Preparation, degradation and in vitro release of ciprofloxacin-eluting ureteral stents for potential antibacterial application

Materials Science and Engineering C 66 (2016) 92–99

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Preparation, degradation and in vitro release of ciprofloxacin-eluting ureteral stents for potential antibacterial application Xiaofei Ma a, Yan Xiao a,⁎, Heng Xu c, Kun Lei a, Meidong Lang a,b,⁎ a b c

Key Laboratory for Ultrafine Materials of Ministry of Education, School of Materials and Science and Engineering, East China University of Science and Technology, Shanghai 200237, China Shanghai Collaborative Innovation Center for Biomanufacturing, 130 Meilong Road, Shanghai 200237, China Collaborative Innovation Center for Petrochemical New Materials, Anqing, Anhui 246011, China

a r t i c l e

i n f o

Article history: Received 15 January 2016 Received in revised form 18 April 2016 Accepted 21 April 2016 Available online 22 April 2016 Keywords: Ciprofloxacin lactate Release mechanism Ureteral stents Degradation Coating Antibacterial Cytotoxicity

a b s t r a c t Drug-eluting stents with biodegradable polymers as reservoirs have shown great potential in the application of interventional therapy due to their capability of local drug delivery. Herein, poly(L-lactide-co-ε-caprolactone) (PLCL) with three different compositions as carriers for ciprofloxacin lactate (CIP) was coated on ureteral stents by the dipping method. To simulate a body environment, degradation behavior of PLCL as both the bulk film and the stent coating was evaluated in artificial urine (AU, pH 6.20) respectively at 37 °C for 120 days by tracing their weight/Mn loss, water absorption and surface morphologies. Furthermore, the release profile of the eluting drug CIP on each stent exhibited a three-stage pattern, which was greatly affected by the degradation behavior of PLCL except for the burst stage. Interestingly, the degradation results on both macroscopic and molecular level indicated that the release mechanism at stage I was mainly controlled by chain scission instead of the weight loss or morphological changes of the coatings. While for stage II, the release profile was dominated by erosion resulting from the hydrolysis reaction autocatalyzed by acidic degradation residues. In addition, ciprofloxacinloaded coatings displayed a significant bacterial resistance against E. coli and S. aureus without obvious cytotoxicity to Human foreskin fibroblasts (HFFs). Our results suggested that PLCL copolymers with tunable degradation rate as carriers for ciprofloxacin lactate could be used as a promising long-term antibacterial coating for ureteral stents. © 2016 Elsevier B.V. All rights reserved.

1. Introduction Indwelling ureteral stents have extensively been used to facilitate urine drainage during surgery of ureteral stones and to treat benign/malignant ureteral obstruction or urogenital disorders [1–3]. In the past few decades, various polymers including silicone, polyurethane and their derivatives have clinically been applied as host materials for ureteral stents [4,5]. For example, silicone as the 1st generation of polymeric ureteral stent was relatively biocompatible but the high frictional coefficient has restricted its function optimization [5]. Currently the most widely applied stent with PU-related materials has encountered problems of bacterial adhesion and biofilm formation on its surface, potentially leading to urinary tract infection (UTI) and encrustation [2,5]. Therefore, the stent surface which directly contacts with urinary tract has played a key role in device-associated complications, for which the infection became the main reason [6,7]. Recently, biodegradable ureteral stents based on PLA, PGA and PCL whose surface was constantly changing as the stent degrades could effectively prevent bacterial adhesion, condition film deposition and resultant encrustation [1–3]. ⁎ Corresponding authors at: East China University of Science and Technology, 130 Meilong Road, P.O. Box: 391, Shanghai 200237, China. E-mail addresses: [email protected] (Y. Xiao), [email protected] (M. Lang).

http://dx.doi.org/10.1016/j.msec.2016.04.072 0928-4931/© 2016 Elsevier B.V. All rights reserved.

However, high content degraded fragments eventually led to inflammation and even ureteral obstruction in some patients [8]. Hence, how to design an ideal stent surface became one of the most challenging issues with the aim to overcome the stent-associated complications. Surface modification has been considered as a direct and efficient approach to achieve the abovementioned surface. Several coatings formulated from different materials to improve the performance of ureteral stents have been tried. For example, hydrogel-coatings were attempted to prevent biofilm formation by a thin hydration layer, which reduced bacterial adhesion on the stent surface [9]. Polyvinylpyrrolidone (PVP) has been used as a coating material for its excellent hydrophilicity and lubricity [10,11]. Tunney et al. [12] found that adherence of hydrophobic E. faecalis and struvite encrustation in vitro on PVP-coated PU was less than that on uncoated PU. Hydrophobic diamond-like carbon (DLC) was also applied as a polymeric coating for ureteral stents with the improved resistance toward bacterial adhesion, infection and encrustation [13]. Other coatings of natural antibacterial substances such as hyaluronic acid, heparin have also attracted much interest to elongate the antimicrobial duration [14–16]. Although a variety of materials have been designed to prevent the surface from bacteria adhesion, they were normally unable to inhibit the bacteria growth and proliferation because the bacteria were kept alive. Contact-active biocidal strategies with silver [17–19], chitosan

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[16] and other bactericidal polymers [20] as coatings had been applied to prevent bacterial infections. Sun et al. [21] prepared an N-halaminebased rechargeable antimicrobial and biofilm controlling polyurethane through covalently grafting. Gultekinoglu et al. [22] had ureteral stents covalently grafted with PEI and constructed a stable non-leaching antibacterial surface, which exhibited a permanent antibacterial activity against K. pneumonia, E. coli and P. mirabilis. Another effective biocidal strategy to overcome stent-related infections was surface modification based on releasing bactericidal agents such as cationic nanostructures [23], salicylic acid [24] and antibiotics [25], among which ureteral stents coated with antibacterial drugs showed great potential to eventually overcome stent related infection and biofilms by preventing bacterial growth or killing bacteria. The originally antibacterial drug-coated stents were simply prepared by coating surfaces with antibiotics without any additives as drug carriers. For example, ureteral stents coated with norfloxacin, ciprofloxacin, gentamicin, and nitrofurazone, prepared by dipping into drug solution had successfully prevented bacterial adhesion for short-term devices [26,27]. However, the initial burst of the drug has increased its local concentration and caused the damage of cells [25,28]. Recently, drug-coated ureteral stents by incorporating antibiotics into bioabsorbable polymer coatings became one plausible approach to overcome stent-induced infection for its sustained drug release profile with polymer degradation [28]. Ofloxacin-blended polylactone-coated SR-PLLA stents may provide an early protection for the stents from uropathogens up to 2 weeks by sustained release of ofloxacin [29]. However, the long-term antibacterial effect had not been studied. Dave et al. prepared an enzyme-embedded polycaprolactone (PCL)-based coating, incorporated with the antibiotic gentamicin sulfate (GS), which exhibited sustained release of GS with PCL degradation triggered by lipase B [28]. Although the feasibility of an antibiotic-plus-enzyme-loaded coating for preventing bacterial multiplication on stents surface had been testified, potential biocompatibility and safety of lipase B need to be further verified. In this study, bioabsorbable copolymers with tunable degradation rate as carriers for antibacterial drugs were expected to achieve a long-term sustainable release with the aim for the prevention of bacterial growth on stents surface. In particular, ciprofloxacin lactate was chosen as the antibacterial model drug, which was encapsulated in PLCL with different compositions as coating, to investigate the drug release profiles of ciprofloxacin-eluting stents. The degradation process of PLCL as both the film and the stent coating was evaluated in artificial urine to understand the impact of coating polymers on drug release mechanism. The effect of drug dose on release profiles and antibacterial assay against E. coli and S. aureus were investigated. Furthermore, in vitro cytotoxicity of CIP-loaded PLCL5050 coatings with four different drug dose was estimated by CCK-8 assay toward HFFs. Overall, these results further advanced the potential antibacterial application of ciprofloxacin-eluting ureteral stents with PLCL as carriers. 2. Experimental section 2.1. Materials Ciprofloxacin lactate was purchased from Aldrich and used as received. Polyurethane ureteral stents from Zhangjiagang Shagong Medical Technology Co., Ltd were cut into pieces with 30 mm length. Dulbecco's modified eagle medium (DMEM) was purchased from Gibco. Fetal bovine serum (FBS) was bought from BioSun. Cell Counting Kit-8 was purchased from DOJINDO. All other commercially available reagents and solvents were used without further purification unless otherwise mentioned. Poly(L-lactide-co-ε-caprolactone) copolymers were synthesized according to a previously reported procedure [30]. Typically, predetermined amounts of L-LA and ε-CL were simultaneously added to a flame-dried flask under argon flow. The flask was then degassed under vacuum at 45 °C for 4 h to remove the excess moisture. Afterwards, Sn(Oct)2 (0.0005 equiv of monomer) in dry toluene

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was added under the protection of argon. The polymerization was proceeded at 140 °C for 24 h. The products were purified by precipitation into cold methanol and further dried under vacuum. 2.2. Preparation of PLCL-coated/CIP-eluting stents A schematic preparation process of PLCL-coated stents and CIPeluting stents was shown in Scheme 1. The details were described as follows. The ureteral stents of 30 mm in length were coated with PLCL copolymers with a solution of 200 mg PLCL dissolved in 5 mL of THF. The stents were dipped into the polymer solutions and taken out quickly. Afterwards, the stents were kept in a fume chamber for 24 h and the coating process was then repeated for four times. After another 24 h of evaporating in the fume chamber, all the coated stents were dried under vacuum at 45 °C for 72 h to completely remove THF. The ciprofloxacin-eluting ureteral stents were prepared by abovementioned dipping method with a polymer-drug solution of PLCL (200 mg) dissolved in THF (2 mL), and ciprofloxacin lactate (15 or 30 mg) dissolved in DMSO (1.5 mL), which were mixed to obtain a homogeneous solution. After coating for four times, the drug-eluting stents were dried in oven at 50 °C for 72 h and then kept in a vacuum chamber for another 72 h. The loaded drug was calculated from the weight of the stent coating and the drug weight percent in the coating solution [31,32]. 2.3. In vitro degradation of PLCL-coated stents The degradation assay was carried out by immersing polymercoated stents and bare ureteral stents as control in artificial urine [24] (12 mL per stent, pH 6.20), at 37 °C with gently shaking to mimic in vivo environment for 120 days. The medium was withdrawn and replaced with fresh AU once a month. At different periods of time, stents were washed with distilled water in order to remove the residual salts and dried to a constant weight. The weight loss of stents and the water absorption of the stents during the degradation were measured by a gravity method and calculated by Eqs. (2-1) and (2-2), respectively. %Weight loss ¼ ðW0 −Wd Þ=W0

ð2  1Þ

%Water absorption ¼ ðWt −Wd Þ=Wd

ð2  2Þ

where W0 is the initial weight, Wd and Wt are the final dried and wet weight at predetermined point, respectively (n = 3). 2.4. Drug release from ciprofloxacin-eluting ureteral stents Ciprofloxacin-eluting ureteral stents were immersed in artificial urine (AU, pH 6.20) and incubated at 37 °C with shaker Incubator (100 rpm) for 120 days. At predetermined time intervals, 5 mL of artificial urine was withdrawn and replaced with the same volume of prewarmed fresh media. The eluted drugs were measured via UV–Vis Spectrometer at 275 nm. A standard curve with known concentrations of ciprofloxacin lactate was used to calculate the amount of released ciprofloxacin. The data were given as mean ± standard deviation (SD) based on three independent measurements. 2.5. Preparation and in vitro degradation of PLCL films PLCL films were prepared by solvent evaporation using dichloromethane as solvent. The polymer solutions with the concentration of 100 mg/mL were transferred to Teflon molds and kept in a fume chamber for 24 h at room temperature, followed by vacuum drying for another 48 h. The in vitro degradation was carried out by immersing square pieces (10 × 10 mm) which were cut from PLCL parent films into the artificial urine (AU, pH = 6.20). The medium was withdrawn and replaced with fresh AU once a month. At each scheduled time,

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Scheme 1. The schematic diagram of PLCL-coated/CIP-eluting stents preparation.

films were washed with distilled water and dried till a constant weight. The changes in number molecular weights (Mn) of PLCL during the degradation were measured by GPC (Waters 1515, USA).

Elx800, USA) at 450 nm. HFFs with blank glass substrates were used as control. The relative cell viability was calculated according to the following formula.

2.6. Scanning electron microscopy (SEM)


Cell viability ð%Þ ¼ ODsample −ODblank =ðODcontrol −ODblank Þ  100%

The surface morphologies of PLCL-coated ureteral stents during degradation were measured by SEM (JSM-6360LV; JEOL Ltd., Tokyo, Japan) with an accelerating voltage of 15 kV. Cross-sections of the stents were examined after fracturing in liquid nitrogen to identify the thickness of coating. 2.7. Antibacterial assessment E. coli (ATCC 25922) and S. aureus were selected as the model bacteria because they are the main cause for biofilms and UTI. Antibacterial activity of the coatings and ciprofloxacin-loaded coatings against E. coli and S. aureus was determined by the plate count method as described in a previous report [33]. E. coli and S. aureus were incubated at 37 °C with shaking at 200 rpm in yeast broth and nutrient broth, respectively. 100 mg of each coating membrane was immersed in a sterile flask, in which 10 mL of bacteria culture was added. The flasks were placed in the shaker incubator for up to 54 h. At selected time intervals, 0.1 mL of bacterial culture was taken out and added into 0.9 mL of sterilized physiological saline, and then decimal serial dilutions were prepared. Properly diluted suspensions of E. coli and S. aureus were spread on an agar media (containing 10 g/L peptone, 10 g/L sodium chloride, 5 g/L yeast and 20 g/L agar). The Petri dishes were sealed and incubated at 37 °C for 24 h. The bacterial cell colonies were imaged using a digital camera and counted. 2.8. Cell culture HFFs generated from human foreskin were cultured in DMEM containing 10% heat-inactivated FBS, 100 U/mL penicillin and 100 U/mL streptomycin on tissue culture flask at 37 °C in a humidified 5% CO2 atmosphere.

ð2  3Þ

where ODsample, ODcontrol and ODblank represented the value from sample, control and blank wells respectively. 3. Results and discussion PLCL copolymers were synthesized by ring-opening polymerization using trace water as initiator and Sn(Oct)2 as catalyst [30,34]. Table 1 summarized the PLCL copolymers which were used as coatings and drug carriers in this study. It was shown that Mn of copolymers with different compositions was in a comparable range from 193 to 211 KDa (and PDI from 1.67 to 1.78), which were characterized by GPC. PLCLcoated stents were prepared by dipping precleaned stents into each copolymer solution with the concentration of 40 mg/mL for four times. As shown in Fig. 1, PLCL5050-coated stents exhibited a smooth, uniform and nonporous surface and approximately a 30 μm thick of the coating was measured by the cross-section image. It was also observed for the other two PLCL copolymer coatings that were homogeneous and ca. 30 μm thick. 3.1. Evolution of PLCL-coated stents To better understand the effect of PLCL coatings on the ciprofloxacin release profiles, we evaluated the degradation progress of the PLCLcoated ureteral stents in artificial urine (pH 6.20, 37 °C) for 120 days. Fig. 2 plotted the weight loss and water absorption with degradation time for three PLCL-coated stents. According to the literature, a typical process of PLCL bulk degradation exhibited a sharp weight loss after the first few weeks when the sample weight remained constant [35, 36]. When coated on tubular ureteral stents, PLCL6040 and PLCL4060 exhibited a clear two-stage of weight loss during the degradation. At

2.9. In vitro cytotoxicity The cytotoxicity of blank PLCL5050 coatings and CIP-loaded PLCL5050 coatings with four different drug dose (3.5%, 7.5%, 11.5% and 15%) was estimated by CCK-8 assay toward HFFs. All the coatings (4 mg) fabricated by spin-coating were sterilized with 75% ethanol for 2 h and UV light irradiation for 30 min before cell seeding. HFFs were seeded at 2 × 104 cells per well on the polymer films, and cultured in DMEM supplemented with 10% FBS at 37 °C in a humidified 5% CO2 atmosphere for 24 h. Thereafter, CCK-8 solution was added to culture for 2 h and the absorbance was recorded by microplate reader (Bio Tek,

Table 1 Characterization data of the PLCL in the study. Coating polymer

PLCL6040 PLCL5050 PLCL4060

%Molar ratioa % LA

% CL

59 50 39

41 50 61

Mnb (KDa)

PDIb

211 193 211

1.78 1.67 1.72

a Calculated by 1H NMR spectra from the relative intensity of the methine of PLLA at 5.15 ppm and the ε methylene signal of ε-PCL at 4.1 ppm. b Determined by GPC (Waters 1515, USA).

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Fig. 1. The surface (a) and cross-section (b) SEM images of stents after coating.

stage I (0–60 days), they degraded gradually with a relatively slow rate. But at stage II (60–120 days), a sharp increase of mass loss was presented in both PLCL6040 and PLCL4060. On the contrary, PLCL5050 coating displayed a relatively fast rate of weight loss throughout the degradation time, with a nearly liner weight loss profile. As can be seen in Fig. 2(b), all the coating copolymers also presented a relevant increment in water absorption with the degradation evolved. At stage II, PLCL6040 exhibited the highest water uptake capacity (WA = 40.8%) followed by PLCL5050 (WA = 30.5%) and PLCL4060 (WA = 25.1%). However, a sharp decrease in water absorption after 90 days can be found in all the coatings due to the collapse of the coated materials from the stents. On day 120, PLCL6040 had a highest weight loss of 85.4% and a WA of 33.3%, while PLCL5050 exhibited values of weight loss of 80.1% and WA of 30.0%. On the contrary, PLCL4060 presented the least WL of 55.4% and WA of 15.1%. Based on the aforementioned results, a gradual WL and WA of all samples at stage I can be mainly attributed to the slow hydrolytic degradation [37]. Moreover, accelerated weight loss and water absorption after 60 days, especially in PLCL6040 with lager LA proportion, perfectly proved that bulk erosion dominated the degradation process at stage II [38]. It was observed from the degradation curves that the order in weight loss and water absorption during stage I was appeared as following, PLCL5050 N PLCL4060 N PLCL6040. At this stage, it was observed from Fig. 3 that tiny cracks accompanying with micro holes were firstly appeared on the surface of PLCL5050-coated stents at 30 days, leading to its fastest changes in WL and WA. In comparison to PLCL6040 coating without any cracks and channels at 60 days, PLCL4060 displayed big fissures on partial surface, resulting in a relatively higher weight loss. Therefore, at stage I the surface morphological change was in good

agreement with the evolution of weight loss and water absorption for all samples. After 60 days, both PLCL6040 and PLCL5050-coated stents lost morphological integrity, consistent with their larger mass loss and water uptake. In contrast, PLCL4060 with the highest CL proportion did not show any significant changes of surface morphology. On the basis of the above, the morphological changes on the stent surface during the degradation visually illustrated the trend in weight loss and water absorption of all three coatings. 3.2. Ciprofloxacin release profile PLCL copolymers with tunable degradability by altering the molar ratio of LA and CL were then used as carriers for the controlled release of ciprofloxacin lactate to prevent bacterial growth on the ureteral stent surfaces. CIP-coated ureteral stents with 7.5% and 15% drug dose of three PLCL copolymers with different compositions were prepared by dipping methods as described in the experimental section. By changing the CIP concentrations in coating solvents, we could prepare CIPeluting stents with relatively low or high drug loading. As shown in Table 2, 630 μg and 1300 μg of ciprofloxacin were approximately encapsulated per stent in 7.5% dose and 15% dose, respectively. The slight standard variation of the drug loading was probably attributed to the faintly different lengths of the stents. Fig. 4 displayed the release profiles of ciprofloxacin with low and high doses from three PLCL coatings on ureteral stents. It was found that the CIP release profiles of all the CIP-coated ureteral stents revealed a typical tri-phase release profile. From 0 to 7 days, a sharp burst release could be witnessed, which was considered as the immediate dissolution of the nonencapsulated ciprofloxacin on the surface or the easily

Fig. 2. The weight loss (a) and water absorption (b) of ureteral stents with different coating polymers.

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Fig. 3. SEM images of ureteral stents with different coatings during the degradation at low (Scale bar, 100 μm) and high (Scale bar, 10 μm) magnification. (a) PLCL6040 coating. (b) PLCL5050 coating. (c) PLCL4060 coating.

accessible hydration of drug particles close to the surface [32]. After the burst elution, there was a nearly linear and slow release phase from 7 to 60 days determined as stage I, which was correspondent to the same stage in the degradation. During this zero-order release stage, ciprofloxacin lactate was mainly delivered by the diffusion both from matrix and micro-porous structures formed by PLCL degradation. The final stage (60–120 days) exhibited a sharp increase in the release rate, also in accordance to the stage II when the degradation was dominated by bulk erosion originated from autocatalytic hydrolysis [39,40]. Therefore, it can be concluded that diffusion and erosion of PLCL coating polymers successively drove ciprofloxacin lactate into the surrounding medium. Fig. 4 also presented in vitro elution of ciprofloxacin with 7.5% and 15% loading contents from three different coatings in detail. It was generally observed that the release rate of ciprofloxacin decreased with drug loading increasing in spite of the different coating compositions, which could be explained by drug-drug interactions that may effectively affect and transform the physical-chemical properties of the delivery system [40,41]. Moreover, when the drug loading fixed, it was found that the sequence of the release profile was identical for high and low doses, i.e. PLCL5050 N PLCL6040 N PLCL4060. The fastest elution rate of PLCL5050-CIP at both stages was perfectly explained by its highest weight loss and water absorption, as well as the most crushed morphology during degradation. Nevertheless, CIP accumulated release percentage for PLCL6040 and PLCL4060 at stage I was different from their degradation behavior on the macroscopic scale, i.e. CIP released faster in PLCL6040 than in PLCL4060, while the WL, WA and morphology changed more dramatically in PLCL4060 than in PLCL6040. To reveal the reason for the difference, we further investigated the degradation of PLCL films in artificial urine from the molecular level. As shown in Fig. 5, PLCL5050 exhibited a fastest decrease in Mn followed by PLCL6040 (moderate) and PLCL4060 (slowest), which was the same as the sequence of the drug release. It could be possible that degradation residues with short polymer chains resulted in an increase in chain mobility that promoted the drug diffusion. Hence, it was reasonably

assumed that the release mechanism at stage I might be mainly controlled by chain scission inside a polymer matrix, instead of the weight loss or morphological change of the coatings. As for stage II, the release profile was perfectly agreed with the degradation behavior in both macroscopic and molecular level due to an erosion-controlled release mechanism. Noticeably, PLCL6040 with 7.5% CIP loaded exhibited a higher release percentage than PLCL5050 after 90 days. This complicated profile at stage II might be attributed to the significant collapse of PLCL6040 coating.

3.3. The antibacterial assay In this study, PLCL5050 coating polymer with ciprofloxacin lactate was further used to evaluate the antibacterial profiles of the stents. Herein, the antibacterial effect of 15% CIP-loaded PLCL5050 coating (PLCL5050-CIP) and PLCL5050 coating were investigated by the spread-plate method. The bacteria cultures without any coating or drugs were defined as the blank. The photographs of the colony forming units (CFUs) in agar plates after incubation with the samples for 54 h were displayed in Fig. 6. Numerous colonies of both E. coli and S. aureus either individually or in small tufts of the blanks can be observed on the agar plates. The PLCL5050 without ciprofloxacin showed

Table 2 The quantity of ciprofloxacin for different ureteral stents (n = 6). Abbreviation

CIP load/μg

CIP in coating (w/w)

PLCL6040-L PLCL5050-L PLCL4060-L PLCL6040-H PLCL5050-H PLCL4060-H

639 ± 33 671 ± 15 621 ± 46 1269 ± 65 1220 ± 100 1376 ± 152

7.5% 7.8% 7.6% 15.1% 15.1% 15.3%

Fig. 4. The ciprofloxacin release profiles of different ureteral stents with different ratio of drug-to-polymer.

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Fig. 5. Evolution of number molecular weights of PLCL films degraded in artificial urine.

no significant effect on reducing bacterial cells. On the contrary, there was hardly any viable colony forming units on the agar plates after 54 h of incubation with CIP-loaded PLCL5050 coatings, which can be ascribed to the inhibiting efficiency of ciprofloxacin released from polymer matrix [42,43]. The number of viable bacterial cells of both E. coli and S. aureus after incubation with PLCL5050 and PLCL5050-CIP coatings for 6 h, 18 h, 30 h, and 54 h were displayed in Fig. 7 (E. coli) and Fig. 8 (S. aureus), respectively. Meanwhile, the inhibiting efficiency was calculated from the bacteria counts of PLCL5050-CIP and the blank. It was found that neither E. coli nor S. aureus cell counts showed any significant difference for PLCL5050 samples as compared with the blanks. On the contrary, PLCL5050-CIP coatings significantly inhibited the bacterial growth within 54 h. As plotted in Fig. 7, the number of CFUs dramatically decreased in the culture containing PLCL5050-CIP films, and an inhibiting efficiency of 90.3% was reached at 6 h. After 18 h of incubation, a higher inhibiting efficiency (91.6%) was obtained, although the number of viable bacteria slightly increased to 4.4 × 109 because of the rapid growth

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Fig. 7. The number of viable colony forming units (CFUs) of E. coli in agar plates after 24 h of incubation. Before plating, the E. coil suspension were incubated with the samples for 6 h, 18 h, 30 h, and 54 h, respectively All measurements were made in duplicate; the standard deviations from the mean CFUs values are represented by error bars.

of E. coli. With the ciprofloxacin release, 99.8% of bacteria were inhibited after 30 h, and all the bacteria lost their viability at 54 h. As presented in Fig. 8, the viable bacterial cells of S. aureus also reduced dramatically during the treatment with PLCL5050-CIP samples. A well antibacterial efficiency of 93.5% was obtained at 6 h followed by 97.9% (18 h), 98.7% (30 h) and 99.6% (54 h). Therefore, a sustained antibacterial activity against E. coli and S. aureus was observed in the PLCL5050-CIP samples, which might be due to the controlled release of ciprofloxacin lactate from ciprofloxacin-loaded PLCL5050. 3.4. In vitro cytotoxicity In vitro cytotoxicity of CIP-loaded PLCL5050 coatings were evaluated against HFFs by CCK-8 assay. Fig. 9 showed the HFFs cell viability after incubation with DMEM containing PLCL5050 coating or CIP-loaded PLCL5050 coatings with different drug dose after 24 h. According to CIP release profile as presented in Section 3.2, approximately 20% of ciprofloxacin was released at 24 h and the concentrations of CIP in DMEM

Fig. 6. The images of CFUs of E. coli and S. aureus in agar plates after 24 h of incubation. Prior to spreading, the bacterial cells were incubated with substrates for 54 h.

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controlled by chain scission inside a polymer matrix, instead of the weight loss or morphological change of the coatings. As for stage II, the release profile was dominated by erosion resulting from autocatalysis hydrolysis. Furthermore, ciprofloxacin-loaded PLCL5050 coating showed well antibacterial efficiency on E. coli and S. aureus without obvious cytotoxicity to normal HFFs. These results suggest that PLCL copolymers with tunable degradation rate as carriers for ciprofloxacin could be used as a promising long-term antibacterial coating for ureteral stents. Acknowledgements

Fig. 8. The number of viable colony forming units (CFUs) of S. aureus in agar plates after 24 h of incubation. Before plating, the S. aureus suspension were incubated with the samples for 6 h, 18 h, 30 h, and 54 h, respectively All measurements were made in duplicate; the standard deviations from the mean CFUs values are represented by error bars.

This work was supported by the National Natural Science Foundation of China (51103041, 21274039), Shanghai Scientific and Technological Innovation Project (14520720600), the Fundamental Research Funds for the Central Universities (WD1414007), 111 project (B14018) and Specialized Research Fund for the Doctoral Program of Higher Education (20130074110007). The authors acknowledged the kind help of Dr. Zhaoyang Ye, Zhimiao Xiong for cytotoxicity tests and Changlin Chen, Yunlong Sun for antibacterial experiments. References

were range from 0.028 mg/mL to 0.12 mg/mL. It was found that both PLCL5050 and CIP-loaded PLCL5050 coatings exhibited no significant cytotoxicity to the HFFs even when the drug dose of coatings had increased to 15% as the viability of cells kept almost 100%. The results perfectly agreed with the previous results that CIP was no cytotoxicity to normal HFFs at the concentrations range from 0 mg/mL to 0.389 mg/mL [44,45]. Hence, it could confirm that CIP-loaded PLCL5050 coatings had great potential in biomedical fields. 4. Conclusions In this work, high and low dose loaded ciprofloxacin-eluting ureteral stents with coatings of three different compositions were successfully prepared. Degradation behavior of the PLCL coatings and ciprofloxacin release profile were tuned from fast to slow by altering the ratio of LA and CL in the copolymers. The effect of the drug doses on release profiles was investigated and showed that the release rate of ciprofloxacin decreased with increasing drug loading. All the CIP-eluting ureteral stents exhibited a three-stage sustainable release profile for 120 days, which was greatly affected by degradation behavior of PLCL except for the burst stage. Interestingly, the release mechanism at stage I was mainly

Fig. 9. Cell viability of Human foreskin fibroblasts treated with PLCL5050 and CIP-loaded PLCL5050 coatings.

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