Tailoring Polyelectrolyte Architecture To Promote Cell Growth and

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Tailoring polyelectrolyte architecture to promote cell growth and inhibit bacteria adhesion Shanshan Guo, Min Yi Kwek, Zi Qian Toh, Dicky Pranantyo, En-Tang Kang, Xian Jun Loh, Xiaoying Zhu, Dominik Ja#czewski, and Koon-Gee Neoh ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b00666 • Publication Date (Web): 13 Feb 2018 Downloaded from http://pubs.acs.org on February 15, 2018

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ACS Applied Materials & Interfaces

Tailoring polyelectrolyte architecture to promote cell growth and inhibit bacteria adhesion Shanshan Guo†, Min Yi Kwek‡, Zi Qian Toh‡, Dicky Pranantyo‡, En-Tang Kang‡, Xian Jun Loh ǁ,⊥, #, Xiaoying Zhu,§, Dominik Jańczewski*,∇, Koon Gee Neoh*,†,‡ †

NUS Graduate School for Integrative Science and Engineering, National University of Singapore, Kent Ridge, Singapore 117576 ‡

Department of Chemical and Biomolecular Engineering, National University of Singapore, 4 Engineering Drive 4, Singapore 119260. E-mail: [email protected]; Tel: +65 6516 2176; Fax: +65 6779 1936 §

ǁ

Department of Environmental Science, Zhejiang University, Hangzhou, China, 310058

Institute of Materials Research and Engineering, A*STAR (Agency for Science, Technology and Research, 2 Fusionopolis Way, Singapore 138634

⊥Department

of Materials Science and Engineering, National University of Singapore, 9 Engineering Drive 1, Singapore 117576 #

Singapore Eye Research Institute, 11 Third Hospital Avenue, Singapore 168751

∇Laboratory

of Technological Processes, Faculty of Chemistry, Warsaw University of Technology,

Noakowskiego 3, 00-664, Warsaw, Poland. E-mail: [email protected]; Tel: +48 22 234 5583; Fax: +48 22 234 5504 KEYWORDS: layer-by-layer assembly, polyelectrolyte multilayer stiffness, surface charge, surface wettability, antimicrobial, cell adhesion and proliferation

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Abstract An important challenge facing the application of implanted biomaterials for tissue engineering is the need to facilitate desirable tissue interactions with the implant while simultaneously inhibiting bacterial colonization, which can lead to implant-associated infection. In this study, we explore the relevance of physical parameters of polyelectrolyte multilayers, surface charge, wettability and stiffness, in tissue cell/surface and bacteria/surface interactions, and investigate the tuning of the multilayer architecture to differentially control such interactions. Polyions with different side chain chemical structures were paired with polyethylenimine to assemble multilayers with parallel control over surface charge and wettability under controlled conditions. The multilayers can be successfully crosslinked to yield stiffer (the apparent Young’s modulus was increased more than 3 times its original value) and more stable films while maintaining parallel control over surface charge and wettability. The initial adhesion and proliferation of 3T3 fibroblast cells were found to be strongly affected by surface charge and wettability on the non-crosslinked multilayers. On the other hand, these cells adhered and proliferated in a similar manner on the crosslinked multilayers (apparent Young’s modulus ~2 MPa) regardless of surface charge and wettability. In contrast, Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli) adhesion was primarily controlled by surface charge and wettability on both crosslinked and non-crosslinked multilayers. In both cases, negative charge and hydrophilicity inhibited their adhesion. Thus, a surface coating with a relatively high degree of stiffness from covalent crosslinking coupled with negative surface charge and high wettability can serve as an efficient strategy to enhance host cell growth while resisting bacterial colonization. 1. Introduction The main barriers to the long-term success of implanted biomaterials for tissue engineering are insufficient tissue cell adhesion and the possibility of bacterial colonization on the surfaces, which can lead to biofilm formation and implant-associated infection.1,2 Thus, it is desirable to design implant surfaces that simultaneously enhance host cell adhesion and suppress bacterial adhesion.3,4 The interactions of solid surfaces with bacteria and tissue cells are greatly influenced by the physical properties of the surface.5-8 These properties include surface charge state,9-12 hydrophilicity8,13 and stiffness.14,15 Therefore, understanding how these physical 2


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parameters can be manipulated to control the adhesion of mammalian and bacterial cells on surfaces will be of great significance in the design of biomaterial surfaces for practical applications. Control over surface properties can be achieved by an appropriately designed coating. A versatile approach for producing thin film coatings with nanoscale control over surface properties and architecture is the layer-bylayer (LbL) assembly.16-19 Interaction of mammalian cells with LbL films has been probed extensively, and as a result, many physical properties of LbL films that control cell behavior have been established. Foremost among these is film stiffness, which has been modulated by covalent crosslinking20-22 or assembly pH.16,23 Fibroblast,16,20 preosteoblast,21 skeletal myoblast22 and endothelial cells23 have been shown to increase their adherence, spreading and proliferation on stiffer LbL films. The processing of mechanical stimuli by mammalian cells is partly mediated by integrins that link the cytoskeleton to the physical environment.24,25 However, when the mechanical properties of the films are changed, variation of other surface properties may also occur. For example, increasing the elastic modulus of multilayers of chondroitin sulfate A and poly(Llysine) after genipin crosslinking led to a significant increase in arithmetic mean roughness from ~20 to ~30 nm for six bilayer films and ~14 to ~160 nm for 12 bilayer films.21 Assembly pH has been used to modulate the elastic modulus of multilayers of poly(acrylic acid) and poly(allylamine hydrochloride), but it is not reported whether other surface properties (surface roughness, wettability and charge) were affected.23 Similarly, the Young’s modulus of multilayers of poly(L-lysine) and thiol-modified hyaluronan was increased by oxidation of thiol groups,20 but possible differences in surface charge were not quantified. Thus, it is difficult to elucidate the contributions of the different parameters in controlling cell behavior on these multilayers. Myoblast cell behavior on multilayers assembled from covalent carbodiimide crosslinking of poly(L-lysine) and hyaluronan without significant changes in film roughness and wettability has been reported.22,26 However, these studies investigated only hydrophilic films (with water contact angle ~25° before crosslinking and < 10° after crosslinking) and did not show how substrate stiffness and competing surface features, surface charge and wettability, act together to influence cell adhesion and growth. Thus, a systematic consideration of the collective effects of these substrate characteristics remains an important topic. 3



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Furthermore, it is still unknown whether the range of substrate stiffness that modulates the adhesion of mammalian cells, would influence bacterial adhesion. Previous studies on multilayers of poly(acrylic acid) and poly(allylamine hydrochloride) considered a wide range of stiffness changes (with the Young’s modulus varied from ~1 to ~100 MPa) by modulating assembly pH and have shown that both bacterial and human microvascular endothelial cells increased their adhesion on stiffer films.23,27 Thus, cell selectivity cannot be achieved using this approach. In the current study, we varied the Young’s modulus of LbL films within a narrow range from ~0.5 to ~2 MPa to investigate the effects on mammalian and bacterial cells. Importantly, the variation in stiffness was coupled with a systematic modulation of surface charge and wettability to investigate how these competing substrate features act together to influence adhesion of mammalian and bacterial cells. Particular attention was also paid to minimize variation in thickness and morphological features of the films. The native LbL films were selected from a library of films investigated in our earlier study.28 These LbL films were built from branched polyethylenimine (PEI) and custom-synthesized derivatives of poly(isobutylene-alt-maleic anhydride) (PIAMAn) bearing either poly(ethylene glycol)- or alkyl-carboxylic side chains. Film crosslinking was carried out using carbodiimide 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC) in the presence of N-hydroxysuccinimide (NHS).22,29,30 These multilayers were then used as model surfaces to assess the effect of manipulating the state and range of the multilayer properties on surface interactions with 3T3 fibroblast cells, Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli). 2. Materials and methods 2.1

Materials

PIAMAn (Mw 60 kDa), PEI (Mw 25 kDa, branched), 6-aminocaproic acid (Mw 131.17 Da), (3-aminopropyl)trimethoxysilane (APTMS, 97%), EDC, NHS, folate-free Dulbecco’s modified Eagle’s medium (DMEM), fetal bovine serum, L-glutamine, penicillin and tryptic soy broth were purchased from Sigma-Aldrich. Amine-PEGcarboxylic acid (NH2-PEG-COOH, Mw 2000) was purchased from Laysan Bio Inc. Toluene and ethanol were purchased from Tedia. Ultrapure water (18.2 MΩ cm) produced by a Millipore Milli-Q integral water 4


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purification system was used to prepare aqueous solutions. Silicon wafers were purchased from Latech Scientific Supply, Pte. Ltd. S. aureus (ATCC 25923), E. coli (ATCC 25922) and 3T3 mouse fibroblast cells (ATCC CRL-1658) were purchased from American Type Culture Collection. 2.2

LbL film fabrication

The polyanions PIAMA-C5 and PIAMA-PEG (Figure 1) were synthesized according to the previously published protocol.28 Briefly, 6-aminocaproic acid (0.25 g, 1.9 mmol) or NH2-PEG-COOH (0.75 g, 0.38 mmol) was dissolved in 5 mL of dimethyl sulfoxide and mixed with PIAMAn (0.5 g, 3.8 mmol of monomeric repeats) for synthesis of PIAMA-C5 and PIAMA-PEG, respectively. The reaction proceeded for 24 h at 40 °C for 6aminocaproic acid and at room temperature for NH2-PEG-COOH. The polymer solution after dissolving in 50 mL of NaOH aqueous solution (10 g L−1) was dialyzed using dialysis membrane (FisherBrand, 6000-8000 MWCO) against ultrapure water for 3 days. For LbL assembly, silicon substrates with dimensions of 1 cm × 2 cm were first cleaned by sonication in ethanol and water, respectively, for 10 min. After drying in a nitrogen stream, they were treated with oxygen plasma at 200 W (Model ATTO, Diener Electronic) for 120 s. APTMS at a concentration of 10 mM in toluene was deposited on the substrates for 3 h and served as a positively charged precursor layer. Subsequently, the films were rinsed successively with toluene and water, and blown dry with a stream of nitrogen. LbL assembly was conducted by alternately immersing the substrates in polycation and polyanion solutions (1.0 mg mL−1) for 5 min in a cyclic manner up to 6.5 or 7 bilayers. 0.5 layer refers to one polyanionic layer, and one bilayer indicates one polycationic layer on top of one polyanionic layer. The polycation was PEI, and the polyanion was either PIAMA-PEG or PIAMA-C5. The pH of the polyelectrolyte solutions was adjusted by adding either 0.1 M HCl or 0.1 M NaOH aqueous solution. The films were rinsed with ultrapure water followed by drying with a nitrogen stream between these immersions. After the desired number of layers was deposited, the films were dried with nitrogen and further dried at room temperature for 24 h before characterization experiments. The LbL assembly was conducted at room temperature (25 °C).

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2.3

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Covalent crosslinking of the films

Covalent crosslinking of the LbL films was carried out via reaction between activated carboxylic sites and primary amine groups. An EDC/NHS mixture (EDC at 5 or 25 mM, EDC:NHS molar ratio was kept at 1:1.6) was freshly prepared in DI water at pH 5.5. The LbL-coated substrates were submerged into the EDC/NHS solution for 12 h at 25 °C, and then washed with DI water followed by soaking in DI water for 1 h and drying with a stream of nitrogen. 2.4

LbL film characterization

Surface chemical composition was determined by X-ray photoelectron spectroscopy (XPS) (Thermo Fisher Scientific Theta Probe, monochromatic Al Kα X-ray source) at a takeoff angle of 90°. Atomic concentrations were determined from high-resolution scans at a resolution of 0.05 eV. Attenuated total reflection-Fourier transform infrared spectroscopy (ATR-FTIR) measurement was carried out using a FTIR spectrometer (Bruker, Vertex 80v) at a resolution of 4 cm-1 for each sample surface. All ATR-FTIR measurements were conducted under vacuum conditions. The zeta (ζ) potential of the surfaces was determined by an electrokinetic analyzer (SurPass, Anton Paar). Four measurements were conducted for each run. Contact angles were measured with a goniometer (250-F1 from Rame-Hart Instrument Co.) at room temperature via the static sessile drop method. The average value of four measurements on different locations of three samples was reported. Atomic force microscopy (AFM) experiments were conducted with a NanoWizard 3 NanoOptics atomic force microscope (JPK Instruments AG). Surface morphology in the dried state was measured at three 5 µm × 5 µm spots using silicon probes in tapping mode (k~40 N m−1, Tap 300AL-G, Budget sensors). The root-mean-square roughness (Rq) was determined by the installed software (JPK Data Processing, 4.3.25). 2.5

AFM force spectroscopy

The apparent Young’s modulus (Eapp) of the films was calculated from force-indentation curves obtained using AFM. Force-indentation curves were collected using colloidal probes (with SiO2 sphere, diameter 2.5 µm, Novascan Technologies, Inc.) with a spring constant ranging between 0.07 and 0.12 N m−1 as determined by the 6


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thermal noise method. Typical measurements were conducted in a force volume imaging mode with a z-ramp size of 1 µm, a scan rate of 1.0 Hz and a scanning area of 8 µm × 8 µm with 64 measurements of the deflection as a function of the piezo z-position. All measurements were conducted in phosphate buffered saline (PBS, pH 7.4). The films were hydrated in PBS for a minimum of 1 h before AFM measurement. Before each experiment, the cantilevers were immersed in medium for 15 min to allow thermal equilibration. ( )

The Eapp values of the films were estimated using the Hertz model for a spherical indenter: E = / / where F is the applied force, ν is the Poisson's ratio, δ is the indentation depth and R is the radius of the colloidal probe. Poisson’s ratio was assumed to be 0.5.31,32 In model fitting, Eapp and the contact point were used as the fit parameters using built-in JPK software (see example in Section 3 of the Supporting Information). For thin films on a hard substrate such as silicon wafer, Eapp is not a constant value but varies according to the indentation depth by the AFM tip. Thus, to minimize substrate stiffness effects, low deformation values were used for the calculation. 2.6

Quartz crystal microbalance with dissipation (QCM-D)

QCM experiments to measure protein adsorption were conducted with a Qsense E4 multichannel instrument. Quartz crystal disks coated with silicon dioxide (fundamental frequency 4.95 MHz) were used as sensor chips. The sensor surface was first coated with LbL films according to the procedure described in Section 2.2 and 2.3. Prior to protein adsorption measurement, PBS solution (pH 7.4) was flowed into the QCM cell until a stable frequency (∆f) and dissipation (∆D) baseline reading was reached. The protein solution (100 µg ml−1) was passed through the measuring chamber until surface saturation, followed by buffer rinse to remove loosely attached molecules. The Sauerbrey equation, ∆ = −

∆

, was used to calculate the mass of adsorbed

proteins (∆m) for rigid adsorbed layers (∆D/∆f

Tailoring Polyelectrolyte Architecture To Promote Cell Growth and

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