VSMC were much higher, for example, the L929 cell viabilities cultured for 4 days in the extracts of phytic acid coated WE43 are 28%-70%[68], the HEK293 cell viabilities seeded on the surfaces of BMS-Br and BMS-g-HPBBEA are 68% and 76%, respectively[69], the 7-day HUVECs proliferation data cultured with calcium silicate extracts ranged between 30% and 70%[70]. ...
1
2013
... The cell viabilities of both EA. hy926 (Fig. 10(a)) and VSMC (Fig. 10(b)) grown on the surface of the various MgZnYNd samples for 1, 3, 5 days were tested by the CCK-8 assay, and results are shown in Fig. 10. Viabilities of both EA. hy926 and VSMC seeded on the surface of MgZnYNd-B-A-P, MgZnYNd-A-P, and MgZnYNd-P samples were significantly improved with the cell viabilities of 96.79%-86.83%, 90.20%-62.66%, and 73.85%-51.36%, respectively, for the 5-day incubation period, in contrast with the bare MgZnYNd substrates (ranging 46.11%-20.27%). Therefore, MgZnYNd with silane and/or PLGA coating could all significantly increase the viabilities compared to the bare MgZnYNd. However, for EA. hy926, MgZnYNd-A-P, and MgZnYNd-P groups showed obvious continuous trend of decline with the incubation time, and reached viabilities of 63.00% and 57.45%, respectively. According to the standard ISO 10993-5 2009, viability between 50.00%-79.00% is considered to be Grade II toxicity, which is not acceptable for clinical use. On the contrary, MgZnYNd-B-A-P group showed much higher viabilities than the MgZnYNd-A-P and MgZnYNd-P, which increased slightly during the whole incubation period with the viability of about 96.79% at the 5th day, indicating the good cytocompatibility of BTSE-APTES bilayer, confirming the potential of more remarkable effect in promoting endothelialization. While for VSMC, the data showed significant advantages of MgZnYNd-B-A-P over both MgZnYNd-A-P and MgZnYNd-P, i.e., 6.59%-15.30% higher than MgZnYNd-A-P and 30.23%-35.51% higher than MgZnYNd-P. Besides, for all the groups with silane or PLGA coatings, the 3rd day viability reached to a peak but declined since then. The gradual reduction in cell viability after 3rd day culture in MgZnYNd-B-A-P was probably due to its potential in prevention of restenosis. When referring to other published studies, the cell viability data of the MgZnYNd-B-A-P group for both ECV304 and VSMC were much higher, for example, the L929 cell viabilities cultured for 4 days in the extracts of phytic acid coated WE43 are 28%-70%[68], the HEK293 cell viabilities seeded on the surfaces of BMS-Br and BMS-g-HPBBEA are 68% and 76%, respectively[69], the 7-day HUVECs proliferation data cultured with calcium silicate extracts ranged between 30% and 70%[70]. ...
1
2014
... The cell viabilities of both EA. hy926 (Fig. 10(a)) and VSMC (Fig. 10(b)) grown on the surface of the various MgZnYNd samples for 1, 3, 5 days were tested by the CCK-8 assay, and results are shown in Fig. 10. Viabilities of both EA. hy926 and VSMC seeded on the surface of MgZnYNd-B-A-P, MgZnYNd-A-P, and MgZnYNd-P samples were significantly improved with the cell viabilities of 96.79%-86.83%, 90.20%-62.66%, and 73.85%-51.36%, respectively, for the 5-day incubation period, in contrast with the bare MgZnYNd substrates (ranging 46.11%-20.27%). Therefore, MgZnYNd with silane and/or PLGA coating could all significantly increase the viabilities compared to the bare MgZnYNd. However, for EA. hy926, MgZnYNd-A-P, and MgZnYNd-P groups showed obvious continuous trend of decline with the incubation time, and reached viabilities of 63.00% and 57.45%, respectively. According to the standard ISO 10993-5 2009, viability between 50.00%-79.00% is considered to be Grade II toxicity, which is not acceptable for clinical use. On the contrary, MgZnYNd-B-A-P group showed much higher viabilities than the MgZnYNd-A-P and MgZnYNd-P, which increased slightly during the whole incubation period with the viability of about 96.79% at the 5th day, indicating the good cytocompatibility of BTSE-APTES bilayer, confirming the potential of more remarkable effect in promoting endothelialization. While for VSMC, the data showed significant advantages of MgZnYNd-B-A-P over both MgZnYNd-A-P and MgZnYNd-P, i.e., 6.59%-15.30% higher than MgZnYNd-A-P and 30.23%-35.51% higher than MgZnYNd-P. Besides, for all the groups with silane or PLGA coatings, the 3rd day viability reached to a peak but declined since then. The gradual reduction in cell viability after 3rd day culture in MgZnYNd-B-A-P was probably due to its potential in prevention of restenosis. When referring to other published studies, the cell viability data of the MgZnYNd-B-A-P group for both ECV304 and VSMC were much higher, for example, the L929 cell viabilities cultured for 4 days in the extracts of phytic acid coated WE43 are 28%-70%[68], the HEK293 cell viabilities seeded on the surfaces of BMS-Br and BMS-g-HPBBEA are 68% and 76%, respectively[69], the 7-day HUVECs proliferation data cultured with calcium silicate extracts ranged between 30% and 70%[70]. ...
1
1990
... NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P). ...
1
2000
... NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P). ...
1
1994
... NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P). ...
1
1999
... NO release from both EA. hy926 and VSMC measured via direct contact with various MgZnYNd substrates is recorded in Fig. 11. For EA. hy926, NO release in the MgZnYNd-B-A-P group were about 6.05 µm L-1, which was significantly higher than MgZnYNd-A-P (4.78 µm L-1), MgZnYNd-P (1.82 µm L-1), and bare MgZnYNd (2.82 µm L-1), even a little higher compared with the 4.40 µm L-1 of the blank control. While for VSMC, similar trend could be detected with a little higher NO release value in each testing group, i.e., the MgZnYNd-B-A-P group also showed the highest value among all the testing groups, which was consistent with our previous data when studying the effects of MgF2/polydopamine coating on MgZnYNd alloy[50]. Preceding reports demonstrated NO capacity in obstructing proliferation of VSMC cell[71], [72] and [73] and promoting endothelial cells proliferation[74] except its role in regulating vascular elasticity as a self-produced vaso-active substance. Therefore, it could be concluded that MgZnYNd-B-A-P had a better ability in retarding VSMC growth and promoting endothelial cells proliferation with the higher NO release level, thus minimizing the risk of endotheliosis-induced restenosis and accelerating recovery of vascular endothelium, both of which is beneficial in coronary stent application. Therefore, combined this NO release data with the cell cytotoxicity value showed in Fig. 10, the MgZnYNd-B-A-P offered not only much better cell proliferation but also biological functionality such as NO release compared to the group without silane modification (MgZnYNd-P) and the group with single silane layer (MgZnYNd-A-P). ...
1
2014
... Coupled the data of cell adhesion, cell cytotoxicity and NO release showed above, it is obvious that the bilayer BTSE-APTES pre-treatment benefits a lot when coating MgZnYNd with PLGA, which is reflected in the ability of MgZnYNd-B-A-P to significantly enhance ECV304 and VSMC viability over MgZnYNd-A-P and MgZnYNd-P. Further, the difference in the chemical composition plus the further physical barrier functionality of cross-linked silane coating between MgZnYNd-B-A-P and MgZnYNd-A-P should result to the total different responses for both types of cells. It was reported by Zheng et al. that MC3T3-E1 attachment, stretching, and proliferation were effectively enhanced on GRGD-modified PEEK surface by tailored silanization technique[75]. Besides, silane-coupling agent aminopropyl triethoxysilane was used as a very effective system for collagen immobilizing onto stainless steel to improve human mesenchymal stem cells adhesion and proliferation[27]. ...
