Biomaterials 45 (2015) 64e71 Contents lists available at ScienceDirect Biomaterials journal homepage: www.elsevier.com/locate/biomaterials Self-defensive antibacterial layer-by-layer hydrogel coatings with pH-triggered hydrophobicity Yiming Lu a, Yong Wu b, Jing Liang b, Matthew R. Libera b, Svetlana A. Sukhishvili a, * a b Department of Chemistry, Chemical Biology and Biomedical Engineering, Hoboken, NJ 07030, USA Department of Chemical Engineering and Materials Science, Stevens Institute of Technology, Hoboken, NJ 07030, USA a r t i c l e i n f o a b s t r a c t Article history: Received 11 October 2014 Accepted 20 December 2014 Available online We report on negatively charged layer-by-layer (LbL) hydrogel films, which turn hydrophobic and bactericidal in response to bacteria-induced acidification of the medium. Single-component hydrogel thin films, abbreviated as PaAALbLs, consisting of chemically crosslinked poly(2-alkylacrylic acids) (PaAAs) with varying hydrophobicity [polymethacrylic acid (PMAA), poly(2-ethylacrylic acid) (PEAA), poly(2-n-propylacrylic acid) (PPAA) or poly(2-n-butylacrylic acid) (PBAA)]. With increasing polyacid hydrophobicity, the hydrogel films showed a decrease in water uptake and an increase in elastic modulus. Both parameters were strongly dependent on pH. At pH 7.4, hydrogels of higher hydrophobicity were more resistant to colonization by Staphylococcus epidermidis, with the PBAA coating showing almost negligible colonization. As the medium became more acidic due to bacterial proliferation, the more hydrophobic PEAALbL, PPAALbL and PBAALbL hydrogels became dehydrated and killed bacteria upon contact with the surface. The killing efficiency was strongly enhanced by the polymer hydrophobicity. The films remained cytocompatible with human osteoblasts, as indicated by the MTS assay and live/dead staining. Our approach exploits bacteria-responsive properties of the coating itself without the involvement of potentially toxic cationic polymers or the release of antimicrobial agents. These coatings thus demonstrate a novel approach to the antibacterial protection of tissue-contacting biomedical-device surfaces. © 2014 Elsevier Ltd. All rights reserved. Keywords: Bacterial adhesion Surface modification Antibacterial Hydrogel 1. Introduction Because of the growing clinical importance of biomaterialsassociated infection [1], significant research has recently focused on various ways to modify implant surfaces using coatings that prevent bacterial adhesion, kill bacteria on contact, or enable controlled elution of antimicrobial compounds [2]. Many such coatings involve highly hydrated polymers, such as poly(ethylene glycol) (PEG) [3]. Such coatings exhibit antifouling properties and significantly reduce both protein adsorption and bacterial adhesion due to the high entropy and strong affinity of PEG for water. In contrast, superhydrophobic surfaces also display antifouling properties and can inhibit microbial adhesion [4]. However, for many tissue-contacting biomedical devices, the potential of these continuous antifouling coatings is somewhat limited since they * Corresponding author. E-mail address: [email protected] (S.A. Sukhishvili). http://dx.doi.org/10.1016/j.biomaterials.2014.12.048 0142-9612/© 2014 Elsevier Ltd. All rights reserved. indiscriminately suppress interactions not only with bacteria but also with mammalian cells [5]. Other antibacterial coatings have been designed with various molecular architectures. Among these, for example, are selfassembled monolayers [6], polymer brushes [7], and LangmuireBlodgett films [8]. Furthermore, polycations, including chitosan [9], are known to be antibacterial, and this property is attributed to their positively charged chains, which can efficiently disrupt negatively charged bacterial walls. A polycation-based approach has been developed by Klibanov and co-workers, who showed that surfaces coated with poly(N-alkyl-pyridinium) or N-alkyl-polyethyleneimine efficiently kill gram-positive and gram-negative bacteria upon contact [10]. Another approach exploits the versatility of the layer-by-layer (LbL) technique to create polyelectrolyte multilayer coatings of diverse chemical compositions, architectures, and functions [11]. LbL films include a powerful portfolio of multifunctional antibacterial and antibiofilm coatings [12], and they can exhibit significant antibacterial activity due to their polycationic constituents [13,14], dual anti-adhesion and antibacterial activity achieved by the Y. Lu et al. / Biomaterials 45 (2015) 64e71 simultaneous incorporation of hydrophilic polymers and polycations [13,15], or dual contact killing and antimicrobial release achieved by the inclusion, for example, of polycations and silverion-releasing species [16]. Despite the ability of such films to provide for controlled antimicrobial release [17,18], continuous-elution strategies face several challenges. Among them is the fact that the antimicrobial supply in continuously eluting coatings will ultimately deplete. Since they elute antimicrobials whether those antimicrobials are needed or not, continuously eluting coatings furthermore contribute to the selection of antibiotic-resistant bacteria [19] and can lead to undesirable cytotoxicity effects. Therefore, alternative strategies have recently been developed that create antibacterial coatings which are responsive to various environmental triggers, including pH [20]. We have recently developed the concept of bacteria-triggered delivery of antibacterial agents from LbL coatings. These coatings release antibacterial agents in response to the local acidification from natural bacterial metabolic processes [20]. In contrast to continuously eluting coatings, these new coatings are selfdefensive because they release their antimicrobial payload only when and where bacteria interact with the surface. Another type of self-defensive coating has been developed by G. Cado et al. This coating responds to enzymatic molecules secreted by growing biofilms [21] rather than to pH. Here, we continue to exploit local acidification as a trigger to activate antibacterial functions. However, in contrast to our previous work where bacteria-induced pH variations trigger the release of antibacterial agents [20], here we exploit a new concept where the polymer matrix itself becomes antibacterial in response to bacterial growth. These new biologically active coatings do not include any antibacterial compounds such as small-molecule antibiotics. They are based on thin-film hydrogels made from a family of polyanionic poly(2-alkylacrylic acids) (PaAAs) whose hydrophobicity varies depending on the length of their alkyl side chain. The net negative charge and relatively large water uptake by these coatings under normal physiological conditions (pH 7.4) make these PaAA films non-toxic to osteoblasts. At the same time, the more hydrophobic members of this family are selectively toxic to staphylococcal bacteria because of pH-triggered changes in coating charge and water content. Previous studies to design drug-delivery systems have exploited the capacity of dissolved hydrophobic PaAA homopolymers and copolymers to become membrane-disruptive in acidic endosomal and tumor environments [22]. However, the potential of this class of polymers as antibacterial agents has not yet been explored. We thus study the pH-dependent antibacterial properties of a series of poly(2-alkylacrylic acid) polymers with varying hydrophobicity when these polymers are either dissolved in aqueous solutions or are included within single-component layer-by-layer hydrogel thin-film coatings, which we designate here as PaAALbL. We have recently reported on similar PaAALbL hydrogel coatings which can control the wettability of micro-structured substrates and detect small-molecule adsorption [23]. The present work focuses on the capability of these coatings to selectively control adhesion and viability of bacteria and mammalian cells on planar surfaces. 65 experiments. Undoped silicon substrates of 525 ± 25 mm thickness were purchased from Virginia Semiconductor Inc. 2.2. Methods 2.2.1. Polymer synthesis and characterization Poly(2-ethylacrylic acid) (PEAA) (Mw ~ 79,000 g/mol, polydispersity index (PDI) 2.52), poly(2-n-propylacrylic acid) (PPAA) (Mw ~ 110,000 g/mol, PDI 2.18), and poly(2-n-butylacrylic acid) (PBAA) (Mw ~ 138,000 g/mol, PDI 2.41) (Scheme 1) were synthesized by free-radical polymerization using a procedure similar to that described elsewhere [24]. Briefly, 1 g of 2-alkylacrylic acid monomers and 0.2 mL DMF were mixed in Schlenk vials prior the addition of 3 mg of AIBN. The vials were flame-sealed under vacuum and submerged in an oil bath at 60 C after 3 freezeethaw cycles. The polymerization was stopped after 24 h by breaking the vial to allow oxygen access, followed by immersion of the vial in liquid nitrogen for 2 min to stop the reaction. The molecular weights of the PaAAs were determined by GPC at a flow rate of 1 mL/min and 60 C using 0.01 mM LiBr DMF as solvent and polystyrene standards for column calibration. 2.2.2. Construction of surface hydrogels Silicon substrates (1 cm 1 cm in size) were cleaned by 12 h exposure to UV light followed by treatment with piranha solution (1:3 v/v H2O2:H2SO4) at 80 C. Prior to multilayer construction, a bilayer of PMAA and BPEI, designated as (PMAA/ BPEI)2, was deposited and crosslinked [20]. PaAA/PVPON multilayers were then deposited by sequential adsorption of PaAA and PVPON from 5 mL of 0.5 mg/mL PaAA and PVPON solutions in methanol, except for PMAA/PVPON films, which were deposited from 0.01 M phosphate buffer solutions at pH 2. The deposition time per polymer layer was 10 min. In between the PaAA and PVPON deposition steps, the wafers were rinsed for 30 s with a corresponding solvent (0.01 M phosphate buffer at pH 2 for PMAA/PVPON films, and methanol for all other PaAA/PVPON multilayers). All films were dried with nitrogen gas flow after completion of multilayer assembly. After EDC (3 mg/mL) activation at pH 5.2, the PaAA components of the multilayers were crosslinked by AADH (3 mg/mL) for 3 h. To release uncrosslinked PVPON, the substrates were immersed in pH 7.5 buffer (for PMAALbL) or in a mixture of pH 6 0.005 M phosphate buffer and DMSO (1:5 v/v) (for PEAALbL, PPAALbL and PBAALbL) for 30 min. The substrates were then washed with water and methanol and dried. 2.2.3. Ellipsometry The thickness of the dry PaAALbL hydrogels was measured by phase-modulated ellipsometry with an incident angle of 65 assuming a constant film refractive index of 1.5. The optical properties of the substrates and the native oxide layer were determined prior to polymer deposition. Measurements of PaAALbL hydrogel swelling were performed using a custom-made cylindrical flow-through glass cell. For wet film measurements, the cell was filled with 0.01 M phosphate buffer at various pH values, and thickness and refractive-index measurements were taken after the hydrogels were allowed to equilibrate for 20 min. 2.2.4. FTIR spectroscopy FTIR spectra of PaAALbL hydrogels were acquired using a Tensor 27 FTIR spectrometer (Bruker Optik GmbH). All samples were pre-titrated by phosphate buffer solutions and dried from solutions at varied pH. 2.2.5. Contact angle (CA) measurements Static СAs were measured under ambient conditions (~25 C, relative humidity ~23%) using a Contact Angle and Surface Tension Meter CAM 101 (KVS Instrument Inc.). Prior to a CA measurement at a certain pH value, samples were pretreated by 0.01 M phosphate buffer solution with the same pH as used in the CA measurements. Measurements were repeated 3 times at different positions of the samples. 2. Materials and methods 2.1. Materials Polymethacrylic acid (PMAA) (Mw ~ 56,000 g/mol), branched poly(ethylene imine) (BPEI) (Mw ~ 101,000 g/mol), polyvinylpyrrolidone (PVPON) (Mw ~ 2500 g/ mol), monobasic sodium phosphate, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), adipic acid dihydrazide (AADH), azobisisobutyronitrile (AIBN), hydrochloric acid, sodium hydroxide, methanol, N,N-dimethylformamide (DMF) and dimethyl sulfoxide (DMSO) were purchased from Sigma. LIVE/DEAD® BacLight™ Bacterial Viability Kit was purchased from Life Technologies Corp. Millipore (Milli-Q) type I water with a resistivity of 18.2 MU was used in all Scheme 1. Chemical structures of poly(2-alkylacrylic acids) (PaAAs). The abbreviations, PMAA, PEAA, PPAA and PBAA stand for polymethacrylic acid, poly(2-ethylacrylic acid), poly(2-n-propylacrylic acid), and poly(2-n-butylacrylic acid), respectively. 66 Y. Lu et al. / Biomaterials 45 (2015) 64e71 2.2.6. Zeta-potential measurements Zeta potentials of bacterial cells were measured using Zetasizer Nano-ZS equipment (Malvern Instruments, Ltd). These measurements were performed with a 106 CFU/mL bacterial concentration in 0.05 M phosphate buffer at 37 C. 2.2.7. Atomic force microscopy (AFM) AFM was used to characterize the stiffness of hydrated PaAALbL hydrogels at both low and high pH using an NSCRIPTOR™ dip pen nanolithography system (Nanoink). Aspire™ CCS conical contact mode short cantilevers (NanoScience Instruments, Inc.) with a force constant of 0.1 N/m were used for nanoindentation. Under dry conditions, the cantilevers were calibrated using the thermal fluctuation technique. For wet measurements, cantilevers were calibrated in 0.05 M phosphate buffer solutions with 0.2 M NaCl. All force curves were collected in situ using a cantilever velocity of 0.01 mm/s. Force curves were analyzed based on the conical geometric shape of the cantilever. Young's moduli of PaAALbL hydrogels measured by nanoindentation, Eind, were calculated using the following equation [25]: F¼ 8Eind tan ad2 3p ( 1 þ 1:7795 2 tan a d d2 þ 16ð1:7795Þ2 tan a 2 h p2 h ) humidity. The complete medium was replaced every 2 days, and confluent flasks were subcultured using trypsin/EDTA. The cells were then seeded in 24-well culture plates (10,000 cells/cm2) onto PaAALbL hydrogels in 2 mL of medium. The medium was refreshed every 2 days. Each experiment was performed in quintuple (n ¼ 5) for each group of substrates. Cell viability was assessed by a colorimetric assay based on the conversion of the tetrazolium salt into a soluble formazan dye using a Cell Titer 96 Aqueous assay (MTS, Promega). After 1, 3, and 7 days, the culture medium was removed from each well, and the specimens were transferred to new cell culture plates. 100 mL of pre-warmed MTS solution with 1 mL culture medium was added to each well with continuous culture for 3 h. 200 mL of supernatant from each well was then transferred to a 96-well plate, and the absorbance was measured at 490 nm. Four specimens for each group were tested, and each test was repeated three times. The cytotoxicity of the PaAALbL hydrogel thin films was also evaluated using a live/dead cell viability assay kit (Invitrogen) after 1 and 7 days. The cells were stained based on membrane integrity and intracellular esterase activity per the manufacturer's protocols. Briefly, cell-seeded substrates were incubated at room temperature for 30 min in PBS containing 2 mM of calcein AM (green) and 4 mM of ethidium homodimer-1 (red). Cells stained green (live) and red (dead) were imaged using a Nikon C1 confocal microscope. where F is the applied force, d is the indentation depth, a is the half angle (15 ) of the conical cantilever, and h is the sample height. This equation takes into account the effects of the incompressible underlying substrate on the measurement of the film's elastic properties. 3. Results and discussion 2.2.8. Antibacterial activity assay Staphylococcus epidermidis (ATCC 35984) was incubated in tryptic soy broth (TSB) with 6 mg/mL yeast extract and 8 mg/mL glucose (TSB) for 18 h. The resulting biofilm was vortexed, sonicated, and filtered (5 mm) in order to create a near singlecell suspension. To test the antibacterial of PaAAs in solution, 1 mL of suspension (2 106 CFU/mL) was mixed with 1 mL of PaAA dissolved in TSB (2.5 mg/mL) at pH 7.5. The bacteria-polymer solutions were then incubated for 24 h at 37 C in 12-well culture plates. Bacterial suspensions (1 106 CFU/mL) in TSB, solutions of PaAAs (1.25 mg/mL), and pure TSB solution were used as controls. The final bacterial concentrations were determined using calibrated optical densities corrected for the absorbance characteristic of the control samples. To test the antibacterial activities of PaAALbL hydrogel thin films, all Si wafers (with or without coatings) were exposed to UV light for 5 min prior to bacterial inoculation. Then 2 mL of bacterial suspension (1 106 CFU/mL) was seeded onto each PaAALbL hydrogel thin film followed by incubation for various time periods (6, 10, 12, 14, 16, 24 h) at 37 C in 12-well culture plates. Samples were then rinsed with phosphate buffered saline (PBS) twice before fluorescence labeling. Using a live/ dead bacterial viability test kit (Invitrogen), stained specimens were characterized using a Nikon C1 confocal laser scanning microscope (CLSM) system. To determine the bacterial coverage on PaAALbL hydrogels, 2D CLSM images of live/dead stained bacteria were analyzed using ImageJ software. Bacterial coverage was defined as the ratio: Abact.,total/Asurface, where Abact.,total is the surface area covered by bacteria, both live and dead, and Asurface is the area of the entire surface. The fraction of dead bacteria was determined as the ratio of the area covered by dead bacteria, Abact.,dead, to the total area covered by both dead and live bacteria, Atotal. Because PEAA, PPAA and PBAA are not soluble in aqueous acidic media, hydrogen-bonded multilayer films of hydrophobic PEAA, PPAA, or PBAA with polyvinylpyrrolidone (PVPON) were deposited using methanol as the solvent. PMAA/PVPON films were deposited from aqueous solutions at pH 2. In all cases, the film thickness increases linearly as the number of deposited bilayers increases, but individual bilayer thickness depends on the type of PaAA. For PMAA (hydrophilic), the PMAA/PVPON bilayer thickness is ~5.3 nm, while the bilayer thicknesses of films involving the other three PaAAs (hydrophobic) are only ~3.6 nm. Because of the coupling of thinfilm mechanical properties to the underlying solid substrates [26], we constructed films all having the same thickness of 54 ± 1 nm by depositing 10 bilayers of PMAA/PVPON or 15 bilayers of the more hydrophobic PaAA/PVPON films (Figure S1A). After crosslinking of the self-assembled films and the subsequent pHinduced release of uncrosslinked PVPON, the film thicknesses decrease by ~35% to ~35 nm when dried (Figure S1A). FTIR spectra of dry hydrogen-bonded PMAA/PVPON multilayers and dry PMAALbL hydrogel thin films show a complete release of PVPON, which is indicated by the significant decrease of the amide band intensity at ~1650 cm1 (Figure S1B). The remaining absorbance at ~1650 cm1 is due to amide bonds introduced through covalent attachment of AADH crosslinker to the PaAA. FTIR spectra of the other crosslinked PaAALbL films demonstrate the same (23e26%) fractional contribution of amide band intensities to the total 2.2.9. Cell culture and cell viability assays Human fetal osteoblasts (hFOB 1.19, ATCC CRL-11372) were cultured in DMEM containing 10% FBS and 1% penicillin/streptomycin at 37 C in 5% CO2 and 95% 3.1. PaAALbL thin-film hydrogels: synthesis and properties Fig. 1. (A) FTIR spectra of PaAALbL hydrogel films. The bands centered at 1550 cm1 and 1700 cm1 represent carboxylate (>COO) and carboxylic (>COOH) groups, respectively. The absorbance band at 1650 cm1 corresponds to amide (eNHeCOe) groups introduced via PaAA crosslinking. (B) Representative curve fitting of a PEAALbL FTIR spectrum. Y. Lu et al. / Biomaterials 45 (2015) 64e71 absorbance in the 1500e1800 cm1 region, which also includes a eCOO band at 1550 cm1 and a eCOOH band at ~1700 cm1 (Fig. 1). This suggests that all PaAALbL films have similar crosslinking degrees and, consequently, that hydrogel swelling is determined primarily by the solubility of the polymer chains between crosslinks. The wetting and swelling properties of the various hydrogel films depend strongly on pH. The static CAs decrease (Fig. 2A) and the swelling increases (Fig. 2B) with increasing pH because of the emergence of negative charges in the PaAA chains. The pH of the wettability transition, pHtr, defined as the pH at the transition halfheight, depends on polyacid hydrophobicity. It increases from 5.8 to 6, 6.4, and 7.8 for the PMAALbL, PEAALbL, PPAALbL and PBAALbL hydrogels, respectively. These transition values roughly correlate with the pKa values of 6.2, 6.5, 6.8 and 7.2 for PMAA, PEAA, PPAA 67 and PBAA, respectively. Compared to previous pHtr data for PaAALbL films in low-salt solutions [23], the transition pHs in salt solutions are smaller by ~0.3 pH units and reflect the screening of chargeecharge interactions by the salt. The PMAALbL hydrogel, whose constituent PMAA chains remain hydrated over the entire pH region, demonstrate the lowest static CAs and the largest swelling among the family of PaAALbL hydrogels. In contrast, the more hydrophobic PEAALbL, PPAALbL and PBAALbL hydrogels, whose constituent chains dehydrate and collapsed at pH values below 5.2, 5.6 and 7, respectively, show less swelling and higher contact angles. In an acidic environment, the most hydrophobic of the PaAALbL films (PBAALbL) contains as little as 8% water. Nanoindentation was used to correlate the PaAALbL film swelling with changes in the surface mechanical properties (Fig. 3). The hydrophilic PMAALbL coating, which exhibits the highest water uptake of the various hydrogel films, exhibits a modulus six orders of magnitude lower than that of Si wafers. Importantly, its modulus decreases by almost 3-fold between pH 3 and 8. A similar decrease in modulus has been reported for more rigid electrostaticallyassembled multilayers with increased hydration [26] and for microgels that swell in response to external stimuli (pH and temperature) [27]. Because of the large changes in the swelling and the stronger electrostatic repulsion at higher pH, the modulus strongly depends on pH in the entire polymer system studied here. Moreover, Eind increases with PaAA hydrophobicity, with the most hydrophobic hydrogel (PBAALbL) having the highest modulus at pH 3 (Eind ~ 338 kPa). The value of Eind was as low as ~72 kPa for the PMAALbL hydrogel at pH 8, which is within the range reported for crosslinked poly(L-lysine)/hyaluronan LbL films (30e150 kPa) [28] but lower than that reported for crosslinked PAH/PAA [29]. The dependence of the elastic modulus on polyacid hydrophobicity reflects the impact of the reduced water content and the increased number of polymerepolymer contacts within the more hydrophobic PaAALbL films. 3.2. Antibacterial activity of PaAAs in solution Prior to exploring the antibacterial activity of PaAA chains immobilized within LbL hydrogels, we studied the interactions of PaAA chains and bacteria in solution. S. epidermidis was chosen for Fig. 2. (A) Static CAs of PaAALbL-coated silicon wafers as a function of pH. (B) pH dependence of the swelling ratios of PaAALbL surface hydrogels, defined as the ratio of wet film to dry film thicknesses, as measured by in situ ellipsometry. All experiments were done in 0.01 M phosphate buffer solutions containing 0.2 M NaCl. The error bars represent the standard deviation of the contact angle data, obtained by measurements at three different spots on each of three identically prepared samples. Fig. 3. Young's moduli measured by nanoindentation, Eind, of PaAALbL coatings deposited on Si wafers and exposed to 0.2 M NaCl PBS solutions at pH 3 and pH 8. Representative force curves for all LbL-coated samples and control Si wafers are shown in Figure S2 in Supporting Information. The error bars represent standard deviations based on five repeated measurements. 68 Y. Lu et al. / Biomaterials 45 (2015) 64e71 Fig. 4. (A) Time evolution of pH during culturing of S. epidermidis (initial concentration 1 106 CFU/mL) in TSB. (B) Bacterial concentration and pH of S. epidermidis TSB culture solutions containing 1.25 mg/mL of various PaAA acids after 24 h of bacteria culture. The shaded color gradient illustrates increased polyacid hydrophobicity. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.) these experiments, because it is frequently implicated in many biomaterials-associated infections. As a result of metabolic activity, these bacteria secret lactic acid during growth, causing acidification of the immediate environment [30]. Fig. 4A shows that growing bacteria, in the absence of any dissolved PaAA, cause a significant drop in solution pH. The pH plateau for t < 10 h is due to the buffering effect of TSB. This is followed by a steep decrease in pH from 7.5 to ~5 over the period of 10 h < t < 20 h (shaded area in Fig. 4A). After 24 h, the solution pH decreases to 4.8 while the concentration of bacteria increases by almost 300%. In contrast, adding PaAA to the culture medium substantially reduces the concentration of bacteria measured after 24 h of culture (Fig. 4B). And, with fewer bacteria, the pH decrease is concomitantly less. These effects increase as the hydrophobicity of the PaAA increases. We hypothesize that this bacteria-induced pH drop drives conformational coil-to-globule transitions in the polyacid chains, reveals hydrophobic moieties within the more hydrophobic PaAAs, and triggers the penetration of PaAA segments through the bacterial wall, ultimately resulting in bacterial death. The data in Fig. 4B corroborate this hypothesis. The bacterial concentration in suspension after 24 h of culture is largest for the PMAA, and there is a systematic decrease in the bacterial concentration when cultured in the presence of PaAAs with increased hydrophobicity. Correspondingly, suspensions cultured with the more hydrophobic PaAAs maintain higher pH values, because these solutions grow a smaller number of media-acidifying bacteria. The strongest inhibition is observed with the most hydrophobic PaAA (PBAA). These results indicate that a new interaction mechanism between the PaAAs and the bacteria is activated at decreased pH values, which leads to the inhibition of bacterial growth. The mechanism is probably similar to that described for interactions between PaAAs and phospholipid membranes, which involves the insertion of hydrophobic segments of PaAA into the membrane, membrane reorganization, and membrane permeabilization [31,32]. Note that these interactions occur despite the negative change of S. epidermidis between pH 5 and pH 7.5 (zeta potential ~ 24 mV, Figure S5). To the best of our knowledge, this is the first study of interactions between bacteria and hydrophobic pH-sensitive anionic polymers, with prior reports exploring the antibacterial activity of cationic polymers [33]. study PaAAs thin films when these polymers are immobilized within LbL coatings. Previous studies indicate that when bioactive polyelectrolyte chains are included within traditional LbL coatings, charge compensation and ionic pairing neutralize the cytotoxicity of polyamines [34]. We hypothesize, however, that segmental 3.3. Antibacterial activity of PaAALbL-coated substrates Fig. 5. (A) pH evolution in S. epidermidis suspensions (initial concentration 1 106 CFU/mL) in TSB above uncoated or coated Si wafers. (B) Fractional bacterial coverage of PaAALbL surfaces at different culture times. The data and standard deviations are calculated from 27 CLSM images of bacteria-covered hydrogels taken from four repeated experiments. Representative CLSM images are shown in Figure S3 and S4. While these experiments indicate that anionic polymers with pH-switchable hydrophobicity exhibit effective antimicrobial properties when dissolved in solution, the main focus here is to Y. Lu et al. / Biomaterials 45 (2015) 64e71 mobility within LbL hydrogels can enable the penetration of hydrophobic polymer segments into the bacterial membrane and therefore introduce antibacterial functionality into an LbL coating. The effect of surface-attached hydrogel thin films on bacterial growth and viability were evaluated by culturing bacteria in suspensions brought in contact with either bare Si wafers (control) or Si wafers coated with PMAALbL, PEAALbL, PPAALbL or PBAALbL hydrogel thin films (Fig. 5A). In all cases, bacterial growth in solution decreases the pH of the TSB medium. Furthermore, the time evolution of solution pH (Fig. 5A) is similar in all suspensions, regardless of the presence or the absence of the PaAALbL coatings on the substrates. This finding indicates that bacteria/substrate interactions do not significantly affect bacterial growth in the surrounding solution. At the same time, the pH changes in the medium induce deswelling and hydrophobicity transitions in the PaAALbL coatings (see Fig. 2). These transitions have a strong impact on the bacteria/substrate interactions. Fig. 5B illustrates the time dependence of the bacterial surface coverage for the various types of PaAALbL coatings. Interestingly, the bacterial colonization of the more hydrophilic and softer PMAALbL coatings is indistinguishable from that of the coating-free Si control surface. In contrast, the more hydrophobic and stiffer PaAALbL coatings resist bacterial colonization. In fact, the most hydrophobic coating (PBAALbL) remains largely bacteria-free until 12e14 h into the 24 h culture period. While the literature reports conflicting observations on the effect of substrate stiffness on bacterial coverage [28,35], our results are in agreement with work by N. Saha et al. [28] who also reported more rapid bacterial growth on softer LbL films. A steep onset in the bacterial colonization of PEAALbL, PPAALbL and PBAALbL hydrogels occurs for times longer than 12 h when the pH decreases below ~6.3. This relatively abrupt increase in bacterial surface coverage reflects an important property of our hydrophobic hydrogel coatings: they reveal their hydrophobic moieties in an acidic environment, which, in turn, dramatically alters their interactions with bacteria. While bacterial colonization increases at the onset of the hydrophilicehydrophobic transition (Fig. 5), live/ dead staining experiments show that a significant fraction of these 69 bacteria are dead (Fig. 6). The fraction of dead bacteria increases with increasing hydrogel hydrophobicity and becomes as high as 0.65 for the case of the PBAALbL coatings (Fig. 6E). In contrast, live bacteria almost exclusively populate surfaces of hydrophilic PMAALbL films (Fig. 6A) and control Si wafers (data not shown). Moreover, none of the coatings show a bacteria-killing ability for bacteria culture times shorter than 12 h, when the pH does not drop below 7.0 and the coatings remain highly hydrated. The bactericidal efficiency of PaAALbL coatings turns on only as pH decreases and reveals the coating hydrophobicity. 3.3.1. Cytocompatibility of PaAALbL-coated substrates A critical objective in the development of bacteria-resistant tissue-contacting implants is to endow the coatings with specific activity against bacteria without compromising their ability to support the adhesion and proliferation of tissue cells [36]. We thus used human fetal osteoblasts to assess the cytocompatibility of the various thin-film hydrogel coatings. Fig. 7 shows the results of MTS proliferation and viability tests. Osteoblast proliferation occurs on all of the PaAALbL coatings, but proliferation on the hydrophilic and more swollen PMAALbL surfaces is ~50% higher than that on the more hydrophobic PaAALbL coatings. The three hydrogels with pHswitchable hydrophobicity (PEAALbL, PBAALbL and PBAALbL) demonstrate similar osteoblast growth properties. In particular, while it has been established that softer, more hydrated films are typically more cytophobic [37], here, in agreement with other earlier report [38], we find that osteoblast proliferation is favored by more hydrophilic surfaces. In addition, Fig. 7A shows HFOB 1.19 cells, visualized by livedead staining, on silicon wafers and on PaAALbL hydrogels. All proliferating cells appear as green (Fig. 7A), indicating that the hydrogels do not induce cellular death. Overall, the data in Fig. 7A and B illustrate that all types of hydrogels studied here are non-toxic to osteoblasts. This finding might not be surprising in the case of PMAALbL, PEAALbL and PPAALbL hydrogels, since these polyacids do not disrupt lipid bilayer or cellular membranes at pH 7.4 [39]. However, the compatibility of Fig. 6. (AeD) Comparison of the total bacteria (black symbols) and dead bacteria (red symbols, shaded areas) coverage for various PaAALbL coatings. Live/dead bacteria staining images are shown in Figure S3 and S4. (E) Fraction of dead bacteria as a function of culture time. The error bars represent the standard deviations of bacterial coverage data, analyzed by ImageJ software on 9 live/dead stained, bacterial CLSM images. Description of how to calculate bacterial coverage data can be found in the Materials and Methods section. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.) 70 Y. Lu et al. / Biomaterials 45 (2015) 64e71 Fig. 7. (A) CLSM images of HFOB 1.19 cell attachment and proliferation on Si wafers and PaAALbL hydrogels at day 1 and day 7. The scale bar is 100 mm. The cells were live/dead stained prior to imaging, but showed exclusively live (green) cells, with no observable dead (red) cells. (B) Mammalian cell metabolic activity studied by the MTS assay of osteoblast proliferation on control Si wafers and wafers covered by PaAALbL hydrogels. Statistically significant differences from the control Si samples and PMAALbL-coated substrates were determined using a student's t-test (p < 0.05) are marked by stars. The magnitude and error bars represent the average and standard deviation, respectively, from 9 different experiments. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.) the PBAALbL hydrogels with human osteoblasts is an important new finding, since free PBAA homopolymer in solution causes red blood cell lysis even at pH 7.4 [39]. While the diminished cytotoxicity of polycations has been shown to be a result of strong binding between polycations and polyanions within an LbL film [34], here we have demonstrated that cytotoxicity of polymer chains can be hampered by their immobilization within a surface network. Obviously, the immobilization of polyacid chains within hydrogels significantly hinders the availability of hydrophobic functional groups for membrane interactions at pH 7.4, rendering these hydrogel coatings non-toxic to osteoblasts, which do not acidify the immediate environment. Being included within a hydrogel nevertheless preserves sufficient segmental mobility that enables these polymers to become bactericidal in acidic environments when the coatings turn hydrophobic and become more dehydrated. 4. Conclusions Here, we have developed the self-defensive concept for bacteria-responsive hydrogel thin films that do not incorporate antimicrobial agents for triggered release. These films instead act via a pH-induced mechanism of antibacterial activity inherent within the thin film itself. We studied a series of four LbL hydrogel thin films, which enable the systematic variation of their pHdependent hydrophobicity. The pH-induced hydrophilic-to-hydrophobic transitions of these coatings appear to be the central enabler underlying their bactericidal properties. The results of these experiments suggest that the antibacterial mechanism involves the insertion of hydrophobic segments from the polymer networks into bacterial walls, which can lead to wall damage and bacterial death. However, regardless of their hydrophobicity, the Y. Lu et al. / Biomaterials 45 (2015) 64e71 coatings are noncytotoxic to osteoblasts. We suggest that the design of antibiotic-free bacteria-triggered films presents a promising path towards the development of bacteria-responsive surface coatings that may prevent biofilm growth on tissue-contacting biomedical devices. 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