Article pubs.acs.org/journal/abseba
Nanocomposite Polymers with “Slimy” Surfaces that Refresh Following Abrasion Wenshou Wang,† Ronald A. Siegel,*,†,‡ and Chun Wang*,† †
Department of Biomedical Engineering, University of Minnesota, 7-105 Hasselmo Hall, 312 Church Street S. E., Minneapolis, Minnesota 55455, United States ‡ Department of Pharmaceutics, University of Minnesota, 9-127E Weaver-Densford Hall, 308 SE Harvard Street, Minneapolis, Minnesota 55455, United States S Supporting Information *
ABSTRACT: Creating polymeric biomaterials with antifouling surface properties that persist after mechanical abrasion is a significant challenge. We report a simple but effective approach based on nanocomposites consisting of a bulk biocompatible polymer, polycaprolactone (PCL), admixed with a minute fraction (1−3 wt %) of nanoparticles consisting of a hyaluronic acid (HA)-PCL graft copolymer (HA-g-PCL). In a nonaqueous solvent such as chloroform, the HA-PCL graft copolymer adopts a reverse-micelle-like structure with a shell dominated by PCL chains, allowing it to be mixed well with high-molecular-weight PCL in the same solvent and cast into a nanocomposite film. Upon exposure to aqueous buffer, the HA-g-PCL nanoparticles reveal the hydrophilic chains of HA to face the outside, conferring a hydrophilic “slimy” or “artificial mucus” layer to the bulk PCL film that resists protein and cell adhesion, without altering bulk mechanical properties. After mechanical abrasion, the nanoparticles replenish the newly exposed material surface with HA, sustaining the surface’s protein/cell resistance. This approach could apply to a wide range of biodegradable polymers to achieve consistent antifouling capacity in the face of mechanical abrasion. KEYWORDS: nanocomposites, surface properties, hyaluronic acid, polycaprolactone
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INTRODUCTION Nonspecific adhesion of endogenous proteins and cells to the surface of materials in a physiological environment, often called “fouling,” can have many negative consequences including thrombosis, chronic inflammation, fibrosis, and infection.1 Fouling can compromise the performance and cause failure of implantable materials and devices. Examples include background interference or change in sensitivity of biosensors and bacterial colonization of contact lenses and indwelling catheters.2,3 It is critical for many biomedical materials applications to achieve persistent antifouling capacity by controlling surface chemistry or topography.2−8 A widely pursued approach to designing antifouling surfaces is to reduce interfacial energy by creating a thin hydrophilic layer at the interface of a bulk material substrate and its biological environment.1 Self-assembled monolayers,9 well-packed zwitterionic compounds 1 0 and polyampholytes or poly(zwitterions),11,12 peptidomimetic polymers,13 cell membranemimicking phosphorylcholine,14,15 ethylene glycol polymers and oligomers,16−18 and many other molecules have been used as coatings on a variety of bulk polymer substrates, endowing them with excellent antifouling capacity.2,3,19−21 However, if either the coating or the polymer substrate is damaged due to mechanical abrasion, the antifouling capacity can be lost. This is a significant © XXXX American Chemical Society
problem insofar as a broad range of biomedical polymers and implants may be subject to mechanical abrasion during processing and handling, and compromise of their antifouling surface properties due to abrasion would compromise their function and biocompatibility. Here, we introduce a simple strategy based on nanocomposites to produce abrasion resistant nonfouling polymeric materials. As a demonstration of principle, we prepared a poly(εcaprolactone) (PCL) matrix, doped throughout with PCLgrafted hyaluronic acid (HA) nanoparticles. These nanoparticles, when at the surface, spread out and form a highly charged, “slimy” surface layer, which resists biofouling. This surface coating refreshes after mechanical abrasion, since the nanoparticles are present throughout the bulk. PCL is a well-known biodegradable and biocompatible polyester22,23 with a wide range of medical applications such as sutures (Monocryl),24 adhesion barriers,25 dental fillings,26 drug delivery systems,27 and tissue engineering scaffolds.28 HA is a negatively charged, highly hydrophilic, natural polysaccharide.29,30 HA-based surfaces have been shown to resist nonspecific Received: April 27, 2015 Accepted: January 5, 2016
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ACS Biomaterials Science & Engineering protein absorption and cell adhesion,31−33 reduce inflammation,34 and prevent postsurgical tissue adhesion.35,36 The approach introduced in this article provides a means to retain the desired bulk features of PCL but endow it with antifouling surface features characteristic of HA, these features being refreshed following abrasion. While studies here are carried out with PCL and HA, the concept of the approach may well extend to many other materials that require a persistent antifouling surface with resistance to mechanical abrasion.
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Dynamic light scattering (DLS) was used to measure average particle size of HA-g-PCL dispersion in water and chloroform. Samples were prepared by dispersing 1 mg of HA-g-PCL into 20 g of water or chloroform by sonication. Average hydrodynamic diameter of the dispersed particles at 25 °C was determined using a ZetaPlus Particle Analyzer (Brookhaven Instruments Corporation, Holtsville, NY; 27 mW laser; 658 nm incident beam, 90° scattering angle). Morphology and average size of the HA-g-PCL particles dispersed in water and chloroform was examined by transmission electron microscopy (TEM) using a JEOL JEM-1210 electron microscope with phosphotungstic acid staining. Differential Scanning Calorimetry (DSC) was carried out over a temperature range of −100 to 100 °C using a TA Q100 calorimeter in a nitrogen atmosphere. Heating and cooling rate was 10 °C/min. Static mechanical tensile stress−strain measurements were performed using a Rheometrics Minimat instrument and MTS Testworks 4 computer software for automatic control of test sequences, data acquisition, and analysis. Dumbbell shaped test specimens were cut from the synthesized PCL/HA-g-PCL composite films and tested at room temperature with a crosshead speed of 20 mm/min according to the ASTM D882-88 standard method. Tests were performed in triplicate. X-ray Photoelectron Spectra (XPS) were collected using a Surface Science SSX-100 spectrometer. The aqueous static contact angle was measured using an image analysis contact angle meter (MCA-3, Kyowa Interface Science Co). Protein Adsorption. FITC-BSA was dissolved in PBS at 10 mg/ mL. PCL and PCL/HA-g-PCL films (5 × 5 × 1 mm, L × W × H) were soaked in the FITC-BSA/PBS solution and placed in a 37 °C incubator. At specified time points, polymer films were removed and washed three times with DI water. Fluorescence micrographs of sample surfaces were recorded using an Olympus IX70 inverted microscope equipped with an Olympus DP72 camera and CellSens software. All sample images were acquired with 10 s of exposure time. The amount of protein adsorbed onto the surfaces was represented by the green fluorescence intensity of each image. Cell Adhesion. PCL and PCL/HA-g-PCL films (5 × 5 × 1 mm) were immersed in PBS at 37 °C for 12 h and then dried and sterilized for 3 h under a UV lamp. NIH 3T3 murine fibroblasts were cultured in DMEM (1 g/L D-glucose, L-glutamine, 110 mg/L sodium pyruvate, Gibco) media supplemented with 10% fetal bovine serum (FBS, heat inactivated, Gibco) and 100 units/mL penicillin/streptomycin (Gibco) and seeded into 24-well plates (0.5 mL media, 20 000 cells per well) in the presence of sterilized PCL or PCL/HA-g-PCL films (5 × 5 × 1 mm, L × W × H, one piece per well). The plate was incubated at 37 °C and 0.5% CO2 for 24 h. The samples were then rinsed in fresh PBS and transferred to a new 24-well plate with 0.5 mL of fresh PBS in each well. Cells attached to the sample surfaces were exposed to Calcein AM for 30 min at 37 °C under CO2, imaged, and counted using an Olympus IX70 inverted microscope. Adherent cells in four different fields of view were counted, and the numbers were averaged.
EXPERIMENTAL SECTION
Materials. HA sodium salt (Na-HA, Mw = 8.88 × 106 g/mol) was purchased from Lifecore Biomedical. Cetyltrimethylammonium (CTA) bromide, ε-caprolactone, 1-butanol, stannous octoate, dibutyltin dilaurate (DBTDL), toluene, hexamethylene-1,6-diisocyanate (HDI), PCL (Mn = 45000), and fluorescein-labeled BSA were all purchased from Sigma-Aldrich. Polymer Synthesis. CTA-HA was synthesized using a method adapted from Pravata et al.37 Briefly, 0.55 g of CTA bromide was dissolved in 7.5 mL of distilled water at 40 °C. This solution was added dropwise to an aqueous solution of 0.6 g of Na-HA (1 wt %) at 40 °C. The formed white precipitate was collected and washed three times with hot water and then dried under a vacuum. To synthesize short chain hydroxy-terminated PCL (PCL−OH), 10 g of ε-caprolactone monomer and 0.11 g of 1-butanol were introduced into a dry flask, followed by 20 mL of toluene and 0.1 g of stannous octoate. After thorough mixing, the flask was placed in an oil bath at a temperature of 130 °C, and the reaction was carried out for 24 h under magnetic stirring. The product was cooled down and poured into 200 mL of diethyl ether, and the precipitate was collected and dried under a vacuum. To synthesize short chain isocyanate-terminated PCL (PCL-NCO), 5 g of PCL−OH and 0.4 g of HDI were introduced into a dried flask, followed by 25 mL of toluene as a solvent and a drop of DBTDL in toluene as a catalyst. The flask was placed in an oil bath at 60 °C with magnetic stirring. After 4 h, the product was precipitated in 100 mL of diethyl ether. The precipitate was collected and dried under a vacuum. To synthesize the graft copolymer HA-g-PCL, 0.3 g of PCL-NCO and 0.3 g of CTA-HA were added into a dried flask with 20 mL of DMSO as a solvent and with a few drops of DBTDL in toluene as a catalyst. The reaction was conducted at 60 °C for 4 h. The product was poured into 80 mL of acetone. The precipitate was collected and washed three times with acetone then dried under a vacuum. The CTA group of the graft copolymer was removed according to Pravata et al.,37 in which 0.3 g of CTA-HA-g-PCL was dissolved in 30 mL of a 2:1 (v/v) phosphate buffer (0.3 M pH = 7.4)/DMSO mixture and dialyzed (MW cutoff = 3500 Da) against DMSO (1 day), ethanol (1 day), water (1 day), ethanol (1 day), and finally water (3 days). The product was recovered by freeze-drying. Preparation of Nanocomposite Films and Mechanical Abrasion. To prepare a nanocomposite film containing 1 wt % of HA-g-PCL nanoparticle, PCL (Mn = 45000, 2 g) and HA-g-PCL (20 mg) were dissolved in 40 mL of chloroform and sonicated for 30 min using a Sonic Dismembrator (model 100, Fisher Scientific). The resulting suspension was poured into a glass Petri dish, and the solvent was evaporated to produce a polymer film, which was then molded by compression at 100 °C to form a smooth film. The thickness of the film was measured using a digital caliper. Nanocomposite films containing 3 wt % nanoparticle and pristine PCL films were prepared similarly. To simulate abrasion, the composite films were polished to various thicknesses using 3 M Wetordry Sandpaper (Extra Fine 320 Grit) and cleaned with ethanol. Characterization of Polymers, Nanoparticles, and Surfaces. 1 H NMR spectra of polymers were recorded on a Varian Unity spectrometer (300 MHz) with CDCl3 or DMSO-d6 as a solvent. FTIR was conducted using a Nicolet Series II Magna-IR System 750 with OMNIC software for data collection and analysis.
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RESULTS AND DISCUSSION Rationale for Material Design. Our strategy for creating a renewable HA-coated antifouling surface is possible due to the dynamic structure of HA-g-PCL nanoparticles. As depicted in Figure 1, in a nonaqueous solvent such as chloroform, HA-g-PCL would self-assemble into nanoparticles similar to reverse micelles with a hydrophobic outer layer of PCL chains. However, upon exposure to aqueous buffer, the HA-g-PCL nanoparticles would “flip” and adopt a regular micellar structure with a hydrophilic shell of HA and a hydrophobic core of PCL. We further hypothesize that in the presence of a solid PCL substrate, the amphiphilic HA-g-PCL would preferentially adhere to the solid− water interface, conferring a hydrophilic “slimy” or “artificial mucus” layer to the bulk PCL film (Figure 1). Figure 2 illustrates the overall approach to an antifouling HAbased surface coating for PCL that renews upon mechanical abrasion. High molecular weight PCL is dissolved in a nonaqueous solvent such as chloroform and mixed with a
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First, short chain PCL was synthesized by ring opening polymerization of caprolactone with 90% yield and an average Mn of 7400 (DP = 66), as determined by 1H NMR (Scheme 1A, Figure S1). Initiation by 1-butanol resulted in semitelechelic PCL with one hydroxyl end-group per polymer chain (PCL−OH). Subsequent conversion of the hydroxyl end group to isocyanate (PCL-NCO) was confirmed by FTIR, showing the characteristic isocyanate vibrational peak at 2270 cm−1 (Figure S2). Second, HA-g-PCL (Scheme 1B) was synthesized by grafting short PCL chains onto HA (Mn = 888 000, equivalent to 2220 disaccharide repeat units). The HA-g-PCL graft copolymer (as cetyltrimethylammonium, CTA, salt) was characterized by 1H NMR in DMSO-d6 and compared with “bare” HA CTA salt. The degree of grafting was calculated to be 1.4% according to 3Ib/(2Ia × 66) × 100%, where Ia and Ib are the areas of 1H NMR peaks of “a” (methylene group of PCL) and “b” (acetyl group of HA), respectively (Figure S3). Thus, the weight fraction of PCL in the graft copolymer was 16%, and on average, there were 31 PCL grafts per HA chain. Taken together, the average molecular weight of the graft copolymer is estimated to be M = 888 000 + 31 × 7400 = 1.12 × 106 (g/mol). HA-g-PCL Copolymer Self-Assembles into Stable Nanoparticles in Both Aqueous and Non-Aqueous Solvents. The amphiphilic graft copolymer HA-g-PCL dispersed in either water (a good solvent for HA) or chloroform (a good solvent for PCL) after ultrasonication. DLS detected the presence of solvent swollen nanoparticles with an average diameter between 475 and 550 nm in water and 500−600 nm in chloroform (Figure 3A). The size of these nanoparticles remained relatively unchanged over several hours in either water or chloroform (Figure 3A), suggesting that the amphiphilic nature of the graft copolymer sustains colloidal stability in both solvents. After drying, the nanoparticles were visualized by TEM, which shows that they are largely round (Figure 3B). The average
Figure 1. Schematic illustration of the amphiphilic HA-g-PCL copolymer (A) and its dynamic self-assembly in nonaqueous and aqueous solvents as well as at solid PCL/water interface (B).
small amount of HA-g-PCL copolymer, which forms reversemicelle-like nanoparticles. The hydrophobic PCL segment of the graft copolymer ensures good miscibility with bulk PCL and even distribution within the PCL film after drying. When exposed to an aqueous environment, the HA-g-PCL reverse micelles “flip” and anchor the hydrophilic HA chains to the surface of the bulk PCL, resisting nonspecific adhesion by proteins and cells. When the material surface is mechanically removed, the HA-g-PCL graft copolymer nanoparticles replenish the new surface with a refreshed hydrophilic “slime” or “artificial mucus” layer that continues to resist protein and cell adhesion (Figure 2). HA-g-PCL Copolymer Synthesis and Characterization. The synthetic pathway for HA-g-PCL is shown in Scheme 1.
Figure 2. A nanocomposite approach to an HA-coated antifouling surface of PCL matrix that refreshes following mechanical abrasion. C
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ACS Biomaterials Science & Engineering Scheme 1. Chemical Synthesis of Short-Chain Semitelechelic PCL Isocyanate (A) and HA-g-PCL Copolymer (B)
approximately 10 times larger in diameter (compare Figure 3A and B), it appears that the polymer chains take up a very small fraction, on the order of 1/1000, of the space in the suspended aggregates. We note, however, that the dominance of scattering by large uniformly composed aggregates may lead to an overestimation of average properties. An improved characterization of composition of the nanoparticles in the dry state and in suspension, in which the several assumptions made are tested more rigorously, is a candidate for future efforts. Bulk and Surface Characterization of Nanocomposite Films. Nanocomposite films were prepared by solvent casting mixtures of pristine high MW PCL (100, 99, or 97 wt %) and the HA-g-PCL graft copolymer (0, 1, or 3 wt %) in chloroform. All the films were macroscopically homogeneous. Thermal and mechanical properties of pristine PCL and nanocomposite films are summarized in Table 1. DSC revealed a melting point of 55− 56 °C for the films. Compared to the pristine PCL, the nanocomposite materials had slightly lower crystallinity, higher tensile strength and modulus, and lower elongation at break. As the content of the dispersed graft copolymer nanoparticles increased from 1 to 3%, changes in these bulk properties were more pronounced but never deviated by more than 15% from their values for pristine PCL. Thus, inclusion of the dispersed graft copolymer nanoparticles in small quantities (up to 3%) did not substantially alter the bulk properties of the matrix material. To confirm the presence of HA coating, XPS (1s) spectra of the films were acquired and shown in Figure 4A. Compared to the pristine PCL surface, the nanocomposite surfaces had slightly
sizes of the dried particles, based on the TEM images, were approximately 50 and 100 nm for particles dried from water and chloroform, respectively. These observations support the hypothesis, illustrated in Figure 1, that the HA main-chains and the PCL grafts populate the outer layer of the nanoparticles to maintain colloidal stability in aqueous and nonaqueous solvents, respectively. Colloidal stability of the HA-g-PCL copolymer in chloroform is particularly important in ensuring even mixing with high molecular weight PCL and the formation of homogeneous nanocomposite films, as described below. It is of some interest to estimate the number of graft copolymer chains per nanoparticle. There is no reason to assume that this number will be the same for nanoparticles that were initially suspended in water versus chloroform. If we may assume that the nanoparticles shown in the TEM images (Figure 3B) are essentially free of solvent, and are approximately spherical in shape (i.e., not oblates), then the following formula can be used to estimate the number of chains, N, in a nanoparticle of diameter d consisting of graft copolymer chains of molecular weight M and density ρp: N = (π/6)d3/[M/(NAρp)], where NA is Avogadro’s number. Based on an estimated molecular weight M = 1.12. × 106 g/mol and a density of approximately 1.1 g/cm3 for the grafted polymer chains, we calculate that N ≈ 40 and N ≈ 300 for dry nanoparticles of diameters 50 and 100 nm, which are representative of deposited nanoparticles from water and chloroform, respectively (see Figure 3B). Assuming that the nanoparticles have the same number of chains in suspension as in the dry state, and noticing that the suspended nanoparticles are D
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Figure 4. Surface characterization of nanocomposites and pristine PCL by XPS (A) and static water contact angle measurement (B). Numbers in A indicate the relative abundance of the elements in %.
To investigate whether the nanocomposite surface was resistant to nonspecific protein adsorption, FITC-labeled bovine serum albumin (BSA) was used as a model protein. After incubation with a 10 mg/mL aqueous solution of FITC-BSA at 37 °C for 32 h, the surfaces of the nanocomposite film containing 3% graft copolymer and the pristine PCL were examined by fluorescence microscopy (Figure 5A, left). The intensity of the green fluorescent signal from representative images of the surfaces was quantified as a measure of the amount of BSA absorbed. Clearly, BSA adsorption to the nanocomposite surface was greatly diminished compared to the pristine PCL surface (Figure 5A, right). The antifouling capacity of the nanocomposite surfaces was further tested using NIH 3T3 fibroblasts. Pristine PCL and the nanocomposite films containing 1 or 3% graft copolymer nanoparticles were immersed in aqueous buffer (pH 7.4) at 37 °C for 1 or 4 weeks, incubated with fibroblasts in complete serum-containing media for 24 h, and surface-adhering cells were counted. While many cells were found adhering to the tissue culture plate (TCP) and pristine PCL surface, very few cells, if any, were found on the surfaces of the nanocomposites, even after 1 and 4 weeks of exposing the films to buffer (Figure 5B,C). Cells found on the TCP and pristine PCl surfaces were spread out, whereas those very few cells on the nanocomposite surfaces were round and loosely bound. These findings demonstrate that the nanocomposite surface is highly resistant to cell adhesion. To simulate a loss of surface coating due to repeated mechanical abrasion, the surface of the nanocomposite film containing 3% graft copolymer nanoparticles was milled and polished to different depths, and the exposed “neo-surface” was characterized for a change in hydrophilicity and cell adhesion (Figure 6A). The values of the water contact angle of the “neosurface” at 0.5, 0.6, and 0.8 mm below the initial surface were consistently around 65° and around approximately 20° lower than that of the pristine PCL surface (Figure 6B), suggesting the presence of HA coating on the “neo-surface” that became more hydrophilic. Compared with the nanocomposite surface without mechanical abrasion, the “neo-surface” had similar degrees of
Figure 3. HA-g-PCL graft copolymer. The copolymer forms nanoparticles in water and in chloroform. (A) Time-dependent changes in the diameter of solvent-swollen nanoparticles monitored by dynamic light scattering. (B) TEM images of graft copolymer nanoparticles dried from different solvents. Arrows point to the polymer nanoparticles.
Table 1. Thermal and Mechanical Properties of Pristine PCL and Nanocomposite Films sample PCL PCL/1% HA-gPCL PCL/3% HA-gPCL
tensile strength [MPa]
modulus [MPa]
elongation at break [%]
melting point [°C]
crystallinity [%]
18 ± 5 20 ± 7
310 ± 25 330 ± 43
950 ± 50 880 ± 41
55 56
58 55
21 ± 4
350 ± 60
790 ± 54
56
50
lower C content and slightly higher O content. A small peak at 392 eV corresponding to N1s was observed for the nanocomposite surfaces but was absent in the pristine PCL. The relative N content of the nanocomposite surfaces (0.2%, 0.5%) was essentially proportional to the theoretical content of HA-gPCL dispersion in the bulk (1%, 3%). These results confirm the presence of the HA layer on the surface of the nanocomposites. The static contact angle of the pristine PCL surface in water was ∼80°. However, the contact angle of the nanocomposite surfaces was reduced to 55−60° (Figure 4B). The increase in surface hydrophilicity is likely due to the presence of the HA chains interacting with the PCL substrate through the shortchain PCL grafts, as illustrated in Figure 2. Contact angle of the nanocomposite surfaces did not change over the course of one month, suggesting that the surface coating of HA chains remained highly stable (Figure 4B). E
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Figure 5. Sustained antifouling properties of the nanocomposite films. (A) Surface adsorption of FITC-BSA protein after 32 h at 37 °C. (B) Representative images and (C) quantification of NIH 3T3 fibroblast adhesion to various material surfaces after immersion in water at 37 °C for 1 and 4 weeks. TCP: tissue culture plate. *p < 0.004, comparing nanocomposites to pristine PCL, two-tailed t test.
and a persistent hydrophilic antifouling surface, with a nearly 20° reduction in water contact angle and much reduced protein adsorption and cell adhesion. The amount of nanoparticles needed for such effects was minute, with minimal impact on the bulk properties of the PCL matrix, and the nanocomposite material withstood several cycles of mechanical abrasion without losing its antifouling surface properties. Our approach potentially provides more flexibility and ease in material fabrication and processing, without requiring extensive chemical modification of the bulk materials as reported in other self-replenishing PCL surfaces. 41 Moreover, the present approach presents a replenishable hydrophilic surface on a hydrophobic bulk substrate, in contrast to previous efforts.41 More extensive analysis is required to further elucidate the structural feature of the nanocomposites. Although there is no doubt that an HA coating has been generated on the nanocomposite surface, it is not clear whether the HA chains, as part of the surface-bound nanoparticles, cover the surface completely or partially. Such information as well as detailed surface topography will require further analysis by methods such as AFM. With demonstrated stability of the nanoparticles without aggregation (Figure 3A), even mixing of the nanoparticles with high molecular weight PCL before film casting should result in homogeneous distribution of the nanoparticles inside the PCL matrix. Furthermore, protein adsorption and cell
hydrophilicity with average water contact angle values of approximately 5° higher but without statistical significance (Figure 6B). At the same time, the exposed “neo-surface” at all milling depths were as resistant to adhesion of NIH 3T3 fibroblasts as the unpolished surface (Figure 6C). Collectively, these results demonstrate that a mechanically abraded nanocomposite surface can maintain the hydrophilicity and antifouling property of the HA coating provided by the HA-gPCL nanoparticles trapped in the bulk of the PCL matrix. In summary, we have shown that a hydrophobic biomedical polymer (PCL) can be endowed with hydrophilic surface properties that persist following mechanical abrasion. This was achieved by admixing a small amount of graft copolymer nanoparticles consisting of a hydrophilic backbone (HA) and grafted side-chains made of the same polymer as the bulk material (PCL). Compared to other surface modification approaches involving polymer blends,38−40 our approach is different in that (1) the HA-g-PCL nanoparticles are present throughout the bulk of the PCL, rather than only on the surface, and (2) the amphiphilic nature of the HA-g-PCL copolymer should allow well distributed mixing of nanoparticles with the hydrophobic PCL matrix and effective display of a hydrophilic coating in aqueous environments. Consequently, we have shown that adding as little as 1% but no more than 3% of the nanoparticles is sufficient to maintain bulk mechanical properties F
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that erode due to chemical or enzymatic degradation. These polymers may include polylactide and polyglycolide and their copolymers, which are bulk eroding, as well as surface eroding polymers such as polyorthoesters and polyanhydrides. On the other hand, other polysaccharides such as heparin, chondroitin sulfate, and dextran,42 grafted with short chains of the host polymer, could replace HA to make up the “slimy” material surface, while avoiding the potential drawbacks of PEGylated surfaces.43 Future studies will also focus on evaluation of antifouling properties of such nanocomposite materials in vivo.
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CONCLUSIONS We have developed a nanocomposite-based approach to create persistent antifouling surfaces for polymers subjected to mechanical abrasion. Blending PCL with a minute amount of HA-g-PCL graft copolymer nanoparticles results in a HA-coated surface that resists protein adsorption and cell adhesion. Even when the surface of the material is lost due to repeated cycles of mechanical abrasion, the HA-g-PCL graft copolymer replenishes the surface and maintains the antifouling properties. Although the results reported here are obtained using PCL and HA, this strategy could be generally applicable to other biomedical polymers and to restore antifouling surface properties following mechanical abrasion due to material processing and handling as well as in vivo application.
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ASSOCIATED CONTENT
S Supporting Information *
The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsbiomaterials.5b00182. NMR and FTIR spectra (PDF)
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AUTHOR INFORMATION
Corresponding Authors
Figure 6. (A) Mechanical abrasion of the nanocomposite film to different depths. (B) Water contact angle of the abrased surfaces. (C) Quantification of fibroblast adhesion to the abrased surfaces. *p < 0.0005, comparing nanocomposite film to pristine PCL film, two-tailed t test.
*E-mail:
[email protected]. *E-mail:
[email protected]. Author Contributions
The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript.
adhesion experiments indicate consistent antifouling properties of the initial surface and the subsequent “neo-surface” at various depths, which is in accordance with the notion of even distribution of nanoparticles in the bulk. Direct visualization of nanoparticles in the bulk by sectional TEM should resolve this issue. The exact mechanism of the renewable antifouling nanocomposite surface deserves further understanding. In contrast to the common “surface modifier” approach,39 the HA-g-PCL nanoparticles are not expected to migrate through the bulk of the PCL matrix, given their large size. A more plausible scenario is that the nanoparticles are dispersed and trapped within the hydrophobic PCL matrix, and mechanical abrasion removes the surface layer to expose other nanoparticles underneath to the aqueous environment, forming a new layer of hydrophilic HA coating on the “neo-surface.” Morphology and stability of the matrix-trapped nanoparticles in extended periods of time in an aqueous medium may provide further insight into the validity of the hypothesized mechanism. We envision that this nanocomposite strategy toward renewable surfaces may be applicable to not only polymeric materials subjected to mechanical abrasion but also to polymers
Notes
The authors declare no competing financial interest.
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ACKNOWLEDGMENTS This work was partially supported by IPRIME at the University of Minnesota. REFERENCES
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DOI: 10.1021/acsbiomaterials.5b00182 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
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DOI: 10.1021/acsbiomaterials.5b00182 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX
