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Biological and Medical Applications of Materials and Interfaces

Thermogel Loaded with Low-Dose Paclitaxel as a Facile Coating to Alleviate Periprosthetic Fibrous Capsule Formation Jiabin Luan, Zheng Zhang, Wenjia Shen, Yipei Chen, Xiaowei Yang, Xiaobin Chen, Lin Yu, Jian Sun, and Jiandong Ding ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b13548 • Publication Date (Web): 13 Aug 2018 Downloaded from http://pubs.acs.org on August 16, 2018

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ACS Applied Materials & Interfaces

Thermogel Loaded with Low-Dose Paclitaxel as a Facile Coating to Alleviate Periprosthetic Fibrous Capsule Formation Jiabin Luan a,1, Zheng Zhang

b,1

, Wenjia Shen a, Yipei Chen a, Xiaowei Yang a, Xiaobin

Chen a, Lin Yu a,*, Jian Sun b,*, Jiandong Ding a a

State Key Laboratory of Molecular Engineering of Polymers, Department of

Macromolecular Science, Fudan University, Shanghai 200438, China b

Department of Breast Surgery, Gynaecology and Obstetrics Hospital, Fudan

University, Shanghai 200011, China

* Corresponding authors. E-mail addresses: [email protected] (L. Yu), [email protected] (J. Sun) 1

Authors contributed equally.

KEYWORDS: thermogel; paclitaxel; fibrous capsule; silicone implant; capsular contracture; sustained drug delivery

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ABSTRACT: Medical-grade silicones as implants have been utilized for decades. However, the post-operative complications, such as capsular formation and contracture have not yet been fully controlled and resolved. The aim of the present study is to elucidate whether the capsular formation can be alleviated by local and sustained delivery of low-dose paclitaxel (PTX) during the critical phase after the insertion of silicone implants. A biocompatible and thermogelling poly(lactic acid-co-glycolic acid)-b-poly(ethylene glycol)-b-poly(lactic acid-co-glycolic acid) (PLGA-PEG-PLGA) triblock copolymer was synthesized by us. The micelles formed by the amphiphilic polymers in water could act as a reservoir for the solubilization of PTX, a very hydrophobic drug. The concentrated polymer aqueous solution containing PTX exhibited a sol-gel transition upon heating and formed a thermogel depot at body temperature. In vitro release tests demonstrated that the entrapped microgram-level PTX displayed a sustained release manner up to 57 days without a significant initial burst effect. Customized silicone implants coated with the PTX-loaded thermogels at various drug concentrations were inserted into the pockets of the subpanniculus carnosus plane of rats. The histological observations performed one month post-operation showed that the sustained release of PTX with an appropriate dose significantly reduced the peri-implant capsule thickness, production and deposition of collagen, and expression of contracture-mediating factors compared with bare silicone implants. More importantly, such an optimum dose had an excellent repeatability for the suppression of the capsular formation. Therefore, this study provides a strategic foothold regarding the sustained release of low-dose PTX to alleviate fibrotic capsule formation after implantation and the microgram-level PTX-loaded thermogel holds great potential as an “all-purpose anti-fibrosis coating” for veiling the surfaces of various implantable medical devices.


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1. INTRODUCTION Over the past decades, the use of implanted medical devices, such as scaffolds, stents, artificial joints, heart valve prostheses and intraocular lenses, for treatment of a variety of diseases has made significant progress to address the needs of an aging population.1-4 These devices remarkably improved the life quality of millions of patients worldwide. The formation of a scar-like capsule isolating medical devices from adjacent tissues is a natural healing response to the presence of foreign objects in vivo, but often results in excessive fibrosis scarring, which can cause various complications that adversely impart or compromise the performance of implants, and even lead to a secondary surgery.5-7 This vexing problem is still not sufficiently controlled and resolved so far. Silicone implants are the most widely used medical devices in the field of cosmetic and reconstructive surgery, especially for breast augmentation and reconstructive surgery.8-9 Capsular contracture that originates from the formation of serious pathological fibrous capsule around implants is the most prominent and specific complication, which causes patients a great deal of pain and distorts the appearance of their breasts. The incidence rate may be as high as 30% for patients receiving augmentation mammoplasty or reconstructive breast surgery.10-12 In serious cases, a secondary surgery is required, which highly impairs the wellbeing of patients.10, 12 Although the mechanism of the formation of capsular contracture has not yet been elucidated in detail, it is generally acknowledged that the level of fibroblast proliferation and the collagen synthesized by fibroblasts around the silicone implant play a prominent role in capsule formation,12-13 in addition to clinical factors such as infection by normal skin flora (usually by Staphylococcus epidermidis), radiotherapy, etc.11, 14



Paclitaxel (PTX) is a famous chemotherapy drug that disrupts the normal microtubule dynamics by stabilizing the microtubule and increasing microtubule polymerization, and eventually induces cell apoptosis.15 It has been extensively used in the treatment of various types of cancers. Interestingly, recent studies indicate that low-dose PTX has good inhibitory effects on the synthesis of collagen, growth and proliferation of fibroblasts, and neovascularization,16-17 and has been tried to treat collagen-induced arthritis, multi-organs fibrosis including liver, lung and kidney, and fibrosis-related systemic sclerosis.18-22 We thus hypothesized that low-dose PTX may be therapeutic for hypertrophic fibrosis scarring around implantable medical devices. No pertinent report has been found, and the present study is aimed to check our hypothesis in combination with the unique thermogelling materials. Compared with systemic administration, localized drug delivery may significantly enhance therapeutic efficacy while minimizing systemic toxicity. In situ forming injectable hydrogels are very attractive as localized drug carriers because of their superior advantages including facile fabrication, target administration, and a sustained drug release manner.23-28 Injectable thermogels are a type of hydrogels that are free-flowing polymer solutions at low or room temperature, but turn into semi-solid hydrogels at body temperature.29-39 Therefore, drugs can be easily entrapped into the polymer solutions by simply mixing them at a low temperature, while drug-loaded thermogel depots can be spontaneously formed upon injection into the target sites. Among them, biodegradable and biocompatible thermogels based on poly(lactic acid-co-glycolic acid)-b-poly(ethylene glycol)-b-poly(lactic acid-co-glycolic acid) (PLGA-PEG-PLGA) triblock copolymers are one of the most popular systems owing to the good profile of biocompatibility and the long persistence in the gel form in vivo, and have been employed to deliver a variety of drugs to treat different diseases, such


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as cancer, osteopenia, etc.30, 40-44 In the present work, we suggest using the local, sustained release of low-dose PTX around silicone implants to mitigate capsular formation and/or contracture, as schematically illustrated in Figure 1. Firstly, a thermogelling triblock copolymer, PLGA-PEG-PLGA,

was

synthesized

and

the

sol-gel

transition

of

the

PLGA-PEG-PLGA aqueous solutions with or without PTX was studied. Subsequently, we examined the in vitro drug release behaviors of the thermogel formulations containing microgram-level PTX. Silicone implants coated with microgram-level PTX-loaded thermogels were inserted into the subpanniculus plane of rats, and the resulting fibrosis capsules were carefully compared with those obtained from bare silicone implants. Various quantitative analysis comparing capsular thickness, inflammatory cellularity, vascularity and amounts of transforming growth factor-β (TGF-β), α-smooth muscle actin (α-SMA) and CD68 were performed to access the effects of the sustained release of PTX with varied drug doses on peri-implant capsular formation. The optimum drug dose window was determined and its repeatability was further validated. This study provides a new way to reduce the excessive fibrosis scarring after implantation, including capsular contracture through the sustained release of low-dose PTX at the local insertion site.



Figure 1. Schematic illustration of the function of a microgram-level PTX-loaded thermogel coating on silicone implants.

2. MATERIALS AND METHODS 2.1. Materials. Poly(ethylene glycol) (PEG) with molecular weight (MW) ~1500 and molar mass distribution (ĐM) ~1.13, and stannous octoate (Sn(Oct)2, 95%) were purchased from Sigma-Aldrich. Monomers

D,L-lactide

(LA) and glycolide (GA) were

products of Jinan Daigang Biomaterial Co., Ltd (China). Paclitaxel (PTX, 99.8%) was obtained from Shanghai Meishidun Biotechnology Co., Ltd (China). Miniature silicone breast implants identical to implants used in clinical breast surgery in terms of materials and texture (textured type) with average weight of 1.38 ± 0.10 g, diameter of 15 mm, and height of 10 mm, were customized from Shanghai Winner Plastic Surgery Products Co., Ltd (China). Clinical Taxol solution (each mL of solution contains 6 mg PTX, which is solubilized by Cremophor® EL ACS Paragon Plus Environment

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(polyoxyethylated castor oil)) was supplied by the Obstetrics and Gynecology Hospital of Fudan University and produced by Hainan Haiyao Co., Ltd (China). All other reagents and materials were used as received. 2.2. Animals. Female Sprague-Dawley (SD) rats (average weight: 220-250 g) were purchased and housed at Shanghai Laboratory, Animal Research Center (China). The animals were raised in an environment of 20-24 °C, relative humidity of 30−70% and 12 h light/dark cycle with free excess to standard rat chow and water ad libitum. All animal experiments were performed under the approval of the ethics committee of Fudan University.

2.3. Polymer Synthesis and Characterization. The triblock copolymer PLGA-PEG-PLGA was synthesized by the ring-opening polymerization of LA and GA initiated by the hydroxy end group of PEG and catalyzed via Sn(Oct)2.45 Briefly, 15 g PEG was first added into a three-neck flask and then dehydrated under vacuum at 120 ºC for 3 h. Next, a given amount of LA, GA (molar ratio 3/1) and catalyst Sn(Oct)2 were introduced to the reaction device at 90 ºC under the protection of argon and then dehydrated for another 30 min. Afterwards, the temperature of device was raised to 150 ºC for 12 h with continuous stirring. After the completion of reaction, the crude products were purified by washing in 80 ºC hot water for three times and then lyophilized to eliminate the residual water. Finally, the products were collected and stored at -20 ºC until use. 1

H NMR was applied for structure and MW estimation of the polymer on a 400

MHz NMR spectrometer (AVANCE III HD 400MHz, Bruker) using CDCl3 as the solvent and tetramethylsilane (TMS) as the internal reference. The MW and ĐM of the polymer were further determined by gel permeation chromatography (GPC, Agilent



1260) equipped with a refractive index detector. The measurement was performed at 35 ºC using tetrahydrofuran (THF) as the eluting solvent at a flow rate of 1.0 mL/min, and monodispersed polystyrene was used as the calibration standard of MW. The sol-gel transitions of 25 wt% PLGA-PEG-PLGA triblock copolymers in normal saline (NS) with or without drugs were monitored using a dynamic stress-controlled rheometer (Malvern, Kinexus) equipped with a cone-plate (diameter: 60 mm, cone angle: 1º). Storage modulus G' and loss modulus G'' were recorded with an oscillatory frequency ω of 10 rad/s and a heating rate of 0.5 ºC/min from 15 to 50 ºC. The morphology of micelles was observed with a transmission electron microscope (TEM, Tecnai G2 20 TWIN, FEI) with an accelerating voltage of 200 kV. Samples were prepared by placing one drop of polymer solution (1 wt%) on a copper grid coated with a superthin carbon film, dried under an infrared lamp and observed without staining.

2.4. In Vitro PTX Release. Polymer aqueous solutions (25 wt% in NS) containing 80, 40 and 20 µg/mL PTX were prepared by directly mixing the polymer solution with the drug and stirring them at 4 ºC. One milliliter of PTX-loaded PLGA-PEG-PLGA copolymer aqueous solution was injected into a 15-mL vial (outer diameter: 22 mm) and then incubated in a shaking bath (37 ºC, 50 rpm) for 15 min to form a stable physical hydrogel containing PTX. Next, 10 mL pre-warmed phosphate buffer saline (PBS, pH 7.4) solution containing 0.025 wt% sodium azide and 0.5 wt% Tween-80 was added as the release medium. Half of the release medium was replaced with the same amount of fresh medium every three days to maintain the sink condition. At specific time points, after decanting the release media in vials (n = 3 for each time point), the remaining hydrogels containing PTX were collected and then


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lyophilized. The residual samples were re-dissolved in acetonitrile (2 mL per sample). After centrifugation, the supernatants were analyzed by high performance liquid chromatography system (HPLC, Waters e2695) equipped with a SunFire™ C18 reverse-phase column (4.6 × 150 mm, 5 µm) and a UV-Vis detector (Waters 2489). The mobile phase consisted of a mixture of water and acetonitrile in 50:50 (v/v) ratio, and was delivered in the isocratic elution mode at a flow rate of 1.0 mL/min. UV-Vis detection at 227 nm was used for the quantification of PTX. The accumulated release amount was obtained based on a PTX standard calibration curve prepared with acetonitrile (R2 > 0.9999) as shown in Figure S1. The release profiles of the PTX-loaded hydrogel formulations (80, 40 and 20 µg/mL) were further fitted via zero-order equation / = .46 Here, means the cumulative release amount at the time of t, is the final release amount at infinite time and k is a constant.

2.5.

In

Vivo

Degradation

of

Thermogel.

In

situ

formation

of

the

PLGA-PEG-PLGA thermogel and subsequent in vivo degradation were examined in a SD rat model. Animals were anesthetized by intraperitoneal administration of chloral hydrate (350 mg/kg). Next, PLGA-PEG-PLGA aqueous solution (25 wt% in NS, 0.4 mL), which was sterilized by irradiation of Co-60, was subcutaneously injected into the neck of rats. Animals were sacrificed at scheduled intervals. The remaining hydrogels at the injection sites were then dissected and recorded using a digital camera.

2.6. Implantation Surgery Procedures. The SD rats were randomly divided into five groups (n = 10 per group): control group (injection of 0.9% normal saline, 100



µL), Blank Gel (injection of polymer solution without drugs, 100 µL), PTX 20, 40 and 80 (injection of 20, 40 or 80 µg/mL PTX-loaded polymer solutions, 100 µL). Silicone breast implants were sterilized by immersing in betadine solution, and then rinsed with ethanol and NS successively. After anesthetization by intraperitoneal injection of chloral hydrate (350 mg/kg), backs of the rats were shaved, cleaned and sterilized by betadine solution. Surgery procedures were conducted under aseptic conditions. Two dorsal symmetrical 1-cm incisions were made perpendicularly to the longitudinal axis in the lumbar-sacral region, and subcutaneous pockets beneath the panniculus carnosus muscle were constructed by sharp and blunt dissection. Next, a total amount of 100 µL polymer solutions with or without drugs were injected into the pockets before and after implantation of silicone breast implants to uniformly coat the whole surface of implants. Each animal received two individual implants belonging to the same group. Finally, muscle and skin incisions were closed with 4-0 Vicryl Plus™ sutures (Ethicon, Inc., USA). The detailed experimental procedures were illustrated in Figure S2. Rats were housed in individual cages to avoid biting wound sites each other. General and wound checkups were performed regularly throughout the experiments. One month later, the animals were sacrificed and the implants with the surrounding tissues were harvested. Samples were fixed in 4 % paraformaldehyde for histological and immunohistochemical analysis as described below. To prove the test-retest reliability, investigations of PTX 40 group versus control group (n = 8 for each group) were carried out again at a different time under the same conditions. An independent control experiment was also performed under the same conditions and the SD rats were randomly divided into four groups (n = 10 per group): control


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group (injection of 0.9% normal saline, 100 µL), Taxol 20, 40 and 80 (injection of 20, 40 or 80 µg/mL PTX solutions diluted from the Taxol solution, 100 µL).

2.7. Micro-Computed Tomography (Micro-CT). Micro-CT (Bruker, Skyscan 1176) was applied to monitor the shape and contour of the implants at 0, 14 and 30 d after implantation. Mice were anesthetized by intraperitoneal injection of chloral hydrate. The scanning parameters were set as follows: spatial resolution of 18 µm pixel size; X-ray tube voltage of 80 kV; X-ray tube current of 300 µA; Al + Cu filter; 0.50° rotation step 360°.

2.8. Histological and Immunohistochemical Analysis. Harvested specimens were embedded in paraffin and sectioned for hematoxylin and eosin (H&E) and Masson’s trichrome staining. Each stained slide was photographed under an inverted fluorescence microscope (Axiovert 200, Zeiss) equipped with a CCD (AxioCam HRC, Zeiss) for later processing, and this process was performed under the circumstance of no access to the sample information. The capsule thickness around implants was measured according to the previous publication with modifications.47 Briefly, the thicknesses of 10 sites which were designated every 150 µm were measured using the software Image J (NIH). Capsule thickness for each specimen was obtained by calculating the average of the thickness. Collagen deposition was observed and compared between groups from Masson’s trichrome staining. Immunohistochemical staining was performed using anti-TGF-β (Abcam, UK), anti-α-SMA (Abcam, UK) and anti-CD68 (Santa Cruz Biotechnology, USA). Stained sections were visualized using a Zeiss microscope (Axiovert 200) and imaged with a CCD camera at 200× magnification (AxioCam HRC, Zeiss) with the same imaging



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parameters. Again, all images were obtained with the sample information masked from the operator. At each biopsy time and for each implant sample, three images were randomly obtained and analyzed from each of the ten implants. The densities of inflammatory cells and blood vasculature were quantified to measure the tissue response to the implant. Specifically, the number of inflammatory cells in the capsule of each image was calculated by Image-Pro Plus 6.0 (NIH) and data were presented as number per unit area (0.01 mm2). The number of blood vessels was counted manually for each image and expressed as a vessel number per unit area (1 mm2). For immunohistochemical staining analysis, the total pixel intensity of the formed capsule was measured using Image-Pro Plus 6.0 (NIH) and data were expressed as optical densities.

2.9. Statistical Analysis. Data analysis was performed with SPSS 16.0 (IBM Corp., USA). Results were expressed as means ± standard deviation (SD) or box-whisker plot. For data which were normally distributed with equal variances, one-way ANOVA with Tukey’s multiple comparison test was used.48 For the comparison between the two groups of the test-retest reliability, an unpaired t-test was used.13 A p value < 0.05 was considered statistically significant.

3. RESULTS 3.1.

Synthesis

and

Characterization

of

PLGA-PEG-PLGA

Triblock

Copolymer. The PLGA-PEG-PLGA triblock copolymer was synthesized via the ring-opening copolymerization of LA and GA in the presence of PEG as the macro-initiator and Sn(Oct)2 as the catalyst, as presented in Scheme S1. The resultant specimen was characterized by

1

H NMR and GPC measurements. All the


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characteristic proton signal peaks of copolymer were well assigned in the 1H NMR spectrum, as shown in Figure S3. The number-average MW (Mn) and LA/GA molar ratio of PLGA-PEG-PLGA triblock copolymer were calculated based on the characteristic peaks at 3.65, 4.80 and 5.20 ppm.43 Meanwhile, the GPC trace of PLGA-PEG-PLGA triblock copolymer exhibited a unimodal manner (Figure S4) with a ĐM of 1.26, indicating the successful synthesis of the desired polymer. The results of 1

H NMR and GPC measurements are listed in Table 1.

Table 1. Parameters of PLGA-PEG-PLGA Triblock Copolymer Investigated in This Study

Sample

PLGA-PEG-PLGA a

Mn (g mol-1) a

Mn (g mol-1) b

ĐM b

5490

1.26

1860-1500-1860

LA/GA (mol/mol) a 3.1

Mn of PEG was provided by Aldrich. Mn of each PLGA block and the molar ratio of

LA/GA were calculated via 1H NMR. b Determined by GPC.

3.2. Sol-Gel Transition of PLGA-PEG-PLGA Polymer Aqueous Solutions with or without Drugs. The PLGA-PEG-PLGA triblock copolymers were soluble in NS, and the obtained polymer aqueous solution underwent a sol-gel transition as the temperature increased. Notably, the polymer aqueous solution turned into a semi-solid gel at body temperature (37 ºC), which is important for its in vivo biomedical application. The sol-gel transition was well maintained for the hydrogel formulation containing 40 µg/mL PTX, as shown in Figure S5A. Dynamic rheology analysis was further used to monitor the change of storage modulus G’ and loss modulus G’’ of polymer aqueous solutions with or without drugs



as a function of temperature. As presented in Figure S5B, at low and room temperatures, the moduli were low and the G’’ was greater than the G’, indicating the sol state of the systems. As the temperature elevated, both G’ and G’’ increased rapidly and the G’ value increased more than 5 orders of magnitude. Generally, the sol-gel transition temperature (Tgel) is defined as G’ equals to G’’.49 The Tgel of polymer aqueous solution obtained from the rheological measurement was 33 ºC. It is apparent that the introduction of microgram-level PTX into the polymer/NS system had no obvious effects on its injectability and thermogelling behavior.

3.3. In Vitro PTX Release from PLGA-PEG-PLGA Hydrogels. Paclitaxel is a very hydrophobic drug with a solubility about 4 µg/mL in water.50 The injectable PLGA-PEG-PLGA thermogel has been proved to be an ideal carrier for PTX with a solubility enhancement of over 2000-fold due to the corona-core micellar structure formed by the amphiphilic copolymers in water.50 Besides, the stability of PTX in thermogel formulation was substantially improved,50 which is important for the maintenance of drug efficacy in a sustained release manner. For the microgram-level PTX-loaded thermogel formulations (PTX 20, 40 and 80 µg/mL), due to the ultra-low drug concentration in the release medium, it is not available to correctly detect the amount of drug released in the medium. Therefore, the drug contents in the remaining hydrogels were examined to determine the in vitro release profiles of PTX. As shown in Figure 2, the loaded PTX on the microgram-level exhibited a sustained release manner up to 57 days. Not only was no initial burst release observed, but also the cumulative release amount was more than 85%. Meanwhile, the release profiles of PTX matched well with the zero-order equation and the fitting results are summarized in Table 2. This feature indicated that the drug release kinetics were


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controlled by the thermogel degradation, regardless of the PTX loading amounts in the formulations.

Figure 2. In vitro release profiles of PTX from PLGA-PEG-PLGA thermogels in PBS at 37 ºC. Polymer concentration was 25 wt%. The results are presented as mean ± standard deviation (n = 3).

Table 2. Fitting Parameters of the PTX Release Data In Vitro Zero-order model Sample

k·M∞

R2

PTX 20 µg/mL

1.71

0.979

PTX 40 µg/mL

1.62

0.976

PTX 80 µg/mL

1.67

0.984

The morphology change of PLGA-PEG-PLGA thermogels as a function of drug release was also monitored throughout the in vitro release experiment. As shown in Figure 3, the PTX-loaded hydrogels remained intact for over 3 weeks. While the



hydrogel was degraded via hydrolysis of PLGA blocks40, 50-51 and the volume of hydrogel gradually swelled, a clear interface between the release medium and hydrogel was well maintained at this stage. Subsequently, tiny hydrogel fragments continuously broke away from the gel matrix and entered into the release medium due to the continuous degradation of hydrogel, and thus the volume of residual gel gradually decreased. At the end of the experiment (57 d), only a few portion of remaining hydrogels were observed at the bottom of the vials.

Figure 3. Optical images of the PTX-loaded PLGA-PEG-PLGA thermogels at the indicated time points in PBS at 37 ºC. Initial polymer concentration was 25 wt %. The dash lines are used to indicate the initial interface between the release medium and thermogel.

3.4. In Vivo Degradation of PLGA-PEG-PLGA Hydrogel. In situ gel formation and in vivo degradation of the PLGA-PEG-PLGA thermogel were confirmed and ACS Paragon Plus Environment

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tracked in SD rats. The copolymer aqueous solutions rapidly transformed into in situ thermogels with round or irregular shape protrusion after subcutaneous injection into the neck of animals. As displayed in Figure 4, the sizes of hydrogels gradually decreased as a function of degradation time due to the hydrolysis of ester bonds of PLGA segments.40, 50-51 The final degradation products are lactic acid, glycolic acid and non-degradable PEG, which are easily absorbed and metabolized in the body.40, 50-51 The in vivo integrity of gel maintained over 3 weeks and no visible residual hydrogel was detected at day 30 post-injection. Therefore, the in vivo persistence of the current PLGA-PEG-PLGA hydrogel was determined to be around one month.

Figure 4. In vivo gel degradation. Optical images were taken at the indicated time points after subcutaneous injection of the PLGA-PEG-PLGA polymer aqueous solutions (25 wt% in NS, 0.4 mL) into the neck of SD rats. Dashed ellipses were used to indicate the presence of gel.

Furthermore, no pathological symptoms such as tissue necrosis, hemorrhaging, hyperemia, edema, or muscle damage were observed during the whole in vivo



experiment, suggesting the good biocompatibility of the PLGA-PEG-PLGA thermogel, which will be further supported by the subsequent histological results. It is worth pointing out that the in vivo maintenance of the PLGA-PEG-PLGA thermogel was shorter than that in vitro. This was attributed to the complicated physiological environment in the body, such as invasion of cells, enzymatic hydrolysis of ester bonds, etc. that accelerated the in vivo degradation of the PLGA-PEG-PLGA thermogel.40, 51

3.5. In Vivo Capsular Formation. To study the effect of the PTX-loaded hydrogel formulations on the formation of capsules around the implants, customized silicone implants whose surfaces were uniformly coated with a layer of hydrogel with or without PTX were inserted into the subpanniculus carnosus plane of rats. All the subject animals survived during the entire examination period. The animals were sacrificed 30 d after implantation. The tissues around the silicone implants were carefully harvested, embedded and stained for further analysis.

3.5.1. In Vivo PTX Dosage-Dependence Study. To begin with, a preliminary in vivo PTX dosage-dependence study was carried out to determine the proper PTX loading amount in the thermogel formulations. Considering the loading amount of PTX in the thermogel formulation (OncoGelTM) for the treatment of cancer generally reached milligram level per mL,38, 50 the PTX dosage was reduced by one order of magnitude (PTX 0.1, 0.2 and 0.4 mg/mL) in the current study. As shown in Figure 5, the fibrous capsules at the material-tissue interface have been formed after the introduction of customized silicone implants for one month, regardless of the treatment of PTX-loaded thermogel formulations. However, severe inflammatory responses were observed in the fibrous capsule tissues after the treatment


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of the three thermogel formulations with varied drug dosages and the degree of inflammatory response increased with the increase of PTX amount in the thermogel system. The thickness of fibrous capsules of rats receiving the treatment of 0.1 mg/mL PTX-loaded thermogel formulation was comparable with that of NS group, yet the fibrous capsule contained a large amount of inflammatory cells (mostly lymphocytes) and fibroblasts. As the PTX concentration increased to 0.2 mg/mL, intense inflammatory response in the capsule tissue resulted in a significantly increased thickness of fibrous capsule. As for the animals treated with the thermogel formulation containing 0.4 mg/mL PTX, inflammatory cells even infiltrated into the adjacent muscle layer and the skin necrosis around the implantation sites was clearly observed during the experiment, indicating the presence of severe tissue toxicity. The preliminary PTX dosage-dependence study revealed that the PTX concentrations ranging from 0.1 to 0.4 mg/mL were still excessive, which caused obvious toxicity against the surrounding subcutaneous tissue. Therefore, the PTX loading amount was further reduced by another order of magnitude in the subsequent experiments.

Figure 5. In vivo PTX dosage-dependence study. Representative histological images (40× and 100× magnification) using H&E staining of the capsule biopsy specimens taken from NS (the control) and the thermogels containing 0.1, 0.2 and 0.4 mg/mL PTX after 30-day implantation in rats. Asterisks (*) mark the site of the silicone



implants and white bars denote the fibrous capsule.

3.5.2 In Vivo Microgram-Loaded PTX Study. The effects of microgram-level PTX-loaded thermogel formulations on the formation of fibrous capsule were further evaluated. Figure 6 shows representative histological images of sections after the implantation of customized silicone implants combined with the different treatments for one month. In the case of NS group, the capsules were rich in fibroblasts, histiocytes and vascular endothelial cells. Meanwhile, the deposited collagen fibers in the capsule were dense and irregularly arranged. In contrast, the collagen fibers that formed in the blank thermogel (Gel) group were loose and showed a relatively parallel arrangement. The structure of fibrous capsule in the PTX 20 group was similar to that of Gel group. Interestingly, as the concentration of PTX increased to 40 µg/mL, the thickness of fibrous capsule decreased notably and the parallelly aligned collagen fibers were highly loose. However, with further increasing the concentration of PTX to 80 µg/mL, both the significantly increased inflammatory cellularity and thickness of fibrous capsule were found compared with that of PTX 40 group.


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Figure 6. Representative histological images (100× and 200× magnification) using H&E and Masson’s trichrome staining of the capsule biopsy specimens taken from NS (the control), blank thermogel (Gel) and the thermogels containing 20, 40 and 80 µg/mL PTX (PTX 20, 40 and 80) after 30-day implantation in rats. Asterisks (*) mark the site of the silicone implants and white bars denote the fibrous capsule.

3.6. Capsule Thickness. The differences in the fibrous capsules that formed in different groups were further analyzed quantitatively. As shown in Figure 7A, the thickness of fibrous capsule was measured every 150 µm and the average value of ten sites was defined as the thickness of the tested capsule.47 In consistency with the



morphological observations shown in Figure 6, the PTX 40 group exhibited the thinnest capsule thickness among the five groups (Figure 7B). Compared with the capsule formed in the PTX 40 group, the average capsule thickness was increased by 62% in the NS group, by 21% in the Gel group, by 33% in the PTX 20 group and by 66% in the PTX 80 group. Meanwhile, significantly statistic differences were observed between the PTX 40 group and the NS group or the PTX 80 group. This finding indicated that the administration of 40 µg/mL PTX-loaded thermogel formulation and the subsequent sustained release of drug remarkably alleviated the thickness of formed capsule after the implantation of silicone implants. In order to further verify the repeatability or test-retest reliability, the effect of the most effective PTX concentration (40 µg/mL) versus NS on the capsule thickness was carried out again under the same experiment conditions. Figure 7C shows that the average capsule thickness in the PTX 40 group was 167 µm, while that in the NS group increased by 36% and reached 227 µm. Also, there was a significantly statistic difference between the two groups. Therefore, the good repeatability of the efficacy of the optimized PTX-loaded thermogel formulation for the alleviation of capsule formation was validated. Also, an independent control experiment was performed by the direct injection of PTX solutions at the same low doses (20, 40, and 80 µg/mL) around the silicone implants to evaluate whether the treatment effects were related to long-acting release or fast release of drug. As presented in Figure S6, the short-acting administration of PTX solutions at the insertion site did not mitigate the capsule formation. Conversely, a significant increase in capsule thickness (19%, 33% and 53% higher than that of the control group in average for Taxol 20, 40 and 80 group, respectively) and inflammatory response were observed as the injection dose increased, indicating that


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the fast release of PTX around the silicone implants caused tissue toxicity and induced excessive fibrosis even if at the low doses. Obviously, this finding further confirmed that the alleviation of capsule formation in the PTX 40 group was attributed to the sustained release of proper dose of drug at the insertion site.



Figure 7. (A) The thicknesses of ten sites every 150 µm were measured and the thickness of the capsule was obtained by calculating the average. 100× magnification. (B) Box-whisker plot of the fibrous capsule thickness around the implants (n = 10): minimum, 25th percentile, median (solid line), mean (dash line), 75th percentile and maximum. ANOVA *p < 0.05 versus NS, #p < 0.05 versus PTX 80. (C) Box-whisker plot of the fibrous capsule thickness around the implants in a test-retest reliability (n = 8): minimum, 25th percentile, median (solid line), mean (dash line), 75th percentile and maximum. *p < 0.05 versus NS.

3.7. Cellularity and Vascularity. Inflammatory cells act as a key indicator in inflammatory reactions, which secrete different cytokines, recruit fibroblasts and activate collagen synthesis, thus result in capsule formation.13 The density of intracapsular inflammatory cells (mostly lymphocytes with a small amount of monocytes and macrophages) was evaluated and presented in Figure 8A. The PTX 40 group showed the narrowest data distribution with the smallest data of median (6.8 counts per 0.01 mm2) and average (7.6 counts per 0.01 mm2) among the testing groups. The declining trend of inflammatory cell counts was consistent with the capsule thickness results. We also compared the vascularity, an indicator of neoangiogenesis, of the fibrous capsules around the implants (Figure S7). Compared with the other four groups, the PTX 40 group exhibited an overall trend of lower vascularity (Figure 8B). Meanwhile, we observed a significant difference between the PTX 40 group and the NS group.


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Figure 8. In vivo analysis of intracapsular inflammatory cell and vascular densities. Box-whisker plot of (A) inflammatory cell density (n = 10) and (B) vascular density (n = 10): minimum, 25th percentile, median (solid line), mean (dash line), 75th percentile and maximum. ANOVA *p < 0.05 versus NS.

3.8. In Vivo Micro-CT. To track the possible changes in shape and contour of implants caused by capsular contracture, the implants were scanned by Micro-CT at designed intervals and then the three-dimensional morphologies of implants were reconstructed. As shown in Figure S8, there was no obvious change both in the shape and contour of implants and their transverse sections during the one-month experiment. The reason of this negative result might be that the capsules formed



around the implants within one month are not sufficient to cause detectable compression by Micro-CT. Though no change in shape of implants was observed, a non-invasive detection method was successfully applied by us to track the contour of implants, which might have some reference for other researchers.

3.9. Immunohistochemical Analysis of Capsular Formation. We further performed immunohistochemical analysis to obtain more information on the capsular formation around the silicone implants. The level of TGF-β expression is a major cytokine secreted by inflammatory cells and fibroblasts. It has been regarded as a key mediator in the progression of fibrosis,12, 21, 52-53 which could promote the recruitment and proliferation of fibroblasts, activate fibroblasts to synthesize collagen and down-regulate matrix metalloproteinases.13,

53-55

One-month after implantation, the

optical density of TGF-β in the PTX 40 group was significantly lower than those in the NS, PTX 20 and PTX 80 groups (Figure 9A, D). We expected that the low level of TGF-β down-regulated the inflammatory response and collagen synthesis, thus markedly alleviated the formation of periprosthetic fibrous capsule. α-SMA, as a sign of the formation of myofibroblasts, is responsible for the increased collagen production, extracellular matrix secretion and fibrotic lesions.56 Although no significant difference was observed among the testing groups, the optical density of PTX 40 group showed a lower data distribution than the NS, Gel and PTX 80 groups (Figure 9B, E). The low titer of α-SMA in the PTX 40 group would contribute to the decrease of collagen deposition (as shown in Masson’s trichrome staining images) and the suppression of capsular formation. CD68, a typical inflammatory cell marker, was also related with the severity of fibrosis.57 There was no significant difference among the testing groups. However, it is


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still worth pointing out that the PTX 40 group presented a relatively low CD68 expression compared with the other four groups (Figure 9C, F). This finding was consistent with the results of the inflammatory cell counts in Figure 8A.

Figure 9. Immunohistochemical analysis of the fibrous capsules around the implants for the indicated groups. Representative images of (A) TGF-β, (B) α-SMA and (C) CD68 immunostaining in the fibrous capsules. Asterisks (*) mark the site of the silicone implants. Box-whisker plot of the optical densities of the expressed (D) TGF-β (n = 10), (E) α-SMA (n = 10) and (F) CD68 (n = 10) amounts: minimum, 25th percentile, median (solid line), mean (dash line), 75th percentile and maximum. ANOVA *p < 0.05 versus NS, #p < 0.05 versus PTX 20, †p < 0.05 versus PTX 80.



4. DISCUSSION Capsular contracture remains the most troublesome post-operative complication after the breast implant insertion.10 Although the underlying mechanism of capsular contracture is still controversial, excessive foreign body reactions have been acknowledged as one of the key factors to cause the capsular contracture.13 Typically, the foreign body reactions involve the inflammation process from the acute to chronic stages, collagenous accumulation, and eventual formation of fibrous tissue encapsulating the implant within 3 weeks post-surgery.13 Surface modifications have been tried to alleviate the excessive foreign body reactions around the silicone implants,13, 57-58 however, only limited success has been achieved in clinical studies. Recently, researchers have shown that low-dose PTX is able to alter microtubule dynamics without causing toxicity and thus offers a potentially appealing option to treat various diseases associated with excessive fibrosis and inflammation.59 For example, Lv et al. found that the pulmonary fibrosis in a bleomycin-treated rat model was effectively ameliorated by intravenous injection of low-dose PTX (0.6 mg/kg daily for two weeks) to suppress TGF-β1/Smad3 pathway.20 Liu et al. reported that the intraperitoneal injection of low-dose PTX (0.3 mg/kg, twice a week) significantly suppressed the tubulointerstitial fibrosis in a rat model of unilateral ureteral obstruction by inhibition of TGF-β/Smad activity.21 Therefore, we hypothesized that local and sustained release of low-dose PTX may be a new therapeutic strategy to mitigate the capsular formation and/or contracture around silicone implants. Meanwhile, compared with the systemic administration, such as intravenous injection and intraperitoneal injection, local drug delivery can enhance the therapeutic potency, while minimize the drug-associated toxicities. To the best of our knowledge, no such long-acting delivery systems of low-dose PTX have ever been reported.


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In the current study, we developed an injectable hydrogel consisting of thermogelling and biodegradable PLGA-PEG-PLGA triblock copolymers to long-termly deliver low-dose PTX around the silicone implants. As presented in Figure S9, the amphiphilic carrier polymers easily self-assembled into corona-core micelles in water. The concentrated polymer aqueous solution exhibited a macroscopic sol-gel transition as the temperature increased (Figure S5), and such a thermogelation process was attributed to the formation of a percolated micelle network via micellar aggregation.29, 60 Due to the solubilization effect of polymer micelles, the hydrophobic PTX was conveniently encapsulated into the polymer aqueous solution simply by physical stirring at a low temperature. Meanwhile, the introduction of PTX did not obviously affect the injectability and temperature-induced sol-gel transition of PLGA-PEG-PLGA thermogel (Figure S5). The

previous

study has

revealed

that the

milligram-level

PTX-loaded

PLGA-PEG-PLGA thermogel (ReGel®) exhibited a sustained release manner of drug for approximately 50 days.50 Such a feature was also verified by us, as demonstrated in Figure S10. Interestingly, as the drug loading amount decreased to microgram scale, the release of PTX from the thermogel matrix likewise lasted up to 57 days and almost showed a constant rate throughout the whole examined period (Figure 2). This finding indicated that the carrier polymer degradation-controlled mechanism governed the PTX release regardless of the drug loading amounts. After insertion of the silicone implants into the subpanniculus carnosus plane of rats, the microgram-level PTX-loaded polymer solution was instilled using a conventional syringe around the implants. Due to the contact with body heat, the polymer solution rapidly transformed into a thermogel depot containing PTX to veil the irregular surface of the implant. Subsequently, low-dose PTX was sustainedly



released along with the gradual degradation of thermogel matrix. As is well known, the first month after implant insertion constitutes the most sensitive period of foreign body responses.13,

57

In this study, we carried out

histological analysis of the tissues around the implants one month after operation. In general, capsule thickness is positively associated with the occurrence of capsular contracture.58 Our results revealed that the thickness of capsule and inflammation level were dramatically dependent on the dose of PTX. When the drug loading amount in the thermogel matrix was greater than 80 µg/mL, the sustained release of PTX resulted in a remarkable enhancement of inflammation response in the capsule tissues (Figure 5). Further increasing the drug loading amount to 200 µg/mL, intense inflammation response even led to a significant increase in capsule thickness. This feature suggested that excessive PTX did not mitigate capsule formation, but caused serious side effects. In fact, previous studies have demonstrated that the dose of PTX plays a key role in maintaining the balance between the activity of fibrosis inhibition and the toxicity against normal tissues, and high-dose PTX induces inflammation and fibrosis.59, 61 Interestingly, coating 40 µg/mL PTX-loaded thermogels on the surface of silicone implants notably reduced the capsule thickness, collagen density and inflammatory cellularity (Figures 6-8). Nevertheless, a relatively low or relatively high loading amount of PTX (20 µg/mL or 80 µg/mL) compromised the therapeutic efficacy so that no significant difference was detected compared with bare implants (Figures 6-9). Although the therapeutic window of PTX was relatively narrow, the repeatability test affirmed that the efficacy of the optimized PTX-loaded thermogel formulation for the alleviation of capsule formation was reliable and reproducible (Figure 7C). The good repeatability of the current study provides the great potential for the possible clinical


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translation in the future. We further explored the mechanism of the efficacy of PTX for the treatment of capsular formation using the immunohistochemical analysis. Capsular contracture is a result of excessive foreign body responses due to acute and chronic inflammation after silicone implantation. In general, acute inflammation lasts up to several days. At this stage, neutrophils are predominant at first and then replaced by monocytes migrating from the vasculature as the predominant cell type.62 Subsequently, the persistent presence of the silicone implant leads to the transition from acute to chronic inflammation

and

the

monocytes

rapidly

differentiate

into

CD68-positive

macrophages.62-63 The macrophages (histiocytes) are known to produce growth factors and fibrocyte stimulating cytokines.57,

64

During the chronic inflammation stage,

TGF-β, a key mediator in the regulation of fibrosis,52, 65 is secreted by persistently activated macrophages, lymphocytes and fibroblasts.52-53 The production of TGF-β not only promotes the recruitment and proliferation of fibroblasts, but also stimulates fibroblasts to secrete collagen.53-55 Also, TGF-β drives the differentiation of fibroblasts into α-SMA-expressing myofibroblasts.53-55 Eventually, a contractile force in the capsule tissue caused by α-SMA-presenting myofibroblasts as well as excessive accumulation of collagen results in the fibrotic capsular contracture.12 Therefore, it is important to suppress the expression and activity of TGF-β, α-SMA and CD68 to prevent or mitigate the formation of capsular contracture around the silicone implants. In the current study, the administration of optimal PTX-thermogel formulation (40 µg/mL) and the subsequent sustained release of proper dose of PTX significantly inhibited the expression of TGF-β, α-SMA and CD68 (Figure 9) and thus effectively alleviated the periprosthetic fibrous capsule formation. In addition, the correlation between vascularity and capsule formation is still



unclear and controversial. Lee et al reported that vascularity was irrelevant to the capsule formation.13 However, some researchers claimed that tissues containing more blood vessels were helpful to form softer capsules and reduce the risk of capsular contracture.66 In the current study, our results (Figure 8B) well coincided with the conclusions of some animal and clinical studies that less vascularized structure contributed to a reduced foreign body response around the implants, leading to the formation of thinner fibrous capsules and a lower probability of capsular contracture in breast augmentation.67-68 In the future, more research is called for to reveal the relationship between neoangiogenesis and capsular formation in implantations. It is also worth pointing out that the capsular contracture normally proceeds over one year.69 Therefore, the measurements at one month post-implantation are relatively short. Nevertheless, the initial process of capsular formation in the present model system provides a good prediction of the overall foreign body reaction process. Besides, due to the complexity of foreign body reaction after implantation, a combination delivery system realizing a sustained codelivery of PTX and other drugs may bring about more effective efficacy and lead to lower incidence of capsular contracture. Finally, overcoming the foreign body response to implanted biomaterials, which consists of inflammatory events and wound-healing process, and eventually leads to fibrosis, is a critical need for developing medical devices and implementing new medical advances.5-6 Fortunately, various host responses following implantation of biomaterials share similar mechanisms.5 Therefore, the current study provides inspiring clues for other biomedical materials to resolve the fibrosis-associated complications following the implantation. This microgram-level PTX-loaded thermogel appears to be tremendously promising as an “all-purpose anti-fibrosis coating” for implantable


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biomedical devices due to the easiness and convenience of coating various surfaces without affecting their performances.

5. CONCLUSIONS We propose the local, long-acting delivery of low-dose PTX as a new strategy to alleviate the capsule formation and/or contracture around silicone implants. A biocompatible and biodegradable thermogel composed of PLGA-PEG-PLGA triblock copolymers was prepared by us and used to encapsulate and deliver low-dose PTX. The microgram-level PTX-loaded thermogel system exhibited an almost perfect zero-order release profile in vitro up to 57 days, and the drug release profiles were not affected by changing the drug loading amounts (PTX 80, 40 and 20 µg/mL). In vivo formed fibrotic capsules around the silicone implants coated with the PTX-loaded thermogel with an appropriate dose (40 µg/mL) were significantly thinner. Besides, the capsules exhibited more regular collagen arrangement, lower inflammatory cellularity, vascularity, and expression of TGF-β, α-SMA and CD68 than the bare silicone implants. Also, this optimum dose window had an excellent repeatability for the suppression of capsular formation. Consequently, we conclude that the sustained release of low-dose PTX at the local insertion site is a promising way to mitigate the capsule formation and/or contracture around silicone implants. This microgram-level PTX-loaded thermogel has great potential as an “all-purpose anti-fibrosis coating” for veiling surfaces of various implantable medical devices without affecting their functionalities.

ASSOCIATED CONTENT Supporting Information



The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.XXXXX. Further experimental details and results, including the HPLC chromatogram and standard curve of PTX (Figure S1), in vivo experimental procedures (Figure S2), synthetic procedure (Scheme S1), 1H NMR spectrum (Figure S3) and GPC chromatogram of triblock copolymer (Figure S4), rheological studies of copolymer aqueous solutions with or without PTX (Figure S5), representative histological images and quantitative data of fibrous capsule thickness of an independent control experiment (Figure S6), vascular evaluation of the fibrous capsule (Figure S7), in vivo micro-CT images of the implants (Figure S8), TEM image of micelles of the triblock copolymers (Figure S9) and in vitro release profile of 1.0 mg/mL PTX from thermogel (Figure S10) (PDF)

ACKNOWLEDGMENTS The work was supported by the National Natural Science Foundation of China (grant No. 51773043, 81772363 and 21474019), the State Key Project of Research and Development (grant No. 2016YFC1100300), and the Science and Technology Developing Foundation of Shanghai (grant No. 15JC1490300).


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