Direct Loading and Tunable Release of Antibiotics from

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Direct loading and tunable release of antibiotics from polyelectrolyte multilayers to reduce bacterial adhesion and biofilm formation Bailiang Wang, Tingwei Jin, Qingwen Xu, Huihua Liu, Zi Ye, and Hao Chen Bioconjugate Chem., Just Accepted Manuscript • DOI: 10.1021/acs.bioconjchem.6b00118 • Publication Date (Web): 22 Apr 2016 Downloaded from http://pubs.acs.org on April 26, 2016

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Bioconjugate Chemistry

Direct loading and tunable release of antibiotics from polyelectrolyte multilayers to reduce bacterial adhesion and biofilm formation Bailiang Wanga, c*, Tingwei Jinb, Qingwen Xua, Huihua Liub, Zi Yea, and Hao Chena, c*

a

School of Ophthalmology & Optometry, Eye Hospital, Wenzhou Medical University,

Wenzhou, 325027, China b

Department of Basic Teaching, City college of Wenzhou University, Wenzhou, 325027,

China c

Wenzhou Institute of Biomaterials and Engineering, Chinese Academy of Sciences,

Wenzhou, 32500, China * Corresponding author. Fax: +86 577 88067962. E-mail: [email protected] (B. L. Wang), [email protected] (H. Chen).

Abstract: Bacteria adhesion on the surface of biomaterials and following biofilms formation are important problems during the biomedical applications. The charged antibiotics with small molar mass can hardly alternately deposit with polymers into multilayer films to load the drug. Herein, the (poly(acrylic acid)- gentamicin / poly(ethyleneimine) )n ((PAA-GS/PEI)n) multilayer film was designed and constructed via layer-by-layer self-assembly method. Low molar mass GS cations were first combined with polyanion PAA and self-assembled with PEI to form multilayer films showing exponentially growth behavior. The GS dosage could be adjusted through changing the layer number of films. Furthermore, thermal cross-linking method was used to control the release rate of GS in PBS buffer. Owing to the diffusion of GS, a zone of inhibition about 7.0 mm showed the efficient disinfection activity of the multilayer film. Also it could be seen from the biofilm inhibition assay that the multilayer film effectively inhibited bacteria adhesion and biofilm formation. As the drug loading dosage was 160 µg/cm2, the multilayer films showed very low cytotoxicity against human lens epithelial cells. The present work provides an easy way to load GS into multilayer films which can be applied to surface modification of implants and biomedical devices. Keywords: Layer-by-layer self-assembly, gentamicin loading, drug delivery, 1

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antibacterial, thermal crosslinking Introduction Bacterial infection caused biofilm formation is a major cause of implant failure in the applications. The adhesion of bacteria on the implants and devices accounts for more than fifty percent of all nosocomial infection in the world (1, 2). Only in the United States, biomaterials related infections caused the death of more than 100,000 people per year(3). Bacteria adhesion on the implants surface is key and the first step which leads to the formation of bacterial colonies and the biofilm(4, 5). As a result, to inhibit bacteria adhesion is critical important in reducing the incidence of biomedical infection. There is a complicated process in bacteria adhesion which basically includes two stages(6): (I) interactions that are initial, reversible and rapid; (II) interactions that are slow, irreversible, nonspecific and specific. After the adhesion of bacteria on the surface, the bacterial adhesion changed from reversible to irreversible. In the next step, the adhered bacterial secrete extracellular matrices, proliferate fast and finally grow into biofilm(7). Bacteria hide in the biofilm show much lower susceptibility to antibiotics and it is extremely difficult to remove biofilm from the implants surface(8, 9). As a result, many kinds of antibacterial surfaces have been designed and developed to reduce the adhesion of bacterial and consequently reduce the biofilm formation(10-12). Based on their operating mechanisms, there are mainly three groups of the antibacterial surfaces (6): bacterial anti-adhesive surfaces, bactericidal surfaces and surfaces leaching biocide agents. Although great effects have been paid to develop the antibacterial surfaces, each method has inherent disadvantages and advantages(13). For example, although bacteria resistance surfaces can reduce or prevent the initial adhesion of bacteria, this kind of surface is lack of bactericidal property and long acting stability of the coating owing to the detachment and degradation (14, 15). In the other way, bactericidal surfaces showing efficient sterilization function, always process high cytotoxicity against mammalian cells. Furthermore, the inflammation or immune responses will be triggered by the remaining dead bacteria on the implants (16). Moreover, although biocide agent leaching surfaces show efficiently antibacterial function against both adhered and planktonic bacteria in a certain concentration range, biocide agent release from the matrix is not easy to control(17, 18). In the present work, we aim to develop an antimicrobial coating that can both resist bacteria adhesion on the surface and be able to local delivery of antibiotics in a controlled manner. Although antibiotic resistant bacteria may potentially develop through using of antibiotics, antimicrobial surfaces releasing antibiotic is still the most effective 2


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approach to prevent or treat biomedical infections(19-21). As the sustained release of antibiotics to provide a high local concentration, bacteria can be killed before biofilm formation(22, 23). The amounts of biocide agents in blood or body fluid can also be effectively controlled through drug delivery systems and further to improve the biocompatibility. In controlled delivery systems, the release of the drug can be tuned into a sustained manner through embedding in the matrices(19, 24, 25). Furthermore, if the matrixes are sensitive to specific stimuli, such as electric and magnetic fields, pH, light, temperature and biological ions systems, stimuli sensitive drug delivery system can be developed(26, 27). To construct new delivery systems in an efficient and safe manner, great concerns have been given to carrier materials to controlled release of drugs. It is recently found that microcapsules and thin films constructed via layer-by-layer (LBL) deposition have been used as drug delivery systems (25, 28, 29). Many kinds of multilayer films have been developed as matrixes to load and triggered release of antibacterial agents. The LBL self-assembly method was developed by Decher and coworkers(30), which was based on the sequential deposition of oppositely charged polymers. The most distinctive feature of multilayer film is multi-functionality and responsiveness to many kinds of stimuli such as ionic strength, enzymes, pH, temperature and electrochemical (31-33). Moreover, the thickness of multilayer film can be tuned by simply changing the layer numbers to precisely control the loading dosage of drug in the film. For most of the drugs, it is not easy to load and controlled release of small molecules into delivery systems due to the weak binding to the matrix(34, 35). As for the antibiotics, the binding problem is particularly serious which would result in the bacterial antibiotic resistance. Importantly, through selecting the functional groups type which serves as adsorption sites for drugs, the release profile can be preprogrammed at the step of matrix prepartion(36). As a well-known aminoglycoside antibiotic, gentamicin sulfate (GS) has broad spectrum of antibacterial activity against some fungus, certain gram-positive and most of the gram-negative bacteria (37, 38). Bactericidal activity of aminoglycosides is concentration dependent and has post-antibiotic effect which is different from other antibiotics impeding protein synthesis of bacteria. However, it is not easy work to directly load small, uncharged, and hydrophobic antibiotics (about 40% of FDA approved drugs) into multilayer thin films due to the nonpolymeric character(39). As a result, how to construct multilayer films that can directly load charged small molecule antibiotics is of great importance(34, 40, 41). Hammond et al.(34) synthesized poly(β-amino esters) (Poly X with different chain length, X=1, 2 and 6A) with positive charges and constructed into multilayer films [Poly 1/Anion/GS/Anion]n tetralayer architecture through alternately self-assembly of anion with GS. The drug could release in a controlled manner from 3



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the multilayer films through hydrolysis of Poly X. The mechanism of the multilayer films hydrolytic degradation made the continuous release of GS and there was no need of enzymic or cellular interaction. In the present work, we aimed to load high capacity of GS into the LBL multilayer film through self-assembly method. The drug loaded films should not only provide high local concentration of antibiotics to resist bacterial infection, prevent biofilm formation, but also low degree of cytotoxicity against mammalian cells. First of all, GS was combined with PAA into a complex solution and then directly incorporated into the (poly(acrylic acid)- GS / poly(ethyleneimine) )n ((PAA-GS/PEI)n) multilayer films via self-assembly method. Specifically, we could tailor the rate of film deconstruction by introducing cross-linking between carboxyl groups PAA and amino groups of PEI via heat treatment to control the release rate of GS. This approach represents a convenient and effective strategy to incorporate charged small molecules into LBL films for antibacterial applications. Results and discussion Construction of the (PAA-GS/PEI)n multilayer films

Scheme 1. Schematic representation of construction, cross-Linking, degradation, and antibacterial properties of the (PAA-GS/PEI)n multilayer films.

A major challenge in the development of advanced drug formulations is the elaboration of delivery systems capable of providing sustained release of bioactive materials. Sustained and controlled-release mechanisms offer greater effectiveness, lower toxicity, and improved patient convenience over conventional formulations. We focus our attentions on the design of multilayered assemblies that sustained release incorporated macromolecular materials in aqueous environments. As shown in scheme 1, PAA-GS complex at pH 2.9 and PEI under pH 9.0 were alternately deposited onto substrates to form multilayer films. A precursor layer of PEI was adsorbed onto all substrates before multilayer assembly (42). This kind of substrate 4


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treatment is simple and independent of substrate species, so it has been already widely used in coating technology. To control the release of GS, the multilayer films were cross-linked by EDC/NHSS or thermal treatment. The obtained LBL multilayer films loaded the small antibiotic molecules with cations, which combined with PAA through electrostatic forces. Then the loaded GS released from the multilayer films when exposed to an aqueous physiological environment (0.1 M PBS). The presence of electrostatic forces between PEI and PAA contributed to the stability of the films. Thermal crosslinking of the multilayer films was implemented to delay the film degradation and GS release. The rapid release of antibiotics could be used in the field of short-term implants such as catheters or as wound dressing. The films that sustained release of GS could be used to modify medium and long term implants such as stents, intraocular lens and orthopedics et al.

Fig. 1. Ellipsometry measurement of the (PAA-GS/PEI)10 multilayer film.

Fig. 2. SEM images of (a) thickness of the (PAA-GS/PEI)10 multilayer films and

topographical

features of the (b) (PAA-GS/PEI)7 and (c) (PAA-GS/PEI)10 multilayer films.

The thickness of the (PAA-GS/PEI)n multilayer film increased exponentially with 5



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1210.61 ±320.23 nm for 10 bilayers number with the electrostatic adsorption as the main driving forces (43). In our previous sturdy, a superhydrophobic/hydrophilic asymmetric free-standing film was created by LBL assembly of PEI-Ag+/PAA, which was very thick (19.5 ±1.6 µm made of 20 layer pairs) after peeling off from the Teflon substrate(44). As shown in Fig 1 and Fig. 2a, the ellipsometry data and cross-section measurements from SEM revealed that the film thickness increased exponentially with the increase of deposition layers because of the diffusion “in” and “out” of PEI chains. The “in” and “out” diffusion of PEI with pH change not only resulted in the exponential growth but also in a change of the morphology of the multilayer. Indeed, surface morphology changed from very rough to relatively flat (Fig 2b). Loading of GS into (PAA-GS/PEI)n multilayer films

Fig. 3. Changes of (a) film thickness and (b) cumulative amount of GS loaded into the (PAA-GS/PEI)10 multilayer film as the increase of bilayer numbers.

It has been challenging to directly incorporate small, uncharged, and hydrophobic therapeutics into multilayer thin films due to the nonpolymeric character. Hammond et al.(29, 34, 45) has constructed LBL films with poly(β-amino esters) with controlled erosion nature which could be used to tune the drug release rate from the films. These films degraded through hydrolytic degradation that enabled the continuous release of drug. However, the loading capacity was not high (0.123 mg/cm2 for [(Poly 1/ HA)1 (GS/HA)1]150) and release rate was fast as the film deconstructed in a few hours (34). Therefore, design and preparation of LBL multilayer films to enhance the loading capacity and release the drug in a controlled manner is of great importance. In the present work, as shown in Fig. 3, the GS loading dosage was 0.160±0.056 mg/cm2 for 10 bilayers of (PAA-GS/PEI)n multilayer film and it steadily increased into 1.461±0.346 mg/cm2 as the increasing to 30 bilayers and could be precisely tuned. At the same time, the film thickness grew from 1.21±0.32 µm to 15.23±4.12 µm. However, 1.461±0.346 mg/cm2 was by no means an upper limit, as depositions of 6


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additional layers were expected to result in progressively higher dosages. Comparing with the reported antibiotic loading multilayer films, the capacity of (PAA-GS/PEI)n multilayer film was a very high loading dosage (46).

Kinetics of GS release from (PAA-GS/PEI)n multilayer films The films were immersed in 0.1 M PBS, an ionic buffer that closely modelled conditions of the human body fluid without the biological components, maintained at 37 °C in a water bath. These release curves were normalized to the total released dosage to allow for release rate comparison. Coupled with the loading behaviour of GS, studies were assessments on the kinetics of drug release from these untreated and crosslinked films by EDC/NHSS or thermal treatment. As indicated in Fig. 4, while the untreated films and cross-linked films by EDC/NHSS completed 90.8 % and 99.2 % of the GS release within 24 h, cross-linked films by thermal treatment displayed more prolonged release for 47.2 %. And even as the release of GS from the thermally crosslinked films continued 168 h, it only had released 83.4% of the total amount.

Fig. 4. Normalized cumulative GS release from the untreated, EDC/NHSS and thermally crosslinked (PAA-GS/PEI)10 multilayer film as the release of GS.

GS is expected to diffuse within the film in and out, since it is a small trisaccharide molecule with a molecular weight of 477 Da. The untreated and EDC/NHSS crosslinked films showed fast release of GS owing to the rapid diffusion of GS and deconstruction of the films. It was obvious that the controlled release effect of the two crosslinking methods was significantly different. It is well known that through EDC/NHSS chemical crosslinking, amide bonds formed between carboxylic groups of PAA and amine groups of PEI (47). However, we found that the films became brittle after chemical crosslinking. As it was found in related studies by Coimbra et al. (48) and Manabe et al.(49), this crosslinking method produced porous 7



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structures in the multilayer films. While thermal treatment reduced the permeability of the films and sustained the release of GS. It was reported by Shen et al. (44) and Hammond et al. (50) that thermal crosslinking greatly enhanced the stability of the films and reduced the degradation rate of the film. Furthermore, the formed amide bonds formed between PAA and PEI was difficult to hydrolyze in water. As a comparison, in was indicated by Hammond et al. that [(Poly X/HA)1(GS/HA)1]n system released almost 95 % of its gentamicin in modified simulated body fluid in several hours due to the rapid hydrolysis of Poly X (34).

Fig. 5. Inhibition zones of (a) (PAA/PEI)10 multilayer films and (b) thermally cross-linked (PAA-GS/PEI)10 multilayer film against S. aureus.

Antibacterial property of the (PAA-GS/PEI)n multilayer films Zone of inhibition (ZOI) assays

8


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Fig. 6. Changes of the inhibition zone sizes of (a) (PAA/PEI)10 multilayer films and (b) thermally cross-linked (PAA-GS/PEI)10 multilayer films against S. aureus.

Because more than 85% of S. aureus species are susceptible to GS, we selected S. aureus as the test organism for our study(51). The bactericidal activity of the GS loaded films was examined by a Kirby-Bauer assay against Gram-positive S. aureus, the most common pathogenic bacteria in implant-related infections. As the films were incubated on the bacterial agar plates, the drug diffused out of the films and inhibited the growth of the bacteria, leaving a circular area free of bacteria. The diameter of a ZOI provided a quantitative measurement of the amount of active GS released and diffused into the agar. Fig. 5 presented a sampling of ZOI data for various film-substrates. The ZOI of the thermal crosslinked film loading GS was around 7.1± 0.9 mm (Fig. 5b), whereas the multilayer film without GS loading did not show any bacterial growth inhibition (Fig. 5a). The long-acting bactericidal activity of the untreated and thermal crosslinked GS loaded multilayer films was also tested after being immersed in PBS (as shown in Fig.6). The inhibition zone size was 7.8±1.3 mm for the untreated multilayer film before being immersed in PBS. However, it fast decreased into 1.3±0.6 mm and 0.2

±0.1 mm after being immersed for 1 d and 2 d, respectively. As for the thermally treated multilayer films, the sizes of inhibition zone gradually decreased to 6.5±1.1 mm and 6.0±0.9 mm after being immersed for 1 d and 2 d, respectively. Even after being immersed for 14 d, the ZOI around the film also maintained 0.6±0.4 mm. So it confirmed that the thermal treatment prolonged the release of GS and retained the antibacterial property of the films for a longer time than untreated films.

Anti-adhesion and biofilm inhibition assay. The anti-adhesion properties and bactericidal efficiency of the thermally crosslinked GS loaded multilayer films were conducted qualitatively by the bacterial LIVE/DEAD stain method. As shown in Fig. 7, the distribution of viable (green fluorescence) and dead (red fluorescence) bacteria on the pristine and films modified PDMS substrates were observed by staining with a combination dye of SYTO 9 and PI. At 4 h, it could be observed that there were many distinguishable bacterial cells with green fluorescence, in small clusters or individually, on the pristine PDMS and (PAA/PEI)10 multilayer films(Fig. 7a and 7b). Meanwhile, the number of the S. aureus adhered on the GS loaded multilayer films was much less than that on pristine PDMS and most of the bacteria had been killed (stained red) (Fig. 7c). It could be due to the rapid release of GS into the bacteria solution and fast killing of the bacteria. It has been proved that hydrophilic multilayer films took up large quantities of free water, 9



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which built up the stable defense layer to resist bacterial adhesion (44, 52). Upon prolonging the exposure time to 72 h, the bacterial clusters with green fluorescence grew denser and aggregated to an initial form of a biofilm on pristine PDMS surface and (PAA/PEI)10 multilayer films(Fig. 7d and 7e). At the same time, a few of bacterial cells with red fluorescence were discernible and they were mainly associated with natural apoptosis during the bacterial growth process (Fig. 7f). As for the GS loaded multilayer films modified PDMS, the number of bacteria was much less than that on pristine PDMS. Moreover, there were a lot of bacterial cells with red fluorescence individually on the surface indicating the bactericidal property of the released GS. As it reported, the minimum inhibitory concentration (MIC) against S. aureus was around 0.13 µg/mL and it accounted for the bactericidal properties of the films (34). These results further demonstrated that the GS loaded multilayer films exhibited high biofilm inhibition efficiency against the Gram-positive S. aureus.

Fig. 7. Fluorescent microscopy images of S. aureus adhesions on (a) PDMS, (b) (PAA/PEI)10 multilayer films, (c) thermally cross-linked (PAA-GS/PEI)10 multilayer films at 4 h, (d) PDMS, (e) (PAA/PEI)10 multilayer films and (f) thermally cross-linked (PAA-GS/PEI)10 multilayer films at 72 h.

Cell viability assays

10


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Fig. 8. The cell viability assay of HLECs cultured on the surfaces of (a) TCPS, (b) pristine PDMS, (c) (PAA/PEI)10 multilayer film, (d) thermally crosslinked (PAA-GS/PEI)10 multilayer films and (e) untreated (PAA-GS/PEI)10 multilayer films for 24 h. The absorbance of the diluted Cell Counting Kit solution has been deducted from each data point and the statistical significance is indicated by different letters (p < 0.05).

Fig. 9. Growth and morphology of HLECs stained with FDA after 24 h of incubation on (a) TCPS, (b) pristine PDMS, (c) (PAA/PEI)10 multilayer film, (d) thermally crosslinked (PAA-GS/PEI)10 multilayer films and (e) untreated (PAA-GS/PEI)10 multilayer films for 24 h. under fluorescence microscopy (the magnification was 10×).

The biocompatibility of modified films is very important for biomedical applications. HLECs were used to test the cytotoxicity by cell morphology and activity evaluation with TCPS and PDMS as negative controls. As shown in Fig. 8, the cell viability of the HLECs on pristine PDMS, (PAA/PEI)10 multilayer films and the thermally treated GS loaded multilayer films was more or less the same as that on TCPS, which suggested the good biocompatibility of the multilayer films. However, the cell viability of the HLECs on the untreated GS loaded multilayer films was the lowest (83.7 % of that on TCPS), which also showed low cytotoxicity toward HLECs. 11



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The rapid release of GS from the untreated (PAA-GS/PEI)10 multilayer films contributed to the reduction of biocompatibility toward HLECs. In compassion, sustained of GS release from the thermally crosslinked (PAA-GS/PEI)10 multilayer films enhanced the biocompatibility. FDA assays were used to evaluate the morphology of HLECs on GS loaded multilayer films. The adhered HLECs on the surface were photographed with an inverted fluorescence microscope after FDA staining and the results were shown in Fig. 9. As it showed, the surfaces of TCPS and PDMS were good for HLECs growth and proliferation (Fig.9a and 9b). As for the (PAA/PEI)10 and thermally treated (PAA-GS/PEI)10 multilayer film, the cell density was less than that on controls but maintained normal spreading morphology. The adhesion of cells on different substrates is affected by various chemical and physicochemical factors, such as roughness and hydrophilicity(12). In this work, hydrophilic PAA and PEI constituents in the multilayer film with the ability to strongly bind water created a high hydrophilic surface. Even so, the good spreading morphology of HELCs on the surface indicated the low cytotoxicity of the drug loaded films.

Conclusions In this article, the (PAA-GS/PEI)n multilayer films were constructed to load small molecule charged antibiotics by pre-loading the antibiotics in polyanion solution and then assembled with polycation. The loading dosage of GS was high and also could be tuned through changing the number of bilayers. Thermal crosslinking method was used to delay the release of GS by introducing crosslinking points into the film to reduce the permeability of the film. The sustained release of GS could last more than 14 days and maintained effective antibacterial ability. The drug loading multilayer films were demonstrated efficacious against S. aureus and nontoxic toward HLECs. This approach represents a generalized strategy for incorporating charged small molecules into LBL films, which can be used as antibacterial surfaces for biomedical applications.

Materials and methods Materials. PAA (Mw: 25 kDa), poly(ethyleneimine) (PEI, branched, Mw: 25 kDa),

N-hydroxysulfosuccinimide

(NHSS),

and

1-ethyl-3-(3-dimethylaminopropyl)

carbodiimide (EDC) were purchased from Sigma-Aldrich. Staphylococcus aureus (S. aureus, ATCC 6538) was kindly provided by Prof. Jian Ji (Zhejiang University, Hangzhou, China). Polydimethylsiloxane (PDMS) was prepared from Sylgard®184 12


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from Dow Corning, according to the manufacturer’s instructions, using 10:1 ratio of elastomer base to curing agent. A Millipore MilliQ system (USA) was used to produce Ultrapure distilled water. Construction of the (PAA-GS/PEI)n multilayer films. Silicon wafers and PDMS used as substrates were cleaned with ethanol, acetone, and water for 10 min, respectively, and then dried with N2. Substrates were first pre-treated with PEI solution (5 mg/mL, 30 min) to form a precursor layer. For the constructing of (PAA-GS/PEI)n multilayer films, the substrates were alternately dipped in a PAA (1 mg/mL, pH 2.9) and GS (0.17 mg/mL ) mixture solution and PEI (1.0 mg/mL, pH 9.0) solution. The substrates were first immersed in the PAA-GS solution for 10 min, and then rinsed three times with buffer solution. The films were dried under a gentle stream of N2. Next, the substrate was immersed in PEI solution for 10 min and then also rinsed with buffer solution. This dipping cycle corresponds to the deposition of one bilayer. The cycle was repeated until the desired number of bilayers was reached.

Release of GS from (PAA-GS/PEI)n multilayer films Release of GS without any crosslinking treatment. The total GS loading dosage in the multilayer films was calculated through deducting the remaining GS in PAA-GS solution and two buffer solutions from input GS amount. For the release process, GS-loaded multilayer films were introduced into a 5 mL of PBS buffer solution (pH=7.4) for 168 h. For each point, 300 µL of the PBS buffer solutions were sampled and the release of GS was measured. Then, solutions were reintroduced in the plate for other points and the release kinetics profiles can be obtained. GS concentrations were obtained by a spectroscopy method(53). Firstly, the o-phthaldialdehyde reagent was formulated and stored in the dark environment. To test the GS concentrations, the GS samples, o-phthaldialdehyde reagent, and isopropanol were equally mixed stored for 30 min in the dark environment. The o-phthaldialdehyde solution was used to react with amino groups in the GS to obtain chromophoric products which have spectral absorption at 332 nm. After that, the absorbance of the mixed solution was measured using a spectrophotometer (Spectronic Instruments, Rochester, NY). A calibration curve ranging from 10 to 120 µg/mL was obtained to calculate the GS concentrations.

Chemical crosslinking method with EDC/NHSS. As it reported, the multilayer films were immersed in EDC/NHSS and GS complex solution (EDC: 10 mmol, NHSS: 20 mmol, GS: 0.1 mg/mL) in the 2-morpholino-ethanesulfonic acid (MES) 13



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buffer (0.1 M, pH 5.5) for 4 hours at 25 ºC to crosslink carboxyl groups of PAA and amino groups of PEI (42, 49). Then the release kinetics profiles were also tested as mentioned above.

Thermally crosslinking to control GS release. As it previous reported, the obtained multilayer films were firstly thermal crosslinking (4, 50, 54) before the implement of GS release. The amide bonds formed through thermally crosslinking at 150 ºC for 2.5 h between carboxyl groups PAA and amino groups of PEI. After thermal crosslinking procedure, the release of GS was done in PBS buffer solution.

Characterization of the multilayer films. The thickness of the self-assembly multilayer films on silicon wafer was measured by spectroscopic ellipsometry (M-2000 DITM, J.A. Woollam). The set parameters and test procedures were that: the wave length ranged from 124 to 1700 nm, both 65º and 70º were used as the angle of incidence. For data analysis, ∆ and Ψ values were set at wavelength ranging from 600 to1700 nm. The thickness of multilayer films was determined using Cauchy model with An and Bn as fit parameters set at 1.45 and 0.01 respectively. Then the thickness of the multilayer films can be automatically calculated. Then the thickness that best fit the multilayer films can be automatically calculated. FE-SEM (SiRion100) was also used to measure surface morphology and thickness of the multilayer film. When the desired number of the film was deposited, after the drying process, the silicon wafers with the multilayer films were dried and snapped. The cross-sectional images of the multilayer films were obtained to measure the thickness. In vitro antibacterial test. Both zone inhibition and bacterial LIVE/DEAD staining methods were conducted to measure antimicrobial properties of multilayer films with S. aureus as model bacteria. For the zone inhibition test, nutrient agar in Petri dishes were seeded with 0.2 mL of 1.0 × 106 cells/mL bacteria suspension before placing of antibacterial coating modified PDMS. After 24 h incubation at 37 ºC, the area clearing surrounded the film was measured as the zone of inhibition (ZOI). Furthermore, bacteria adhesion and viability were also determines using the LIVE/DEAD BacLight bacterial viability kit (L-7012, Invitrogen). For this method, bacterial structural integrity on the surfaces live or dead could be evaluated. Before testing, 10 mL 1.0×105 cells/mL S. aureus suspensions in PBS were incubated with the antibacterial coating modified and pristine PDMS for 4, 72 h. After being stained according to the kit protocol and washing, the PDMS sheets were kept in the dark and observed by fluorescence microscope investigation (Zeiss, Germany). 14


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Cytotoxicity assays Cell cultivation. The DMEM/F12 (1:1) cell culture medium containing 10 % fetal bovine serum, 100 U/mL penicillin, and 100 µg/mL streptomycin was used to incubate the human lens epithelial cells (HLECs, from ATCC, SRA01/04) in a 5 % CO2 incubator at 37 ºC. Confluent cells were digested to harvest the cells using 0.25 % trypsin and 0.02 % EDTA, followed by centrifugation (1000 g for 3 min). And then the cell concentration was calculated using haemacytometer and resuspended for incubation on the surfaces of materials. The HLECs with a density of 1.0×104 cells per sample were cultivated with the specimens in 96-well tissue culture plate for 24 h. Subsequently, the viability and morphology studies of HELCs on the samples were measured by fluorescein diacetate (FDA) and Cell Counting Kit-8 (CCK-8) methods. Cell viability. In this experiment CCK-8 (Beyotime, China) assay was employed to quantitatively evaluate the cell viability of multilayer films toward HLECs. After inoculating with the samples for 24 h, the HLECs were replaced by 100 µL 10 % FBS DMEM/F12 (1:1) mixed medium containing 10 µL CCK-8. The mixed medium was incubated to form water dissoluble formazan at 37 °C for 2 h. Then 100 µL of the formazan solution were aspirated from each sample with pipette and added to a new 96-well plate. The absorbance at 450 nm (calibrated wave) was examined using a microplate reader (Multiskan MK33, Thermo electron corporation, China). Tissue culture plates (TCPS) without any modified films were used as a control and six parallel replicates were prepared.

Cell morphology. Stock FDA solutions (5.0 mg/mL) were prepared by dissolving FDA in acetone. The working solution with the concentration of 5.0 µg/mL was freshly prepared by adding FDA stock solution into 0.1 M PBS. The membrane integrity and cytoplasmic esterase activity of cells on the surface could be examined using FDA (Sigma) as indicator for fluorescence microscope investigation (Zeiss, Germany) at 10× magnification in fluorescein filter, 488 nm excitation. After incubation with the specimens in the 96-well tissue culture plate for 24 h, FDA solution (20 µL) was added into the HLECs solution and incubated for 5 min. After washing with PBS twice, the fluorescence microscope examination was taken at the wavelength of 488 nm for each sample. TCPS that did not contain samples were used as controls.

Statistical analysis 15



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All experiments were conducted in triplicate, and data points were expressed as the mean. Two sample t test in origin 8.0 (Microcal, USA) were used to compare data obtained with the different samples under identical treatments. A value of p < 0.05 was considered significant.

Acknowledgements National Natural Science Foundation of China (51403158, 81271703, 31570959), the International Scientific &Technological Cooperation Projects (2012DFB30020), Natural Science Foundation of Zhejiang Province (LY12H12005) and Science & Technology Program of Wenzhou (S20140005) are greatly acknowledged.

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Direct Loading and Tunable Release of Antibiotics from

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