Loading of Antibiotics into Polyelectrolyte Multilayers after Self

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Loading of Antibiotics into Polyelectrolyte Multilayers after Self-Assembly and Tunable Release by Catechol Reaction Bailiang Wang, Zi Ye, Yihong Tang, Huihua Liu, Quankui Lin, Hao Chen, and Kaihui Nan J. Phys. Chem. C, Just Accepted Manuscript • DOI: 10.1021/acs.jpcc.6b00957 • Publication Date (Web): 07 Mar 2016 Downloaded from http://pubs.acs.org on March 10, 2016

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The Journal of Physical Chemistry C is published by the American Chemical Society. 1155 Sixteenth Street N.W., Washington, DC 20036 Published by American Chemical Society. Copyright © American Chemical Society. However, no copyright claim is made to original U.S. Government works, or works produced by employees of any Commonwealth realm Crown government in the course of their duties.

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The Journal of Physical Chemistry

Loading of Antibiotics into Polyelectrolyte Multilayers after Self-Assembly and Tunable Release by Catechol Reaction

Bailiang Wang,a, b* Zi Ye,a Yihong Tang,a Huihua Liu,b Quankui Lin,a, b Hao Chen,a, b Kaihui Nana, b*

a

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

Wenzhou, 325027, China

b

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] (K.H. Nan).

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ABSTRACT: The colonization of bacteria on biomedical implants and subsequent forming of biofilm often leads to serious infections. In this work, poly(acrylic acid) modified by dopamine (PAA-dopa) and poly(ethyleneimine) (PEI) self-assembled into multilayer films were used to load and sustained release of antibiotics. The obtained multilayer films were immersed in tris-buffered saline containing gentamicin sulfate (GS) to load the antibiotics and were cross-linked via catechol chemistry of dopa. Through raising pH value, more carboxyl groups of PAA were protonated and combined with GS through electrostatic forces. After cross-linking, the multilayer films were used as drug delivery system to embed GS in the matrices which showed a two-stage sustained release profile. In vitro, the release of GS into phosphate-buffered saline (PBS) from the multilayer films, as well as the antibacterial activity, was determined. The drug loaded multilayer films showed excellent bactericidal function against S. aureus and good cell compatibility toward human lens epithelial cells when the loading dosage of GS was 255 µg/cm2. In vivo, multilayer films modified and uncoated polydimethylsiloxane (PDMS) were implanted into subcutaneous incisions of New Zealand white rabbits. Both wound healing appearance evaluation and histopathology analysis demonstrated that implantation of the antibacterial coating modified PDMS had promoted wound healing and showed anti-inflammatory effect. In summary, this study presents a convenient and effective method for the incorporation of antibacterial agents into multilayer films.

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INTRODUCTION

The colonization of bacteria on implants and biomedical devices and forming infections remains a serious health care problem1, 2. Once bacteria adheres to the surface of implants and biomedical devices, bacterial colonies and biofilm will gradually form on the surfaces and damage the functionality of biomaterials. In the USA, at least one million persons died from implant-associated infections per year3. Bacteria can easily adhere on solid surfaces, form colonies and subsequently biofilms. The bacteria biofilm is a tight and organic structure which consists of exopolysaccharides, proteins, minerals and also nutrients and waste discharge channels4, 5. Due to the protection of the biofilm, bacteria in the matrix becomes more resistant to antibiotics than bacteria in the solution1,

6, 7

. In many cases, the

contaminated implants have had to be extracted to eliminate the infections because there is no suitable treatment method to deal with the situation. However, construction of antibacterial coatings on the surface of implants is an important approach to inhibit biofilm growth8-10. Furthermore, cationic polymers on the surface can not only directly kill bacteria through contact but also can be used as matrices for the delivery of antibacterial agents, such as silver ions or antibiotics11. Different techniques, involving chemical grafting12, surface impregnation13,

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, Langmuir–Blodgett

deposition15, 16, or physical entrapment17, 18, have been explored to immobilize active antimicrobial substances onto solid surfaces. To construct antibacterial surfaces, layer-by-layer (LBL) self-assembled multilayer films can be used as a multifunctional platform19, 20. It is a simple and 3



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versatile way to fabricate LBL films on virtually any substrates which involves alternate deposition of polymers with complementary noncovalent interactions21-23. According to the recent review by Rubner et al., there are mainly three strategies to design antibacterial surfaces through constructing LBL films: anti-adhesive, biocide leaching, and contact-killing surfaces10. Although three kinds of antibacterial surfaces have unique advantages and functions, each method rarely is able to play an ideal and long-lasting effect24. As for the surface with anti-adhesive property, many kinds of hydrophilic polymers such as poly(ethylene glycol) (PEG)1, 25, poly(ethylene oxide) (PEO) brushes26, and polyvinylpyrrolidone (PVP)27 have been used to create hydrophilic and antifouling coatings. However, it is impossible for such a surface to completely resist bacterial adhesion for a long time as it will inevitably become contaminated by various types of proteins and further bacteria adhesion1, 28, 29. For the bactericidal surfaces, it can efficiently kill bacteria, but bacteria corpse will permanently retain on such surfaces due to the lack of antifouling property30, 31. Biocide leaching surface can release antibiotics32-34 or silver ions35, 36 in a controlled manner to reduce the incidence of infections. However, continuous release of antibacterial agents also causes serious problems such as the emergence of resistant bacteria against antibiotics37. Therefore, locally and sustained delivery of antibacterial agents is a promising approach to efficiently kill bacteria with locally high concentration antibiotics and also to reduce the bacterial resistance issue 5, 38, 39. Mussels, through secreting adhesive proteins, can adhere to almost any kind of organic and inorganic surface in a liquid environments40, 4


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. Many kinds of

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low-molecular-weight catecholamine mimics have been artificially synthesized, such as dopamine (dopa) which can self-polymerize into an adherent coating on many kinds of substrates42, 43. The catechol moiety can strongly adhere to solid surfaces through many reactions, such as dismutation, Michael addition, or Schiff base reaction44, 45. The reactive catechol moiety can be introduced into polymer chains for further chemical modification, including cross-linking and further grafting of biomolecules. Yang et al. have synthesized poly(acrylic acid) (PAA-dopa) to construct LBL thin film with poly(vinylpyrrolidone) based on hydrogen bonding, which could be cross-linked via catechol chemistry of dopa41. A free-standing film which had a reversible swelling-shrinking behavior was obtained through changing the pH value making the hydrogen bond in the film break and reconstruct. Messersmith et al. synthesized catechol-containing polymers PEI-3-(3,4-dihydroxyphenyl) propionic acid (PEI-C) and dopa-HA (HA-C) to develop substrate-independent LBL films and in-situ reduction of Ag(I) to Ag(0) in the film by utilizing catechol functional groups as a reducing agent for antibacterial activity46. Gentamicin sulfate (GS) is an aminoglycoside antibiotic which is effective against gram-negative bacteria and certain gram-positive species, such as Staphylococcus aureus (S. aureus)32, 47. This antibiotic can also effectively kill methicillin-resistant S. aureus and inhibition the formation of biofilms. 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 character48, 49. Therefore, design and preparation of LBL multilayer films to enhance the loading 5



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capacity and to release the drug in a controlled manner is of great importance. The aim of the present study was on developing LBL multilayer films as nanocarriers to load a high capacity of GS by raising pH value and control the release behavior through crosslinking via catechol chemistry of dopa. 1H NMR and Fourier transforms infrared spectroscopy (FT-IR) were used to confirm the synthesis of PAA-dopa that consisted of a biomimetic adhesive side chain. Spectroscopic ellipsometry was used to measure the thickness of the multilayer film. The effects of pH value on GS loading dosage and the cross-linking of multilayer films on GS release rate were investigated. The antibacterial properties of drug loading multilayer films were investigated in detail both in vitro and in vivo.

MATERIALS AND METHODS

Materials. PAA (Mw: 25 kDa), dopa hydrochloride, poly(ethyleneimine) (branched PEI, Mw: 25 kDa), 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysulfosuccinimide (NHS) were purchased from Sigma-Aldrich. S. aureus (ATCC 6538) was kindly provided by Prof. Jian Ji (Zhejiang University, Hangzhou, China). Polydimethylsiloxane (PDMS) was prepared from Sylgard®184 from Dow Corning, according to the manufacturer’s instructions, using 10:1 ratio of elastomer base to curing agent. A Millipore Milli-Q system (USA) was used to produce Ultrapure distilled water. Synthesis of PAA-dopa. PAA-dopa was synthesized similarly to the method established by Sun et al41. PAA powder (3 mmol), EDC·HCl (0.6 mmol) and dopa 6


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hydrochloride (0.9 mmol) were dissolved in 20 mL of 0.1 M phosphate buffer solution (PBS, pH 6.0). The reaction continued for 24 h under continuous stirring and N2 gas protection. The product was obtained after dialyzing against deionized water using cellulose dialysis tube with a molecular weight cutoff of 8000 until there was no dopa detection in the washing solution. The dialysate was stored in the dark after being diluted with 0.1 M PBS solution. Construction of the (PAA /PEI)n and (PAA-dopa/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 pretreated with PEI solution (5 mg/m, 30 min) to form a precursor layer. For the underlying (PAA-dopa/PEI)n LBL film, the substrates were alternately immersed in PAA or PAA-dopa (pH 3.2, 1.0 mg/mL) and PEI (pH 10.8, 1 mg/mL) solutions. The substrates were first immersed in the PAA solution for 10 min, and then rinsed three times in buffer solution. The films were dried under a gentle stream of nitrogen gas. Next, the substrates were immersed in PEI solution for 10 min and then also rinsed with buffer solution. This dipping cycle corresponded to the deposition of one bilayer. The cycle was repeated until the desired number of bilayers was reached. Characterization of PAA-dopa and 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 such that the wave length ranged from 124 to 1700 nm, and both 65º and 70º were used as the angle of incidence. For data analysis, ∆ and Ψ values were set at 7



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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. The thickness of the multilayer films then can be automatically calculated. Proton Nuclear Magnetic Resonance (1H NMR) was measured with a Bruker 400 NMR spectrometer at 25 °C, using deuterium oxide (D2O) as solvent and tetramethylsilane (TMS) as the internal standard. FT-IR was measured on a FT-IR spectrometry (Bruker Optics). The samples were prepared as KBr disk. Loading and Release of GS into the Multilayer Films. The obtained multilayer films were cut into 1×2 cm2 and immersed in 5 mL GS (100µg/mL) solution at different pH value for 24h. To crosslink the multilayer films based on catechol chemistry, they were immersed in 5 mL GS (100µg/mL) solution in tris buffer (pH=8.5). For each point, 300 µL of the GS solutions were sampled and the loaded GS dosage was measured. Then, solutions were reintroduced in the plate for other points and the loading kinetics profiles obtained. GS concentrations were obtained by a spectroscopy method50. First, 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 and 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 8


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curve ranging from 10 to 120 µg/mL was obtained to calculate the GS concentrations. For the release process, GS-loaded multilayer films were introduced into a 5 mL of PBS buffer solution (pH=7.4) for 48 h. For each point, 300µL 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 could be obtained. Antibacterial Tests in vitro. 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 surrounding the film was measured as the zone of inhibition (ZOI). Furthermore, bacteria adhesion and viability were also determined 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 24, 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). 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 9



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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). Then the cell concentration was calculated using haemacytometer and re-suspended 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 10


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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. Antibacterial Tests in vivo. We obtained the approval of the local laboratory animal committee for this study. New Zealand White rabbits (weighing between 2.5 and 3.5 kg) were used in this experiment, which were obtained from Animal Administration Center of Wenzhou Medical University. The rabbits were treated according to the guidelines set forth by the Association for Research in Vision and Ophthalmology. Studies of GS-loaded (PAA-dopa/PEI)10 multilayer films coated PDMS and uncoated PDMS were implemented using twelve animals in each group, respectively. After being sterilized, all PDMS implants were stored in sealed petri dishes prior to implantation. The animal model used in this study has been similarly reported by Hansen51. Before the surgery of PDMS implantation, isoflorane delivered by facemask was used to anesthetize the rabbits. A broad area of the back was shaved, washed with surgical scrub, wiped with alcohol and betadine, and draped for surgery. Under sterile conditions, multilayer films-coated and uncoated PDMS implants were implanted into two subcutaneous symmetrical on either side of the spine. In order to eliminate 11



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cross-contamination risks between the implants there should be broad physical separation. In the rabbits, 10 mL of the 108 cells/mL S. aureus suspension was then inoculated into the subcutaneous pocket. 4-0 interrupted nylon suture was used to close the incisions in a single layer and the epidermis was cleaned with 2% H2O2. Adequate amounts of food and water were given to the individually housed rabbits. Throughout the next seven days, rabbits were observed daily before they were returned to the operating room. Each pocket was opened via a small incision distinct from the prior wound and two sterile cotton swabs were inserted through the incisions. Rabbits were then euthanized, and the incisions were extended. A closed centrifuge tube containing the implant and 5 mL PBS was vigorously vortexed for 30 s, followed by sonication for 5 min. The solution (0.2 mL) was plated onto the triplicate solid agar and incubated for 24 h. The number of viable bacteria was counted and the results were expressed as mean colony forming units (CFU) per mL. After obtaining culture samples, each pocket was excised en bloc, and transmural sections from representative areas taken. 10% formalin was used to fix in the specimens before embedding in paraffin blocks. Tissue sections about 5 mm thick were mounted onto slides and stained using Hematoxylin and Eosin. Statistical Analysis. 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.

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RESULTS AND DISCUSSION

Construction of the (PAA-dopa/PEI)n and (PAA/PEI)n Multilayer Films. Biomedical infections could be efficiently prevented through immobilization and controlled release of antimicrobial agents from biomedical surfaces. Improved drug release systems have been developed through selecting appropriate carriers that are capable of controlling the delivery of the drugs. There are two phenolic hydroxyl groups in one dopa molecule, which can be grafted on PAA. As shown in Scheme 1, synthesis of dopa-grafted PAA was carried out through carbodiimide chemistry. As shown in Scheme 2, PEI and the obtained PAA-dopa alternately were self-assembled into multilayer film. As a small trisaccharide molecule, GS can diffuse within the film in and out. The (PAA/PEI)n-GS multilayer films showed fast release of GS owing to the rapid diffusion and release of GS. On the other hand, (PAA-dopa/PEI)n-GS multilayer films cross-linked via catechol chemistry of dopamine displayed more prolonged release of the antibiotics.

Scheme 1. Schematic representations for the synthesis of PAA-dopa.

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Scheme 2. Schematic representation of construction, cross-linking, GS loading, and antibacterial properties of the (PAA-dopa/PEI)n multilayer films. The 1H NMR spectrum (Figure 1) and FT-IR spectrum (Figure 2) proved that dopa was successfully linked onto PAA chain. The 1H NMR had been extensively used to determine the grafting efficiency of dopa groups onto PAA. Figure 1 showed the typical 1H NMR spectrum of PAA-dopa sample. Peaks found at δ 4.79 ppm corresponded to deutriated water that was used for sample dissolving. Two peaks at δ 1.58-1.88 and 2.3 ppm corresponded to -CH(CH2-)- of H in the polymeric backbone. Peaks found at δ 2.7-2.8 and 3.0-3.2 ppm corresponded to protons of -CH2-CH2- in dopa. Peaks found between δ 6.6 and 6.8 ppm corresponded to C6HH2(OH)2 of dopa, which demonstrated the presence of dopa. Furthermore, the dopa grafting proportion on PAA was calculated by the formula of f= A/A0. The amount of H in the aromatic rings of grafted dopa molecules was represented as A through calculating the integral area of the peaks at δ 6.6–6.8. A0 was the integral area of the peaks at δ 1.4-2.5 ppm representing the amount of H in the polymeric backbone. The ratio of dopa grafted on PAA was about 14.5% according to this method.

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Figure 1. 1H-NMR spectrum of PAA-dopa.

Figure 2. FT-IR spectra of PAA-dopa. The peak at 1717 cm-1 was attributed to the C=O stretching vibration of carboxylic groups. The shoulder at 1627 cm-1 resulted from the C=O stretching vibration of secondary amide group. The peak at 1565 cm-1 was attributed to the

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combination of N-H deformation and C-N stretching vibration. The peak at 1452 cm-1 was assigned to backbone -CH2 deformation vibration. The appearance of an amide peak in the FTIR spectrum of PAA-dopa indicated that an amide bond has formed between PAA and dopa.

Figure 3. Ellipsometry measurement of the (PAA-dopa/PEI)10 and (PAA/PEI)10 multilayer films.

As shown in Figure 3, PAA-dopa at pH 3.2 and PEI under pH 10.8 were alternately deposited onto substrates to form multilayer films. The thickness of the (PAA-dopa/PEI)10 multilayer film increased exponentially into 1319.6 ± 92.5 nm for 10 bilayers number with the electrostatic interactions as the main driving force. As for the self-assembly of (PAA/PEI)10 multilayer films, the thickness of the multilayer film was a little thinner (1119.6 ± 132.5 nm), also showing exponential growth. The obtained multilayer films were immersed into GS solution to load the small GS molecules with cations, which combined with PAA through electrostatic forces. To change the penetrability of the multilayer films and to control the diffusion of the drug, 16


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GS was loaded in tris buffer at pH 8.5 and simultaneously catechol chemistry was applied to crosslink the film. At pH 8.5, carboxyl groups in PAA dissociated to combine with GS in the multilayer films. The presence of electrostatic forces between PAA-dopa and PEI contributed to the stability of the films. As the stability of the films is an important factor in its application. According to the literature, self-assembled multilayer films through electrostatic forces possess high stability against certain ionic strength and pH value which can even be used to construct free standing films52, 53. Loading of GS into (PAA-dopa/PEI)n Multilayer Films

Figure 4. Cumulative amount of GS loaded (a) into the (PAA-dopa/PEI)10 and (PAA/PEI)10 multilayer films as time evolution and (b) into (PAA-dopa/PEI)10 multilayer film with different pH values.

The loading of GS into the multilayer films was very fast which completed 80% of GS loading within 2 hours. GS was expected to diffuse within the film in and out as it had a small molecular weight. As indicated in Figure 4a, the GS loading dosage was 193 ± 48 µg/cm2 for (PAA-dopa/PEI) 10 multilayer film at 2 h and it steadily increased into 255 ± 11 mg/cm2 at 24 h (pH 8.5). And the GS loading dosage into (PAA/PEI) 10 17



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multilayer film was a little higher than that in (PAA-dopa/PEI) 10 multilayer film (265 ± 13 mg/cm2 at 24 h). It could be anticipated that as the increasing of bilayers number the loading dosage of GS could be precisely tuned. At the same time, the crosslinking of films based on dopa auto-polymerization in alkaline environment (tris buffer, pH 8.5) had also completed44. As indicated in Figure 4b, the loading pH value had great effect on the loading dosage. The GS loading dosage was 300 ± 34 µg/cm2 for (PAA-dopa/PEI) 10 multilayer film at pH 7.0 after 24 h. Through increasing pH value, more carboxyl groups in PAA dissociated to combine with GS in the multilayer films. However, a much higher pH value severely decreased the loading dosage of antibiotics, which could be due to the changes of the structure and permeability of the films. It has been challenging to directly incorporate small, uncharged, and hydrophobic therapeutics into multilayer thin films due to the nonpolymeric character. Only some groups have developed LbL multilayer films to direct load small molecular antibiotics as assembly components. Sukhishvili et al. constructed LBL multilayer films through self-assembly of tannic acid with one of the cationic antibiotics (tobromycin, gentamicin, and polymyxin B)22. The adhesion of bacteria on such films can lead to acidification of the environment around the films to trigger the release of the antibiotics. Hammond et al. has constructed LBL films using synthesized poly(β-amino esters) which controlled erosion and provided tunable drug release in physiological buffers48, 54. These films would degrade through a hydrolytic process that enabled continuous release of antibiotics55. 18


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Kinetics of GS Release from the Multilayer Films

Figure 5. Normalized cumulative GS release from the (PAA-dopa/PEI)10 and (PAA/PEI)10 multilayer films.

Below, we explored GS release behavior from multilayer films during long term exposure to physiological pH (pH 7.4).The films were immersed in 0.1 M PBS, an ionic buffer that was similar to the conditions of the human body fluid. These release curves normalized to the total released drugs displayed the difference of release rate from different kinds of films. As indicated in Figure 5, the (PAA/PEI)10-GS films completed 90% of its release of GS within 2 h. While the (PAA-dopa/PEI)10-GS films cross-linked by dopa auto-polymerization displayed more prolonged release of GS (32% at 2 h). As the time prolonged to 8 h, the (PAA/PEI)10-GS films completed 99% of the GS release. On the other hand, the cross-linked films only completed 46% of the GS release at 8 h and 88% at 168 h showing sustained release of the antibiotics. The rapid release of GS from the (PAA/PEI)10-GS films could be due to the rapid diffusion of small molecular drug in PBS with high ionic strength. It is well known that through auto-polymerization dopa could also react with amino groups in PEI forming multiple 19



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cross-linking points in the films. Crosslinking reduced the permeability of the films and prolonged the diffusion of the drug. The large advantages of using LbL systems are the ability to load high capacity of drug and to coat complex material surfaces such as implants and microscopic size structures. The rapid growth of self-assembly multilayer film with exponential growth behavior is a promising class of LBL films which can be constructed using weak polyelectrolytes. Furthermore, the loading dosage and release process can be regulated by changing the type of polyelectrolyte component that serve as adsorption centers for antibiotics56, 57. For most of the therapeutic agents, incorporation and controlled release of low molecular weight drugs becomes an issue due to the weak binding forces. Meanwhile, cross-linking reduced the permeability of the films and sustained the release of antibiotics. It was found by Shen et al. 53 and Hammond et al. 58

that thermal crosslinking greatly enhanced the stability of the films and reduced the

degradation rate of the PEM films. Cell Viability Assays

Figure 6. The cell viability assay of HLECs cultured on the surfaces of (a) TCPS, (b) pristine PDMS, (c) (PAA/PEI)10 multilayer film, (d) (PAA/PEI)10-GS multilayer films 20


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and (e) (PAA-dopa/PEI)10-GS 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).

Figure 7. 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) (PAA/PEI)10-GS multilayer films and (e) (PAA-dopa/PEI)10-GS multilayer films for 24 h. under fluorescence microscopy (the magnification was 10×).

The biocompatibility of modified films has a great impact on the biomedical applications. After incubation with HLECs, both CCK-8 and FDA assays were used to evaluate the cytotoxicity of the GS-loaded multilayer films. As shown in Figure 6, the cell viability of the HLECs on the pristine PDMS, (PAA/PEI)10, (PAA/PEI)10-GS and (PAA-dopa/PEI)10-GS multilayer films modified PDMS was more or less the same as that on TCPS. It indicated the low cytotoxicity of the GS-loaded multilayer films. FDA assay was used to evaluate the morphology of HLECs on multilayer films with PDMS and TCPS as negative controls. According to Figure 7, the surface of TCPS 21



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was favorable to HLECs growth and proliferation. On the other hand, the cell density on the GS-loaded films was less than that on TCPS but maintained normal spreading morphology that showed the low cytotoxicity of the drug loading films.

Antibacterial Property of the Multilayer Films Zone of Inhibition (ZOI) Assays

Figure 8 Inhibition zones of (a) PDMS, (b) (PAA-dopa/PEI)10-GS film, (c) (PAA/PEI)10-GS film, (d) (PAA/PEI)10-GS and (e) (PAA-dopa/PEI)10-GS films after being immersed in PBS for 168 h against S. aureus.

The ability of antibacterial films coated with PDMS to inhibit the growth of S. aureus was examined using a modified Kirby-Bauer test on a bacteria coated agar plate. 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 ZOI of the (PAA-dopa/PEI)10-GS film was around 7.0-8.0 mm (Figure 22


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8b), whereas the PDMS did not show any bacterial growth inhibition (Figure 8a). Furthermore, the ZOI of (PAA/PEI)10-GS film was 9.5 ± 10.5 mm which was larger than that of (PAA-dopa/PEI)10-GS film (Figure 8c). The long-acting bactericidal activity of the drug loaded films was also tested after being immersed in PBS for 168 h. As shown in Figure 8d, there was no ZOI around the (PAA/PEI)10-GS film indicating the absence of GS release from the film. On the other hand, as for the (PAA-dopa/PEI)10-GS multilayer films, the sizes of inhibition zone decreased to 3.5-4.5 mm showing the continuous release of GS (Figure 8e). It confirmed that the crosslinking reaction via catechol chemistry of dopa prolonged the release of GS and retained the antibacterial property for a much longer time than (PAA/PEI)10-GS film. Either drug concentrations below the minimum inhibitory concentration (MIC) or inappropriate delivery timescales will lead to the development of antibiotic resistant bacteria59, 60. As for the GS-loading multilayer films in this work, there were two release phases: rapid release early and slow release later of the GS. The corresponding antibacterial function originating from the release behavior of GS could well meet the demands of the implant. The drug delivery system developed in this work were comparable to other drug delivery platforms developed by other researchers48, 49. Furthermore, the high loading dosage of drug into the antibacterial film could lead to a high GS concentration above the MIC of GS against common bacteria. As reported, the MIC against S. aureus was about 0.13 µg/mL and it accounted for the bactericidal properties of the films61. These results demonstrated that the GS loaded multilayer films had high bacterial inhibition efficiency against the Gram-positive S. aureus. 23



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Bacterial LIVE/DEAD Staining Methods

Figure 9. Fluorescent microscopy images of S. aureus adhesions on (a) PDMS, (b) (PAA/PEI)10-GS film, (c) (PAA-dopa/PEI)10-GS film at 24 h and (d) PDMS, (e) (PAA/PEI)10-GS film, (f) (PAA-dopa/PEI)10-GS film at 72 h.

In addition, the viability of bacteria on films modified PDMS was evaluated by a LIVE/DEAD BacLight bacterial viability kit. The surface treated PDMS displayed a promising reduction in bacterial adhesion. As shown in Figure 9, with an appropriate mixture of the SYTO 9 and propidium iodide stains, bacteria with intact cell membranes stained fluorescent green, whereas bacteria with damaged membranes stained fluorescent red. After 24 h of culture, a large number of live bacteria with green fluorescence, either individually or in small clusters was found on pure PDMS, as shown in Figure 9a. In Figure 9b and c, a lot of adhered bacteria was noticed on (PAA/PEI)10-GS and (PAA-dopa/PEI)10-GS films. However, majority of the single or colonized bacteria was found dead with only few of them surviving on the surface. It 24


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could be due to the rapid release of GS into the bacteria solution and fast killing of the bacteria. At the time of 72 h, the bacteria cluster with green fluorescence on pristine PDMS surface grew more and aggregated to early biofilm structure (Figure 9d). There was a strong reduction (approximatively 80%) of S. aureus on the (PAA/PEI)10-GS film (Figure 9e). However, the number of bacteria surviving on the surface was much larger than that at 24 h which could be due to the absence of GS release. As the sustained release of GS from (PAA-dopa/PEI)10-GS film, much less bacteria was found on the surface (approximatively 95% reduction comparing with PDMS) that was probably related to the bactericidal function of GS (Figure 9f). Overall, the results showed a reduction of biofilm forming on all modified surfaces, thus confirming the success of controlled release of GS. Animal Experiments of Antibacterial Activity

Figure 10. Photographs of pockets implanted with pristine PDMS (a) after operation, (b) at day 3, (c) at day 7 and pockets implanted with (PAA-dopa/PEI)10-GS films (d) after operation, (e) at day 3 and (f) at day 7.

25



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Figure 11. Photographs of implant pockets, opened at seven day. The images of internal of the pockets: (a) pristine PDMS, (d) (PAA-dopa/PEI)10-GS film coated PDMS and images of hematoxylin and eosin stained sections of surrounding connective tissues of (c) pristine PDMS and (d) (PAA-dopa/PEI)10-GS film coated PDMS.

Staphylococci have mainly lead to the infections both on temporarily and permanently implanted biomaterials62. The clinical experience with foreign body associated infections clearly shows that with the use of antimicrobial films with proven in vitro activity, it is often impossible to deal with those infections occurred on the implanted device63. Therefore, we developed a New Zealand white rabbit model for surgical site infections to assess the potential of the (PAA-dopa/PEI)10-GS film using in vivo. Briefly, subcutaneous incisions were made in the abdominal subcutaneous of anesthetized rabbits and closed with silk sutures after implanting of samples. There were 24 rabbits in total with twelve animals each for pristine and 26


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(PAA-dopa/PEI)10-GS film modified PDMS for the seven days observation period. Additionally, when we opened the surgical site and pulled back the skin to examine for subcutaneous inflammation, all twelve rabbits implanted with pristine PDMS had visible signs of inflamed skin and redness in the hide and tissue beneath the sutures at day 0, 3 and 7 (Figure 10a-c). Comparatively, the (PAA-dopa/PEI)10-GS film treated animals had no or minimal visible signs of inflammation within or under the sutures (Figure 10d-f). Specifically, the healing of the surgical sites implanted with antibacterial film modified PDMS was very good without any redness. As indicated in Figure 11a-b, at day 7, significant amounts of purulence were found in all bacterially challenged surgical sites implanted with pristine PDMS, which indicated severe infections. Whereas wounds into which PDMS coated with antibacterial films were free of purulence and culture negative. A significantly lower infection rate was observed for antibacterial film coated PDMS that released GS. The subcutaneous implantation enabled an accurate examination of the tissue response. As shown in Figure 11c-d, after 7-day incubation, tissues adhering to the implants were then processed for hematoxylin and eosin stained sections (H&E staining). As shown in Figure 11d, the surrounding connective tissues of multilayer films modified PDMS appeared vascularized with little indication of residual inflammation. The number of multinucleated giant cells showed statistically significant differences between antibacterial films modified and pristine PDMS. It was worth noting that the wound that implanted with pristine PDMS was dominated by acute inflammatory response, which well corresponded to its poor wound healing effect and increased size of 27



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wound as well. The cavity was surrounded by a relatively thick layer of chronic active inflammation and necrosis (Figure 11c). Therefore, the incorporation of GS to the LbL film gave a clear inhibitory effect of the local infection post-implantation.

CONCLUSIONS

The present study demonstrates convincingly the efficacy of (PAA-dopa/PEI)n film as a drug delivery system. It is a facile and versatile approach to control the release of antibiotics by exploiting the self-crosslinking property of catechol containing polymers. The (PAA-dopa/PEI)n multilayer film grew exponentially into a thick film which can be used to load high dosage of antibiotics. The sustained release of GS from the multilayer films endows the films with satisfactory antibacterial properties with noncytotoxicity toward HLECs. Implants of (PAA-dopa/PEI)n multilayer film modified PDMS loaded GS showed promising ability for treating infection in rabbit by providing a sustained release of the antibiotics. Overall, the bio-inspired LbL assembly approach based on catechol containing polymers will provide a general mean to deliver active therapeutics in biomedical applications for therapeutic surface coatings with controllable release properties. Acknowledgements National Natural Science Foundation of China (51403158, 81271703, 31570959), and Natural Science Foundation of Zhejiang Province (LY12H12005) are greatly acknowledged.

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Loading of Antibiotics into Polyelectrolyte Multilayers after Self

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