Article pubs.acs.org/journal/abseba
Comparison Study on Four Biodegradable Polymer Coatings for Controlling Magnesium Degradation and Human Endothelial Cell Adhesion and Spreading Wensen Jiang,† Qiaomu Tian,‡ Tiffany Vuong,‡ Matthew Shashaty,‡,§ Chris Gopez,†,∥ Tian Sanders,‡ and Huinan Liu*,†,‡ †
Materials Science and Engineering, ‡Department of Bioengineering, and §College of Natural and Agricultural Sciences, University of California at Riverside, 900 University Avenue, Riverside, California 92521, United States ∥ Narco College, 2001 Third Street, Norco, California 92860, United States ABSTRACT: Magnesium (Mg)-based bioresorbable cardiovascular scaffold (BCS) is a promising alternative to conventional permanent cardiovascular stents, but it faces the challenges of rapid degradation and poor endothelium recovery after device degradation. To address these challenges, we investigated poly(Llactic acid) (PLLA), poly(lactic-co-glycolic acid) (PLGA) (90:10), PLGA (50:50), and polycaprolactone (PCL) coatings on Mg, respectively, and evaluated their surface and biological properties. Intact polymer coatings with complete coverage on Mg substrate were achieved. The biological performance of the materials was evaluated by culturing with human umbilical vein endothelial cells (HUVECs) in vitro using the direct culture method. The pH of the culture media and Mg2+ and Ca2+ ion concentrations in the media were measured after culture to characterize the degradation rate of the materials in vitro. The results showed that the PLGA (50:50) coating improved the adhesion and spreading of HUVECs the most among the four polymer coatings. Moreover, we found three possible factors that promoted HUVECs directly attached on the surface of PLGA (50:50)coated Mg: (1) the higher concentration of Mg2+ ions released into culture media with a concentration range of 9−15 mM; (2) the lower Ca2+ ion concentration in culture media at 1.3−1.6 mM; and (3) the favorable surface conditions of PLGA (50:50), when compared with the other sample groups. This in vitro study provided the first evidence that the PLGA (50:50) is a promising coating material for Mg-based biodegradable metals toward potential cardiovascular or neurovascular applications. KEYWORDS: bioresorbable magnesium implants, polymer coatings, bioresorbable cardiovascular scaffold, human umbilical vein endothelial cells, in vitro direct culture method
1. INTRODUCTION Cardiovascular stents made of nondegradable biomaterials are implanted in the blood vessels permanently. Their long-term presence in the vessels often leads to complications, such as thrombosis, neointimal hyperplasia, and chronic inflammations.1−6 Drug-eluting stents (DES) have been developed to overcome the acute complications;1,7 however, their long-term risks are even greater than that of bare metal stents (BMS) due to the delay of endothelialization caused by the drugs.1,3,6,8 If fractured or dislodged, both the DES and BMS must be retrieved from the body and the removal procedure is highly invasive. A bioresorbable cardiovascular scaffold (BCS) is a promising alternative to current permanent stents. Ideally, biomaterials for BCS should meet the following requirements. First, their degradation products (intermediate and final products) should have minimal toxicity to the human body. Second, the degradation rate of biomaterials should match the recovery rate of vascular tissue. Rapid degradation could lead to the loss © XXXX American Chemical Society
of radial support from the stent too early, while a slow degradation could increase the risks of late complications induced by the materials. Third, biomaterials should induce rapid endothelialization to restore the functions of vascular tissue. Full degradation of the BCS and complete recovery of the endothelium would minimize the chronic complications and liberate the repaired vessels from the restriction caused by a permanent stent. Aliphatic polymers and copolymers have attractive properties to serve as biomaterials for BCSs, and they were initially investigated as the biodegradable materials for coronary stents in humans.9 Studies indicated that aliphatic polyesters are potentially beneficial for vascular tissue recovery.10−15 The final degradation products of aliphatic polyesters are CO2 and H2O, which can be safely eliminated from the body through natural Received: April 6, 2017 Accepted: May 5, 2017
A
DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
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ACS Biomaterials Science & Engineering Table 1. Degradation Property of Representative Biodegradable Polymers, Mg, and Mg Alloya materials
degradation
PLLA PLLA PLLA PLGA (90:10) PLGA (DL-85:15) PLGA (75:25) PLGA (50:50) PLGA (50:50) PLGA (DL-50:50) PCL Mg Alloy Pure Mg
Complete degradation at 36 months Near complete degradation at 24 months Complete degradation >36 months Near complete degradation at 20 weeks >80% mass loss at 23−30 weeks >50% mass loss at 2 weeks >50% mass loss at 1 week Near complete degradation at 4 weeks >40% mass loss at 5.5−7.5 weeks Remain intact at 24 months Complete degradation at 4 months 39% mass loss at 6 weeks
condition In In In In In In In In In In In In
vivo (human) vivo (human) vivo (ewes) vivo (rats) vitro vivo (rats) vivo (rats) vitro vitro vivo (rats) vivo (human) vivo (rats)
mol wt or composition
ref
183 kDa N/A 360 kDa N/A 73−91 KDa 50 kDa 46 kDa 200 kDa 81−87 KDa 66 kDa AE21 (2% Al, 1% Rare Earths) Purity >99.9%
31 32 33 34 35 36 36 41 35 37 38, 39 40
a
Note: Not only molecular weight or composition could influence the degradation rates, but also other factors played a role in the degradation results, such as the dimension and geometry of the samples, processing conditions of the samples, the media and culture conditions used for in vitro testing, and the implanted location for in vivo testing.
polycaprolactone (PCL) were selected because they represent a wide range of degradation rates as summarized in Table 1.31−41 The degradation rate of Mg is also included in Table 1 for comparison. The degradation and biological properties of polymer-coated Mg were investigated in vitro by culturing with human umbilical vein endothelial cells (HUVECs) using the previously established direct culture method.42,43 It is expected that the polymer coatings will reduce the degradation rate of Mg differently due to their different degradation properties as summarized in Table 1. It is hypothesized that the polymer coatings could also promote the adhesion and spreading of HUVECs. Moreover, nondegradable titanium alloy substrates were included as controls in the experimental design of HUVEC culture, which enabled us to isolate the effects of polymer coatings versus substrates. The methods established in this study could serve as a standard for future research on surface treatment of biodegradable metals for vascular applications.
metabolism. However, their intermediate degradation products, such as low-molecular-weight alcohols and acid, may reduce local pH and cause severe inflammation.16 In addition, their expansion in blood is slow and their implantation procedure may require heating of the balloon, which could cause injuries to the blood vessel.9 Magnesium (Mg) and its alloys are a category of biodegradable metals, which have attracted great interest in cardiovascular applications. Mg reacts with water, releasing Mg2+, OH−, and H2, as shown in Reaction 1.17 Mg (s) + 2H2O(l) → Mg 2 + + 2OH − + H2(g )
(1)
Mg is an essential element and the fourth most abundant cation in the human body,18,19 and the concentration of Mg2+ ions in serum is about 0.9 mM.20 Mg2+ ion also participates in over 300 enzymatic metabolisms.21 Moreover, the deliverability of a Mg-based stent is much closer to that of the widely used stainless steel stent, in contrast to the polymeric stent.22 Unfortunately, pure Mg and current Mg alloys degrade too rapidly to meet the clinical requirements for cardiovascular applications, and the Mg-based BCS has not yet achieved sufficient endothelium restoration in clinical trials.23,24 Applying polymer coatings on Mg could potentially address the aforementioned problems. The effectiveness of polymer coatings in reducing the degradation rate of Mg has been widely reported.15,23−29 In addition, polymer coatings could potentially improve endothelium recovery.10−15 Due to the merits of polymer coating, substantial progress13,15,25−28,30 has been reported for polymer-coated Mg-based materials, but most studies only involved one or two types of polymers, and few pieces of literature reported their biological performances with endothelial cells. The comparison between different polymers reported in different articles could not converge into a convincing conclusion because the conditions set for those experiments were significantly different. To allow direct comparison, therefore, there is a need to systemically investigate various polymer coatings on Mg, and their biological performances with endothelial cells for cardiovascular or neurovascular applications. The objectives of this study are to investigate four major biodegradable aliphatic polyesters and directly compare them in vitro as representative coating materials on Mg for potential vascular device applications. Poly(L-lactic acid) (PLLA), poly(lactic-co-glycolic acid) (PLGA 90:10), PLGA 50:50, and
2. MATERIALS AND METHODS 2.1. Preparation of Mg and Nitinol (NiTi) Substrates. Mg and Nitinol (NiTi alloy) substrates were prepared following the same procedure. A Mg sheet (as rolled, 99.9% purity, Goodfellow Cambridge Limited, 250 μm thick) was cut into 5 mm × 5 mm squares and used as the biodegradable model substrate. A NiTi sheet (flat annealed, pickled surface, consisting of 55.75% Ni and 44.25% Ti by weight, Alfa Aesar, 250 μm thick) was cut into 5 mm × 5 mm squares and used as the nondegradable model substrate. All the Mg and NiTi substrates were cut using a notcher (Whitney Metal Tool Co.). The Mg or NiTi squares were degreased in acetone for 15 min using an ultrasonic cleaner (VWR symphony Ultrasonic Cleaners), followed by ultrasonic cleaning in 100% ethanol for 15 min. The samples were then dried in air, and mounted in epoxy resin (lowviscosity castable resin, Pace Technologies Inc.). Specifically, the lowviscosity castable hardener was added to the epoxy resin at a weight ratio of 1:10, and the mixture was manually stirred and degassed in a vacuum oven for 15 min. After degassing, the mixture was precured at 70 °C for 15 min to increase its viscosity, and then the Mg squares were placed on top of the precured epoxy resin mixture. For NiTi squares, the precuring time was 20 min. After precuring, the epoxy resin mixture was then fully cured at 70 °C for another 1 h. The optical image of the fully cured sample before coating is shown in Figure 1a1. Figure 1a2 confirmed the boundary between the metal substrate and epoxy resin was smooth without visible defects. The mounted Mg and NiTi substrates were polished using SiC abrasive paper and diamond paste. Specifically, Mg or NiTi substrates were polished with SiC abrasive cloth of 600 grit, 800 grit, and 1200 B
DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
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a strong organic solvent, and thus chemical-resistant Teflon tips were used for pipetting the polymer solutions. The spin coating was carried out at a spin speed of 7000 rpm, an acceleration rate of 1000 rpm/s, and a spinning duration of 60 s. The coated substrates were dried in air for 24 h and then dried in vacuum for 48 h to allow complete evaporation of chloroform. After solvent removal, the PCL coatings were heated at 80 °C for 5 min, which is 20 °C higher than the melting point of PCL (Tm, 60 °C49). The image of a coated sample is shown in Figure 1b1. The polymer coatings are transparent, and thus a green color is used to illustrate the coated region on the metal substrate in Figure 1b1. Figure 1b2 illustrates the cross section of the coated sample. 2.3. Characterization of Surface Microstructure and Coating Thickness. Surface microstructures of the polymer-coated and noncoated Mg samples were characterized using a scanning electron microscope (SEM; Nova Nano SEM 450, FEI). The surface microstructures of the polymer-coated NiTi controls were also characterized and confirmed to be intact using SEM, but the images are not included due to the similarity of the polymer coatings on NiTi to their counterparts on Mg samples. Surface morphologies of Mg after different polishing steps were imaged using a 3D laser scanning microscope (VK-X150, Keyence), and the image area was 140 μm × 105 μm in size for each image. The surface roughness (Sq) of the entire image area, i.e., root-mean-squared height, was calculated using MultiFileAnalyzer (VK-H1XME) by Keyence. The horizontal midplane between the highest point and the lowest point in the image was set as reference plane for calculating Sq. The Sq was calculated according to eq 2, where Z is the deviation in height from the reference plane, and A is the image area projected to the horizontal plane:
Figure 1. Sample configuration and the design of the HUVEC culture. (a1) Optical image of a metallic substrate mounted in the epoxy resin before coating. (a2) Optical image of the magnified interface between the metallic substrate and the epoxy resin. (b1) Top view of a polymer-coated metallic substrate mounted in the epoxy resin, where the green color illustrates the transparent polymer coating. (b2) Illustration of the cross section of a polymer-coated metallic substrate mounted in the epoxy resin, where the green color represents the polymer coating, the gray color represents the metallic substrate, and the blue color represents the epoxy resin. (c) Illustration of HUVECs in direct and indirect contact with the sample in direct culture in vitro. Scale bar = 5 mm for (a1) and (b1). Scale bar = 0.5 mm for (a2).
Sq =
grit (Ted Pella Inc.) in 100% ethanol, and then with diamond paste of 6 μm, 3 μm, 1 μm, and 0.25 μm grade (Physical Test Solutions). The Mg or NiTi substrates were sonicated in 100% ethanol for 5 min. The samples were sonicated in between each polishing step with the respective grades of diamond paste and after the final polishing with 0.25 μm grade diamond paste. The as-polished Mg or NiTi substrates with a mirror-like reflective surface were degreased by gently rinsing their surface with acetone. The substrates were then cleaned again by sonicating in 100% ethanol for 15 min in the ultrasonic cleaner. The polished Mg or NiTi substrates were used as noncoated controls, and as the substrates for polymer coatings. 2.2. Preparation of Polymer Coatings by Spin Coating. The polymers selected for this study were PLLA (No. AP007, Mn = 125,000−150 000 Da, acid end-cap, Polyscitech), PLGA (90:10) (No. AP049, Mn = 125,000−150 000 Da, acid end-cap, Polyscitech), PLGA (50:50) (No. AP089, Mn = 75,000−85 000 Da, acid end-cap, Polyscitech), and PCL (No. AP009, Mn = 150,000−200 000 Da, acid end-cap, Polyscitech). The polymers were spin-coated onto Mg or NiTi substrates. We chose spin coating as the model coating method to compare different polymers as the coating materials, because spin coating is versatile for nonconductive polymers, and highly repeatable and tunable for producing consistent surfaces with complete coating coverage for direct comparison,27,44,45 which is difficult to achieve with other major coating technologies, such as electrospinning46 and electrodeposition.47,48 To prepare the polymer solutions for spin coating, polymers with a mass of 0.1 g were first dissolved in 2 mL of chloroform to achieve a concentration of 0.05 g/mL (i.e., wt/vol % = 5%). The process of dissolving PLGA (50:50) and PCL was accelerated using a speed mixer (FlackTek Inc.) at 2500 rpm for 5 min and an ultrasonic bath (VWR symphony Ultrasonic Cleaners). The polymer solutions were sonicated at a low power in the ultrasonic bath for 1.5 h for PLLA, PLGA (90:10), and PLGA (50:50) solutions and for 6 h for the PCL solution. The water bath in the ultrasonic cleaner was maintained at 40 °C during the entire sonication process. A 25 μL polymer solution was pipetted onto the Mg or NiTi substrate mounted in epoxy resin that was secured onto the sample stage of a spin coater via vacuum (PWM32, Headway Research). Chloroform is
1 A
∬A Z2(x , y)dxdy
(2)
The linear profiles on the surface of Mg with different polishing conditions were also obtained using the 3D laser scanning microscope (VK-X150, Keyence). The linear profile was shown as relative height versus scanning distance, where the lowest point in the vertical direction was set as the reference point (relative height = 0 μm). The linear roughness Rq over the profiled line, i.e., root-mean-square roughness, was calculated using MultiFileAnalyzer (VK-H1XME) by Keyence. The horizontal midplane between the highest point and the lowest point in the linear profile was set as reference plane for calculating Rq. The Rq was calculated following eq 3, where Z is the deviation in height from the reference plane, and L is the total scanning distance in horizontal direction: Rq =
1 L
∫0
L
Z2(x)dx
(3)
The effects of polishing on the surface morphologies of the polymer coatings, as well as the effects of heat treatment on PCL coatings, were characterized using SEM. The SEM images were taken with low vacuum mode at an acceleration voltage of 2 kV to 3 kV. The thickness of the polymer coatings on Mg substrates was measured using a stylus surface profilometer (Vecco system Dektak 8). To measure the thickness, half of the polymer coating was removed using a razor blade, and the thickness of the coating was determined by the linear profiling from the surface profilometer. The average thickness was obtained by measuring three different locations on each sample. 2.4. Characterization of Surface Wettability by Water Contact Angle. Surface wettability was characterized by measuring water contact angle of the samples using a drop shape analyzer (EasyDrop, Kruss). Specifically, optical images of the droplets of 0.5 μL deionized (DI) water were taken at 10 s after being dropped on the samples, and were used to measure the water contact angle in the drop shape analysis software (Kruss). 2.5. Characterization of Corrosion Resistance Using the Tafel Method. The corrosion resistance of the polymer-coated and noncoated Mg samples in revised simulated body fluids (rSBF)50 was determined using the Tafel method. Potentiodynamic polarization C
DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
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ACS Biomaterials Science & Engineering (PDP) curves of the polymer-coated and noncoated Mg were obtained using a potentiostat/galvanostat (Model 273A, Princeton Applied Research). The 5 mm × 5 mm × 1 mm samples were mounted in epoxy resin, and the mounted sample had a thickness of 3 mm. One surface of the sample was connected with a bundle of copper wires and sealed with epoxy resin, and the other surface was exposed in rSBF. The exposed area (A) of each sample was measured based on their optical images using the analysis tools in ImageJ. The bundle of copper wires from the sample was connected to a working electrode in a three-electrode configuration, where platinum was used as a counter electrode, and Ag/AgCl electrode was used as a reference electrode. The copper wire bundle was completely isolated from rSBF and had no contact with the liquid. PDP curves were obtained by scanning from 0 V to −3 V at a scan rate of 20 mV/s. Corrosion potential (Ecorr) and current (Icorr) were determined by extrapolating the tangent line of the anodic half and cathodic half, following the ASTM G102-89 standard. The corrosion current density (Jcorr) was calculated as Jcorr = Icorr/A, where A is the exposed surface area of the sample in rSBF. 2.6. HUVEC Culture with the Polymer-Coated and Noncoated Mg and Controls. 2.6.1. Direct Culture of HUVECs with the Samples. HUVECs were purchased from Lonza (No. C2519A, pooled donor, Lonza Walkersville Inc.). HUVECs at passage 2 were thawed and cultured in a T-75 flask with 25 mL of media (EGM-2, endothelial growth media, prepared from an EGM-2 Bullet kit, No. cc-3162, Lonza) under standard cell culture conditions (5% CO2, 95% air, 37 °C, sterile, humidified). The commercial EGM-2 media contained vascular endothelial growth factor for rapid growth and proliferation of HUVECs. HUVECs with 80% to 90% confluency were collected at passage 3 and seeded onto the samples. The samples mounted in epoxy resin were attached onto a glass slide using medical adhesives (NO. TM7810A, medical grade double-coated film tape, Mactac) and used for HUVEC culture to prevent the samples from floating in the culture media. The polymer-coated and noncoated Mg samples and NiTi controls were disinfected in 100% ethanol for 1 h, and dried in air at room temperature to remove the residual ethanol. We did not use the standard 70% ethanol for disinfection because water reacts with Mg upon contact. Polymer-coated Nitinol (NiTi) samples were included as NiTi control groups. Noncoated Mg, noncoated NiTi, epoxy resin, and glass samples were included as reference groups. The glass reference was purchased from Fisher Scientific (nonculture treated glass slides, 1 mm thick, No. 12-544-1) and cut into 10 mm × 10 mm squares, and the epoxy resin reference was cured following the same procedure as previously described for sample mounting, and had a dimension of ø10 mm × 1 mm. We also included cell-only and media-only groups as reference groups. HUVECs collected from T-75 flasks were seeded onto the samples at a seeding density of 6000 cells/ cm2, which is recommended by the vendor (Lonza), and incubated for 24 h under standard cell culture conditions. 2.6.2. Characterization of HUVEC Adhesion and Morphology. After 24 h of direct culture, the culture media were collected for analyses and the cells were fixed in 4% methanol-free paraformaldehyde for 20 min. HUVECs were stained with Alexa Fluor 488 Phalloidin (A12379, Life technologies) for F-actin, and 4′6-diamidino2-phenylindole dilactate (DAPI, Invitrogen) for nuclei. HUVECs directly attached on the samples (direct contact) and attached on the culture plate around the samples (indirect contact) were imaged using a fluorescence microscope (Eclipse Ti, Nikon). HUVECs in direct contact and indirect contact with the samples were defined previously for assessing the in vitro cytocompatibility of Mg alloys.42 Figure 1c illustrates the cells in direct contact and indirect contact conditions. For each sample, five areas under direct contact condition and nine areas under indirect contact condition were imaged to quantify the number of adhered cells and calculate the cell adhesion density. Three representative images were used to calculate the spreading area of HUVECs. The HUVEC adhesion density and spreading area were calculated based on the fluorescence images using the analysis tools in ImageJ. Specifically, the number of HUVECs on each imaged area was quantified by manually counting DAPI-stained nuclei, and the adhesion density was calculated as the number of HUVECs per unit area. The spreading area was outlined based on the F-actin stain using
the automatic tools in ImageJ. The spreading area per HUVEC was calculated as the total area stained by Alexa Fluor 488 divided by the number of adherent HUVECs for each image. 2.6.3. Postculture Media Analysis. The pH value of the postculture media was measured immediately after collection using a precalibrated pH meter (Symphony, Model SB70P, VWR). Mg2+ and Ca2+ ion concentrations in the collected media were measured using an inductively coupled plasma optical emission spectrometer (ICP-OES, Optima 8000, PerkinElmer). The collected media was diluted to 1:100 solutions in deionized (DI) water in order to minimize the matrix effects. ICP-OES was calibrated using the standard solutions of Mg2+ and Ca2+ that were both diluted to the ranges of 0.5−5.0 mg/L. 2.6.4. Postculture Analysis of Ca Deposition onto the Sample Surface. The mass of total Ca deposited onto the sample surface was quantified using ICP-OES. Specifically, the samples after HUVEC culture were carefully moved into new 12-well plates and completely dissolved in a 2% HNO3 aqueous solution. The dissolving process using 2% HNO3 solutions was repeated three times to ensure that all Ca was collected into the solution. Briefly, 3 mL of a 2% HNO3 solution was added into each well containing the postculture samples or controls. The first round of dissolving lasted 90 min, and the solution was collected into a 15 mL tube. After the first 90 min of dissolving, all samples with Mg as substrate showed no visible remains except for the epoxy resin in the sample mount. For the second round of dissolving, 3 mL of 2% HNO3 solution was added into the well for 15 min, and the solution was collected into the tube. For the third round of dissolving, 4 mL of 2% HNO3 solution was added into the well for 15 min, and the solution was subsequently collected. The total volume of the collected solution after three rounds of dissolving in 2% HNO3 solution was recorded. A similar process was performed for the cell-only and media-only groups for comparison. The Ca2+ ion concentration, designated as Conc(Ca2+), in the collected solution was measured using the ICP-OES (Optima 8000, PerkinElmer). ICP-OES was calibrated using the standard solutions of Ca2+ that were diluted to the range 0.5−5.0 mg/L. The mass of total Ca deposited on the sample surface was calculated as Mass(Ca) = Conc(Ca2+) × V. 2.7. Statistical Analysis. The numerical data were examined using one-way analysis of variance (ANOVA) followed by posthoc test. The statistical analysis was performed using GraphPad Prism 7 software. Statistical significance was considered at p < 0.05. Each sample group and control group were run in triplicate in the experiments for wettability, corrosion resistance, and HUVEC culture. The measurements for Ca deposition were repeated twice.
3. RESULTS 3.1. Effects of Polishing on the Surface Morphologies of Mg Substrates, Roughness of Mg Substrates, and Morphologies of Polymer Coatings. Figure 2 shows the surface morphologies and roughness of the Mg substrates at different polishing steps and the polymer coatings on the polished substrates. Specifically, the surface morphology and surface roughness (Sq) of Mg after polishing with a 600 grit cloth are shown in Figure 2a1; the surface morphology and surface roughness of Mg after sequential polishing with 600, 800, and 1200 grit cloths are shown in Figure 2b1; the surface morphology and surface roughness of Mg after sequential polishing with 600, 800, and 1200 grit cloths followed by diamond pastes of 6, 3, 1, and 0.25 μm grade are shown in Figure 2c1. The fine polishing process with diamond pastes significantly improved the smoothness of the substrates and the consistency and uniformity of the polymer coatings on these substrates. Specifically, the fine-polished Mg substrate in Figure 2c1 was much smoother than the 600 grit-polished Mg substrate (Figure 2a1) and the 1200 grit-polished Mg substrate (Figure 2b1); the surface roughness of fine-polished Mg substrate (Sq = 0.08 μm) in Figure 2c1 was much smaller than those of 600 grit-polished Mg substrate (Sq = 0.70 μm, Figure D
DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
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remove pores in the PCL coating after spin coating. The SEM image of the spin-coated PCL showed many pores before the heat treatment, as highlighted by the red dashed circles in Figure 3a1 and a2, which possibly formed during the process of
Figure 2. Effects of polishing on the surface morphologies of Mg substrates and polymer coatings. (a1, b1, c1) Surface morphologies of noncoated Mg after (a1) polishing with 600 grit SiC paper, (b1) sequentially polishing with 600, 800, and 1200 grit SiC paper, (c1) sequential polishing with 600, 800, and 1200 grit SiC paper, and then with diamond pastes of 6, 3, 1, and 0.25 μm grade. The Sq values shown in (a1), (b1), and (c1) are the root-mean-squared heights of (a1), (b1), and (c1), respectively. (a2, b2, c2) Linear profiles of height along the respective lines drawn in the (a1, b1, c1) laser scanning optical images. All linear profile graphs in (a2, b2, c2) have the same range of Y-axis from 0 to 3 μm. The minimum height in each linear profile (a2, b2, c2) was set as 0 μm; and the maximum height in each linear profile is indicated by the red dashed lines. The linear rootmean-square roughness (Rq) values are shown in (a2), (b2), and (c2). (a3) SEM image of PLGA (50:50) coating on (a1) 600 grit-polished Mg. (c3) SEM image of PLGA (50:50) coating on (c1) fine-polished Mg. Scale bar = 20 μm for (a1), (b1), and (c1). Scale bar = 50 μm for (a3) and (c3). Original magnification = 480× for (a1), (b1), and (c1). Original magnification = 1000× for (a3) and (c3).
Figure 3. SEM images of PCL-coated Mg before and after heat treatment. (a1, a2) SEM images of PCL coating before the heat treatment at (a1) low magnification (1000×) and (a2) high magnification (5000×). The red dashed circles in (a1) and (a2) highlight the pores on the PCL coating. (b1, b2) SEM images of PCL coating after the heat treatment at (b1) low magnification (1000×) and (b2) high magnification (5000×). Scale bar = 50 μm for (a1) and (b1). Scale bar = 10 μm for (a2) and (b2).
crystallization. In contrast, the SEM image of the spin-coated PCL after the heat treatment showed no pores (Figure 3b1 and b2). The removal of the pores on the PCL coating limited the exposure of the Mg substrates to water. 3.3. Characterization of the Polymer Coatings on Mg Substrates. The PLLA, PLGA (90:10), PLGA (50:50), and PCL coatings completely covered the Mg substrates, and all the coatings showed a uniform morphology, as shown in the SEM images in Figure 4a−d. The PLLA, PLGA (90:10), and PLGA (50:50) coatings exhibited an amorphous morphology (Figure 4a−c), and the PCL coating exhibited a crystalline morphology (Figure 4d). The surface morphology of the noncoated Mg appeared uniform and consistent, as shown in Figure 4e, and the polishing traces were less than 1 μm in width after finepolishing with diamond pastes of 0.25 μm. All polymer coatings have a thickness between 0.5−2 μm, as shown in Figure 4f. The coatings of different polymers had different thicknesses when the parameters of spin coating were fixed; that is, the spin speed was set at 7000 rpm; the acceleration was set as 1000 rpm/s; the spin duration was 60 s; and the polymer/solvent ratio was 5%. Specifically, PLLA coating showed a statistically significant greater thickness than all the other polymer coatings, and PLGA (90:10) coating showed a statistically significant greater thickness than PLGA (50:50) coating. PLGA (50:50) coating showed the lowest thickness on average among all the polymer coatings. Figure 4g illustrates the molecular structures of the biodegradable polymers used in this study. The monomer of PLLA is lactide. PLGA is a copolymer of lactide and glycolide. The x and y represent the molar percentage of the lactide and glycolide monomer, respectively. For example, PLGA (90:10)
2a1) and the 1200 grit-polished Mg substrate (Sq = 0.15 μm, Figure 2b1). The respective linear profile of relative height in Figure 2(a2, b2, and c2) confirmed that the fine-polished Mg substrate had a much smaller variation in relative height ( 10 Years) Clinical Outcomes of First-in-Human Biodegradable Poly-lLactic Acid Coronary StentsClinical Perspective. Circulation 2012, 125 (19), 2343−2353. (32) Serruys, P. W.; Ormiston, J. A.; Onuma, Y.; Regar, E.; Gonzalo, N.; Garcia-Garcia, H. M.; Nieman, K.; Bruining, N.; Dorange, C.; Miquel-Hébert, K. A bioabsorbable everolimus-eluting coronary stent system (ABSORB): 2-year outcomes and results from multiple imaging methods. Lancet 2009, 373 (9667), 897−910. (33) Walton, M.; Cotton, N. J. Long-term in vivo degradation of poly-L-lactide (PLLA) in bone. J. Biomater. Appl. 2007, 21 (4), 395− 411. (34) Luis, A. L.; Rodrigues, J. M.; Amado, S.; Veloso, A. P.; ArmadaDa-silva, P. A.; Raimondo, S.; Fregnan, F.; Ferreira, A. J.; Lopes, M. A.; Santos, J. D. PLGA 90/10 and caprolactone biodegradable nerve guides for the reconstruction of the rat sciatic nerve. Microsurgery 2007, 27 (2), 125−137. (35) Ramchandani, M.; Pankaskie, M.; Robinson, D. The influence of manufacturing procedure on the degradation of poly (lactide-coglycolide) 85:15 and 50:50 implants. J. Controlled Release 1997, 43 (2), 161−173. (36) Miller, R. A.; Brady, J. M.; Cutright, D. E. Degradation rates of oral resorbable implants (polylactates and polyglycolates): rate modification with changes in PLA/PGA copolymer ratios. J. Biomed. Mater. Res. 1977, 11 (5), 711−719. (37) Sun, H.; Mei, L.; Song, C.; Cui, X.; Wang, P. The in vivo degradation, absorption and excretion of PCL-based implant. Biomaterials 2006, 27 (9), 1735−1740. (38) Heublein, B.; Rohde, R.; Kaese, V.; Niemeyer, M.; Hartung, W.; Haverich, A. Biocorrosion of magnesium alloys: a new principle in cardiovascular implant technology? Br. Heart J. 2003, 89 (6), 651− 656. (39) Erbel, R.; Di Mario, C.; Bartunek, J.; Bonnier, J.; de Bruyne, B.; Eberli, F. R.; Erne, P.; Haude, M.; Heublein, B.; Horrigan, M. Temporary scaffolding of coronary arteries with bioabsorbable magnesium stents: a prospective, non-randomised multicentre trial. Lancet 2007, 369 (9576), 1869−1875. (40) Yang, W.; Zhang, Y.; Yang, J.; Tan, L.; Yang, K. Potential antiosteoporosis effect of biodegradable magnesium implanted in STZinduced diabetic rats. J. Biomed. Mater. Res., Part A 2011, 99 (3), 386− 394. (41) Gu, X.; Mao, Z.; Ye, S.-H.; Koo, Y.; Yun, Y.; Tiasha, T. R.; Shanov, V.; Wagner, W. R. Biodegradable, elastomeric coatings with controlled anti-proliferative agent release for magnesium-based cardiovascular stents. Colloids Surf., B 2016, 144, 170−179.
(42) Cipriano, A. F.; Sallee, A.; Guan, R.-G.; Zhao, Z.-Y.; Tayoba, M.; Sanchez, J.; Liu, H. Investigation of magnesium−zinc−calcium alloys and bone marrow derived mesenchymal stem cell response in direct culture. Acta Biomater. 2015, 12, 298−321. (43) Cipriano, A. F.; Sallee, A.; Tayoba, M.; Alcaraz, M. C. C.; Lin, A.; Guan, R.-G.; Zhao, Z.-Y.; Liu, H. Cytocompatibility and early inflammatory response of human endothelial cells in direct culture with Mg-Zn-Sr alloys. Acta Biomater. 2017, 48, 499−520. (44) Wang, L.; Li, H.; Chen, S.; Nie, C.; Cheng, C.; Zhao, C. Interfacial Self-Assembly of Heparin-Mimetic Multilayer on Membrane Substrate as Effective Antithrombotic, Endothelialization, and Antibacterial Coating. ACS Biomater. Sci. Eng. 2015, 1 (11), 1183− 1193. (45) Wang, R.; Chua, K. L.; Neoh, K. G. Bifunctional Coating with Sustained Release of 4-Amide-piperidine-C12 for Long-Term Prevention of Bacterial Colonization on Silicone. ACS Biomater. Sci. Eng. 2015, 1 (6), 405−415. (46) Low, K.; Horner, C. B.; Li, C.; Ico, G.; Bosze, W.; Myung, N. V.; Nam, J. Composition-dependent sensing mechanism of electrospun conductive polymer composite nanofibers. Sens. Actuators, B 2015, 207, 235−242. (47) Tian, Q.; Liu, H. Electrophoretic deposition and characterization of nanocomposites and nanoparticles on magnesium substrates. Nanotechnology 2015, 26 (17), 175102. (48) Yan, Y.; Zhang, M.; Moon, C. H.; Su, H.-C.; Myung, N. V.; Haberer, E. D. Viral-templated gold/polypyrrole nanopeapods for an ammonia gas sensor. Nanotechnology 2016, 27 (32), 325502. (49) Elzubair, A.; Elias, C. N.; Suarez, J. C. M.; Lopes, H. P.; Vieira, M. V. B. The physical characterization of a thermoplastic polymer for endodontic obturation. J. Dent. 2006, 34 (10), 784−789. (50) Oyane, A.; Kim, H. M.; Furuya, T.; Kokubo, T.; Miyazaki, T.; Nakamura, T. Preparation and assessment of revised simulated body fluids. J. Biomed. Mater. Res. 2003, 65 (2), 188−195. (51) Cipriano, A. F.; De Howitt, N.; Gott, S. C.; Miller, C.; Rao, M. P.; Liu, H. Bone marrow stromal cell adhesion and morphology on micro-and sub-micropatterned titanium. J. Biomed. Nanotechnol. 2014, 10 (4), 660−668. (52) Jamshidi, K.; Hyon, S.-H.; Ikada, Y. Thermal characterization of polylactides. Polymer 1988, 29 (12), 2229−2234. (53) Lan, P.; Zhang, Y.; Gao, Q.; Shao, H.; Hu, X. Studies on the synthesis and thermal properties of copoly (L-lactic acid/glycolic acid) by direct melt polycondensation. J. Appl. Polym. Sci. 2004, 92 (4), 2163−2168. (54) Mainardes, R. M.; Gremião, M. P. D.; Evangelista, R. C. Thermoanalytical study of praziquantel-loaded PLGA nanoparticles. Rev. Bras. Cienc. Farm. 2006, 42 (4), 523−530. (55) Witecka, A.; Yamamoto, A.; Idaszek, J.; Chlanda, A.; ́ Swięszkowski, W. Influence of biodegradable polymer coatings on corrosion, cytocompatibility and cell functionality of Mg-2.0 Zn-0.98 Mn magnesium alloy. Colloids Surf., B 2016, 144, 284−292. (56) Kirkland, N.; Birbilis, N.; Staiger, M. Assessing the corrosion of biodegradable magnesium implants: a critical review of current methodologies and their limitations. Acta Biomater. 2012, 8 (3), 925−936. (57) Li, Z.; Gu, X.; Lou, S.; Zheng, Y. The development of binary Mg−Ca alloys for use as biodegradable materials within bone. Biomaterials 2008, 29 (10), 1329−1344. (58) Kokubo, T. Formation of biologically active bone-like apatite on metals and polymers by a biomimetic process. Thermochim. Acta 1996, 280, 479−490. (59) Kokubo, T. Apatite formation on surfaces of ceramics, metals and polymers in body environment. Acta Mater. 1998, 46 (7), 2519− 2527. (60) Banai, S.; Haggroth, L.; Epstein, S. E.; Casscells, W. Influence of extracellular magnesium on capillary endothelial cell proliferation and migration. Circ. Res. 1990, 67 (3), 645−650. (61) Zhang, K.; Chen, J. The regulation of integrin function by divalent cations. Cell adhesion & migration 2012, 6 (1), 20−29. N
DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
Article
ACS Biomaterials Science & Engineering (62) Leitinger, B.; McDowall, A.; Stanley, P.; Hogg, N. The regulation of integrin function by Ca 2+. Biochim. Biophys. Acta, Mol. Cell Res. 2000, 1498 (2), 91−98. (63) Xiong, J. P.; Stehle, T.; Goodman, S.; Arnaout, M. Integrins, cations and ligands: making the connection. J. Thromb. Haemostasis 2003, 1 (7), 1642−1654. (64) Knedler, A.; Ham, R. G. Optimized medium for clonal growth of human microvascular endothelial cells with minimal serum. In Vitro Cell. Dev. Biol. 1987, 23 (7), 481−491. (65) Moore, L.; Pastan, I. ENERGY-DEPENDENT CALCIUM UPTAKE BY FIBROBLAST MICROSOMES. Ann. N. Y. Acad. Sci. 1978, 307 (1), 177−194. (66) Afroze, T.; Yang, G.; Khoshbin, A.; Tanwir, M.; Tabish, T.; Momen, A.; Husain, M. Calcium efflux activity of plasma membrane Ca2+ ATPase-4 (PMCA4) mediates cell cycle progression in vascular smooth muscle cells. J. Biol. Chem. 2014, 289 (10), 7221−7231. (67) Whitfield, J.; Boynton, A.; MacManus, J.; Sikorska, M.; Tsang, B. The regulation of cell proliferation by calcium and cyclic AMP. Mol. Cell. Biochem. 1979, 27 (3), 155−179. (68) Brini, M.; Calì, T.; Ottolini, D.; Carafoli, E. The plasma membrane calcium pump in health and disease. FEBS J. 2013, 280 (21), 5385−5397. (69) Lipskaia, L.; Lompré, A. M. Alteration in temporal kinetics of Ca2+ signaling and control of growth and proliferation. Biology of the Cell 2004, 96 (1), 55−68. (70) Tada, M.; Yamamoto, T.; Tonomura, Y. Molecular mechanism of active calcium transport by sarcoplasmic reticulum. Physiol. Rev. 1978, 58 (1), 1−79. (71) Tian, Q.; Deo, M.; Rivera-Castaneda, L.; Liu, H. Cytocompatibility of Magnesium Alloys with Human Urothelial Cells: A Comparison of Three Culture Methodologies. ACS Biomater. Sci. Eng. 2016, 2, 1559. (72) Rohman, G.; Pettit, J. J.; Isaure, F.; Cameron, N. R.; Southgate, J. Influence of the physical properties of two-dimensional polyester substrates on the growth of normal human urothelial and urinary smooth muscle cells in vitro. Biomaterials 2007, 28 (14), 2264−2274. (73) Discher, D. E.; Janmey, P.; Wang, Y.-l. Tissue cells feel and respond to the stiffness of their substrate. Science 2005, 310 (5751), 1139−1143. (74) Park, J. J.; Yu, E. J.; Lee, W. K.; Ha, C. S. Mechanical properties and degradation studies of poly (D, L-lactide-co-glycolide) 50:50/ graphene oxide nanocomposite films. Polym. Adv. Technol. 2014, 25 (1), 48−54. (75) Jacobsen, S.; Fritz, H.-G. Plasticizing polylactidethe effect of different plasticizers on the mechanical properties. Polym. Eng. Sci. 1999, 39 (7), 1303−1310. (76) Torres, F.; Nazhat, S.; Fadzullah, S. S. M.; Maquet, V.; Boccaccini, A. Mechanical properties and bioactivity of porous PLGA/ TiO 2 nanoparticle-filled composites for tissue engineering scaffolds. Compos. Sci. Technol. 2007, 67 (6), 1139−1147. (77) Washburn, N. R.; Yamada, K. M.; Simon, C. G.; Kennedy, S. B.; Amis, E. J. High-throughput investigation of osteoblast response to polymer crystallinity: influence of nanometer-scale roughness on proliferation. Biomaterials 2004, 25 (7), 1215−1224. (78) Park, A.; Cima, L. G. In vitro cell response to differences in poly-L-lactide crystallinity. J. Biomed. Mater. Res. 1996, 31 (1), 17−30. (79) Arima, Y.; Iwata, H. Effects of surface functional groups on protein adsorption and subsequent cell adhesion using self-assembled monolayers. J. Mater. Chem. 2007, 17 (38), 4079−4087. (80) Brodbeck, W. G.; Patel, J.; Voskerician, G.; Christenson, E.; Shive, M. S.; Nakayama, Y.; Matsuda, T.; Ziats, N. P.; Anderson, J. M. Biomaterial adherent macrophage apoptosis is increased by hydrophilic and anionic substrates in vivo. Proc. Natl. Acad. Sci. U. S. A. 2002, 99 (16), 10287−10292. (81) Franz, S.; Rammelt, S.; Scharnweber, D.; Simon, J. C. Immune responses to implants−a review of the implications for the design of immunomodulatory biomaterials. Biomaterials 2011, 32 (28), 6692− 6709.
(82) Lee, P. P.; Desai, T. A. Nitinol-Based Nanotubular Arrays with Controlled Diameters Upregulate Human Vascular Cell ECM Production. ACS Biomater. Sci. Eng. 2016, 2 (3), 409−414. (83) Metavarayuth, K.; Sitasuwan, P.; Zhao, X.; Lin, Y.; Wang, Q. Influence of Surface Topographical Cues on the Differentiation of Mesenchymal Stem Cells in Vitro. ACS Biomater. Sci. Eng. 2016, 2 (2), 142−151.
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DOI: 10.1021/acsbiomaterials.7b00215 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
