Micelle-Embedded Layer-by-Layer Coating with Catechol and

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Applications of Polymer, Composite, and Coating Materials

Micelles Embedded Layer-by-Layer Coating with Catechol and Phenylboronic Acid for Tunable Drug Loading, Sustained Release, Mild Tissue Response and Selective Cell Fate for Re-endothelialization Jiang Lu, Weihua Zhuang, Linhua Li, Bo Zhang, Li Yang, Dongping Liu, Hongchi Yu, Rifang Luo, and Yunbing Wang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.9b01253 • Publication Date (Web): 12 Feb 2019 Downloaded from http://pubs.acs.org on February 14, 2019

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ACS Applied Materials & Interfaces

Micelles Embedded Layer-by-Layer Coating with Catechol and Phenylboronic Acid for Tunable Drug Loading, Sustained Release, Mild

Tissue

Response

and

Selective

Cell

Fate

for

Re-

endothelialization Jiang Lu,† Weihua Zhuang,† Linhua Li, † Bo Zhang,† Li Yang,† Dongping Liu,‡ Hongchi Yu,† Rifang Luo*,† and Yunbing Wang*,† †

National Engineering Research Center for Biomaterials, Sichuan University, Chengdu 610064,

China ‡

West China School of Pharmacy, Sichuan University, Chengdu 610041, China

ABSTRACT： Tunable/sustained drug loading/releasing are of significance in addressing low cytotoxicity, long-term performance and localized mild healing response in biomedical applications. With an ingenious design, a self-healing sandwiched layer-by-layer (LBL) coating was constructed by using chitosan/heparin as the adopted polyelectrolytes with embedding of micelles, in which chitosan backbone was grafted with catechol and the micelle was modified with exposed phenylboronic acid, endowing the coating with enhanced stability by abundant interactions among each coating component (e.g. boric acid ester bond formation, weak intermolecular cross-linking, π-π interactions and Hbonding). Moreover, rapamycin and atorvastatin calcium were selected as drug candidates and loaded into micelles, following with drug releasing behavior study. It was found that the LBL coating maintained a linear growth mode up to 30 cycles, giving a favourable tunability of coating construction and drug loading. The coating could also support sustained release of payloads, and provide wild tissue response. With the systematic in vitro and in vivo study, such catechol-phenylboronic acid enhanced LBL coating

with

drug

loading

would

also

address

enhanced

anti-platelet

adhesion/activation and direct cell fate of endothelial cells and smooth muscle cells via tuning the coating cycles and loaded drugs. With the modular assembling, such coating indicated potential in achieving enhanced re-endothelialization for vascular implants.

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KEYWORDS: endothelialization, layer-by-layer (LBL), micelles, sustained release, catechol, phenylboronic acid

1. INTRODUCTION Cardiovascular disease (CAD) has been the leading cause of death, accounting for more than 40% of all deaths all over the world.1, 2 Cardiovascular stents, especially drug-eluting stents have been a mainstream therapeutic method for the treatment of CAD.

3,4

However, clinical failures occurred frequently after the stents implantation,

such as late stent thrombosis and in-stent restenosis, 5, 6 because of the acute vessel-wall injury caused by the implantation of stents, which resulted in the increased incidence of thrombosis, intimal hyperplasia and restenosis.7,

8

Non-selective drug, like

rapamycin and paclitaxel used in drug eluting stent (DES), can effectively inhibit smooth muscle cells (SMCs) hyperplasia, while on the other hand also suppressed the growth of endothelia cells (ECs) and thus cause delayed endothelial healing which is deemed to be the main reason for late/very late thrombosis. 9, 10 Bioresorbable vascular stents (BVS) have many potential advantages compared with a non-degradable DES and can permit advanced vascular remodeling without a metallic cage and leave only healed native blood vessel tissue after full stents resorption.11 Current study suggested that BVS are safe and effective in patients with ISR.12 However, due to the coatings strategy applied in BVS, which is similar to DES, clinical concerns like delayed healing and late-thrombosis are still major issues. Thus, fabricating a new class of stent coatings that eliminate the adverse effect of anti-proliferative drugs and focused on reendothelialization and anti-thrombus deposition is highly desirable while modulating the proliferation of SMCs.13,

14

As the neointimal growth are caused by SMC

proliferation and macrophages recruitment, coatings that own anti-inflammatory effect are also of interests.15 As commonly accepted, functionalization of surfaces has emerged as a convenient method

for

controlling

interactions

between

materials

and

surrounding

microenvironment.16 In relation to biomaterials, surface properties played important roles in acquiring desired surface bio-functionalities in terms of hydrophilicity,17


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surface roughness,18 functional groups,19 and linked biomolecules.20 Those features could be implemented via surface modification techniques, including chemical, physical and biological approaches.21 Among them, polyelectrolyte multilayers constructed via layer-by-layer (LBL) assembly has been treated as a versatile and easy technique for fabrication of multifunctional coatings in biomedical applications.22 Wohl et al also reported that LBL films could also be used as delivery systems via combining payloads inside and controllable release could be achieved via enzymatic or hydrolytic degradation.23 However, most of the LBL films deliver some payloads in a burst mode once encountered variable environmental stimuli, like ionic strength and pH change.24, 25

As such, there is a growing demand to modify LBL techniques to achieve enhanced

coating stability so as to obtain long-term delivery of payloads.26 Catechols moieties in mytilus edulis foot proteins (Mefps, e.g. Mefp 3 and Mefp 5) are considered as key factors in contributing excellent adhesive properties to various substrates, which strongly implied that biomolecules could also present robust adhesive properties in the presence of catechols.27 In the last decades, polyelectrolytes were also modified with catechols, so as to mimic the components and adhesive functions of mussel adhesive proteins. Lee et al. had modified polyethylenimine and hyaluronic acid with catechols and formed a substrate independent layer-by-layer coating, demonstrating the coating stability.28 Hammond et al. fabricated a layer-by-layer coating using catechol modified polyethylenimine and found it would afford the sustained release of loaded dextran.29 We had also previously constructed a fishing-net like model with valid crosslinking between catechol modified heparin and PEI, with prolonged releasing period of heparin.30 It is now commonly accepted that catechol enhanced coating strategies would benefit for constructing multifunctionalities and would possibly support long-term performance. Though LBL films have widely been investigated, polyelectrolytes used are mainly water soluble, which means it is not an idea platform for the administration of hydrophobic drug.31 Besides, in relation to DES, anti-hyperplasia drug, such as rapamycin, is a hydrophobic drug that could not be easily handled in hydrophilic polyelectrolyte multilayers, due to the solvent features with accompanying uncontrolled



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drug loading and releasing behavior.32 Micelles are formed via the self-assemble of amphiphilic block copolymers in an aqueous solution and have been widely used for drug loading, where hydrophobic drugs are stored in the hydrophobic core.33 Micellar structure is a favorable carrier for tuning drug loading and releasing and thousands of papers are done using smart micelles to facilitate targeted, responsive and precise treatment in cancer therapy, where ‘smart’ are exquisitely designed via modifying the backbone or side chain via introducing pH/Redox sensitive chemical bonds.33,

34

Considering this, integrating micelles into the LBL coating might be a potential approach to facilitate tunable drug loading and controlled release via proper modification of the micelles. As for currently applied DES, the drug release profile significantly impact the overall performance.35 As commonly accepted, the SMCs response, is thought to be largely responsible for stent restenosis, and could last for weeks or months after implantation.36, 37

No inhibition of restenosis occure with too little drug release, reversely, unwanted

effects like delayed re-endothelialization would happen due to too much drug release.38 Thus, sustained release with proper dose of anti-hyperplasia drug is considered important during the healing process.39 Biomolecules that support potential for mediate vascular cell growth could also be brought into coating design. For example, beyond the well known anti-thrombosis formation property of heparin, it could also show antiproliferative effect on SMCs, because the selectin on the SMC’s surface will be blocked by its acceptor.19 Based on this, constructing coatings with combined antihyperplasia drug and anti-coagulant components also present potential for treating cardiovascular diseases. Upon above considerations, herein, we developed a relatively stable LBL assembled coating inspired by sugar-responsive boronate-catechol interactions. Briefly, the polyelectrolytes adopted are heparin and catechol modified chitosan (supporting catechol enhanced LBL coating stability), and the micelles were embedded into the polyelectrolytes in a sandwich-like model. Noteworthily, the micelles are formed via an amphiphilic block copolymer functionalized with phenylboronic acid, which would form borate ester bonding with catechols.40,

41

After that, rapamycin (Rapa) and


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atorvastatin calcium (AC), which could play the role as anti-hyperplasia42 and antiinflammation drug43 are selected as candidates and were further encapsulated into micelles, separately. The study on the coating construction and drug loading/releasing profile was done. Moreover, the stability of LBL coating was done in vitro, to ensure the platform as a coating capacity. The schematic draw is shown in Scheme. 1, by this way, there might be abundant interaction force among the coating components, like physical adsorption, electrostatic adsorption, covalent binding and enriched hydrogen bonding force, giving the coating with enhanced stability. And with the help of such ingenious design, drug loading and release behavior was considered to be better tunable, in terms of diverse drug loading and sustained release profile. In relation to cardiovascular coatings, a detailed vascular cell interaction with coatings was investigated. With the in vitro and in vivo study of the functional coatings, such model could present potential applications for modifying cardiovascular implants. 2. EXPERIMENTAL SECTION 2.1. Materials and reagents 316L stainless steels (316L SS, with the size of 10 mm ×10 mm) were mirror-polished in this study as the model substrate for biocompatibility test. Besides, silicon wafer was used in AFM study due to its high smoothness. The 316L SS wire with a diameter of 100 µm was also used in the in vivo implantation evaluation. Heparin (185 U/mg) was obtained from Macklin. Chitosan (CS; Mw 100 k Da, 80% deacetylated), NaBH4, hyaluronic acid (HA;10 k Da), cholesterol, 2-Hydroxymethylphenylboronic acid (2HMPBA), 4-dimethylaminopyridine (DMAP), N,N'-dicyclohexylcarboimide (DCC), atorvastatin calcium (AC), rapamycin (Rapa), 3,4-Dihydroxybenzaldehyde (3,4-DHB) and cell counting kit-8 (CCK-8) were purchased from Sigma-Aldrich.And

all the

other analytical grade reagents were used or prepared from our lab. 2.2 Synthesis of catechol conjugated chitosan and phenylboronic acid modified micelle molecules Catechol-conjugated chitosan (CS-C) was synthesized using chitosan and 3,4-DHB following previously published method by Lee et. al.44 Details were shown in



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Supporting Information (M1). Micellar molecules are amphiphilic block polymers and synthesized

using

hyaluronic

acid

as

the

backbone,

modified

with

2-

hydroxymethylphenylboronic acid, and using cholesterol to form the hydrophobic core. Detailed protocol could be found in Supporting Information (M2). After synthesis of CS-C and micelle molecule, they were analyzed using 1H-NMR. 2.3 Drug-loading into synthesized micelles Micelle molecules (40 mg) were completely dissolved into DMSO (10 mL) at 80℃, and then atorvastatin calcium (AC, 8 mg) and rapamycin (Rapa, 8 mg) were added into the micelle molecules solution, respectively. After that, DI water (10 mL) was slowly added into above solution drop by drop under vigorous stirring. Finally, the mixture was dialyzed in DI water for 2 days (dialysis membrane bag MWCO = 500 Da). The concentration of the mixture solution was adjusted to 1 mg/mL by evaporation using rotary evaporators and attenuation using DI water, the pH was adjusted to 6.5 by 1M HCl. The drug loaded micelles were denoted as MIC@AC (AC loaded micelles), MIC@Rapa (Rapa loaded micelles). Besides, the co-loading of multi-drugs or micelles into LBL coating layers could also be achieved. 2.4 Characterization of micelles Synthesized MIC, MIC@AC and MIC@Rapa were examined by dynamic light scattering (DLS) to evaluate the zeta-potential and size distribution of the micelles. Prior to test, micellar solutions were incubated at 37℃ for 2 h, followed with the test using Zetasizer (Malvern Zetasizer Nano-ZS90 apparatus, Malvern, UK) at 37℃for 3 times. Besides, the size and morphology of MIC, MIC@AC and MIC@Rapa were further determined by transmission electron microscopy (TEM, Hitachi H-600). Sample solutions were then dropped onto a copper grid coated with carbon, with airdrying for 5 min. After that, the grid was stained with 2 wt% phosphotungstic acid at pH 6.5, the surface water was removed with filter paper. 2.5 Sandwiched LBL (layer-by-layer) coating fabrication The sandwiched LBL coating fabrication process was shown in Scheme. 1. Briefly, the cleaned substrates (316L SS or silicon wafer) were immersed in a dopamine solution


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(2 mg/mL in 10 mM Tris buffer, pH 8.5) for 2 h. After that, they were thoroughly rinsed with DI water for three times and dried with nitrogen gas (the substrates covered with polydopamine coating are denoted as PDA). After that, the alternative layer-by-layer operation was done as following. Firstly, the substrates were immersed in CS-C solutions (pH 6.0) for 10 min. Later on, the samples were immersed in the solution of MIC (pH 7.0, micelles without or with drug loading of AC or Rapa.) for 20 min. After that, the substrates were immersed into CS-C solutions (pH 6.0) again for 10 min. Finally, samples were immersed in negatively charged heparin (2 mg/mL, pH 6.0) solution for 10 min. Above four step treatment is defined as one coating cycle. Besides, after each coating step, samples were placed on a 3D shaker, rinsing with DI water for three times (3ⅹ3 min) and dried with nitrogen gas. The as-prepared coating samples were named as LBLx, where x represents the number of coating cycles. Moreover, micelles embedded samples were labeled as LBLx@AC or LBLx@Rapa or LBLx@AC/Rapa stands for the types of drug loading as AC or Rapa or co-loading of AC and Rapa. 2.6 Characterization of coating construction The coating construction process was monitored by UV-Vis spectrophotometer (UV2401PC, Shimadzu) on a quartz plate and the ultraviolet adsorption between 180 and 600 cm−1 was recorded. Coatings prepared on silicon wafer were examined by scanning electron microscopy (SEM) (Hitachi S-3400N) to study the surface morphology. The cross-sectional view of sample was made via breaking the silicon wafer into two parts with froze by liquid nitrogen for 2 min, and the fracture surface morphology and thickness was studied using SEM.45 Furthermore, the thickness of PDA coating was detected by Elliptical polarization spectrometer accurately. The surface roughness of the coatings was tested using atomic force microscope (AR research MFD 3) with tapping mode. The surface hydrophilicity was measured by the water contact angle (WCA), using a DSA25 contact angle goniometer (Krüss GmbH, Germany). 2.7 Drug loading and releasing test The coatings of LBL30, LBL30@AC, LBL30@Rapa, LBL30@AC/Rapa was done on the



SS substrates and were immersed into 20 mL phosphate buffer saline (PBS, pH = 7.4) at 37 ℃ with continuous shaking at 50 rpm for pre-determined time. At a certain interval, 3 mL of the above immersion solution were taken out and the same volume of fresh buffer medium was added back into the immersion tube immediately. Coatings immersed for 60 days was named as LBL30-60D, LBL30@AC-60D, LBL30@Rapa-60D, LBL30@AC/Rapa-60D. Besides, drug release was also evaluated during immersion. The release properties in vitro of the LBL30@AC; LBL30@Rapa, LBL30@AC/Rapa, and LBL10-10-10@Rapa /N/AC were investigated by puting the coating of these (1 cm × 1 cm) in a dialysis membrane bag (MWCO = 1000 Da) and immersed in 20 mL phosphate buffer saline (PBS, pH = 7.4) at 37℃ with continuous shake at 50 rpm to measure the released drug from the coating. And after a pre-determined interval, the solution (3 mL) were taken out and then the same volume of fresh buffer medium was compensated. The amount of the released drug (AC or Rapa) in the collected medium solution was measured via High Performance Liquid Chromatography (HPLC, Waters 1525 chromatography equipped with a C18 ( 250 mm × 4. 6mm, 5μm)). The mobile phase of the detection of AC was a mixture solution of acetonitrile and water ( 40∶60) and the detection wavelength was at 241 nm. For Rapa, it was a mixture solution of methanol, acetonitrile and water ( 67∶15∶18) and detected at 278 nm (flow rate of 1 mL/min). Especially in this work, the self-healing ability of constructed LBL coating was also done. Firstly, LBL30 coated silicon wafer was treated with a cross scratch on the surface using a scalpel and then was fixed on a glass slide with double faced adhesive tape and viewed under the optical microscope. After recording the morphology of the scratch, the LBL30 coating was transferred and immersed into PBS solution for 1 h at room temperature, following with further optical microscope viewing on the same site of the original scratch. 2.8 Blood compatibility evaluation More than five samples were used for statistical count, and each test was done for more than three times. The whole blood was collected from rabbit in negative pressure tubes containing sodium citrate as the anticoagulant. The hemolysis ratio was detected


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according to previously method.46 DI water was set as the positive control and NaCl solution (0.9 w/v%) was set as the negative control. Prior to platelet adhesion and activation test, the platelet rich plasma (PRP) was obtained by centrifuging whole blood at 1500 rpm for 15min. Briefly, samples were incubated with PRP for 30 min at 37℃, and after fixing and dehydration treatment, the platelets morphologies were viewed using SEM. 2.9 In vitro vascular cell compatibility test 2.9.1 ECs’ viability evaluation EC (endothelial cells) were cultured in RPMI 1640 medium with 10% fetal bovine serum (FBS) and 1% antibiotic Penicillin-Streptomycin. Before seeding, the samples were sterilized by UV for 30 min. The samples were then incubated with 1 mL of ECs suspension solution (with a density of 5 × 104 cells/mL) at 37 °C under 5% CO2 for 4 h, 1 day and 3 days. Four hours before the collecting time day 1 and day 3, the culture medium was removed, then a fresh culture medium containing RPMI 1640 and CCK8 (v/v = 9/1) was added and incubated for another 4 h. Then, the supernatants (0.15 mL) were taken out into 96-well plates, following with the recording at 450nm using Multiskan microplate reader. Besides, samples cultured with ECs were washed completely with saline solution for 3 times and soaked in 2.5% glutaraldehyde solution for 12 h, and then stained with rhodamine (0.1μg) for 20 min. After fully washing, the cells were observed via fluorescence microscope (Nikon, TE2000). 2.9.2 SMCs’ viability test Smooth muscle cells (SMCs) were cultured using DMEM medium containing 1% antibiotic Penicillin-Streptomycin and 10% fetal bovine serum (FBS), and the other process and methods are similar to above protocols mentioned in the ECs viability test. 2.10 Macrophage cell isolation and seeding Prior to macrophage adhesion test, the peritoneum of Sprague−Dawley (SD) rat (abtained from Dashuo Co., Ltd., Chengdu) was used to isolate macrophage cells. After that, macrophage cells were further cultured according to our previously work.47 Firstly, the samples were incubated in 1 mL of macrophage suspension solution (5 × 104



cells/mL) for 1 day at 37 °C with 5% CO2. Then, the samples were washed completely with saline solution for 3 times and soaked in 2.5% glutaraldehyde for another 12 h. After similar treatment for cell culturing and fixing to ECs investigation, cells were viewed by SEM. 2.11 In vivo study All in vivo test were done in compliance with the China Council on Animal Care. Besides, Sichuan University animal use protocol also direct strict animal use, following all the ethical guidelines for experimental animals. 2.11.1 In vivo tissue response test In vivo tissue response test was done based our previous work.48 The implanted and stained methods were shown in Supporting Information (M5), and the implanting schematic diagram was shown in Figure S1 (Supporting Information). 2.11.2 Intravascular SS wire implantation In order to verify the function of the sandwiched like drug loading LBL coatings, 316 SS wire covered with LBL coatings were implanted into the aorta abdominalis of adult SD rats. The implanting schematic diagram was shown in Figure S2 (Supporting Information) and the implanting staining methods48 were shown in Supporting Information (M6). 2.12. Statistical analysis For each test, more than three independent experiments were done, unless described elsewhere. The date was obtained and analyzed using SPSS 11.5 and expressed as a mean± standard deviation (SD). The statistical significance between and within groups was determined using a one-way ANOVA. The values P < 0.05 were considered for having significant difference (*P < 0.05, **P < 0.01, ***P < 0.001). 3. RESULTS AND DISCUSSION 3.1. Synthesis of catechol conjugated chitosan (CS-C) and phenylboronic acid modified micelle molecules In order to give the adhesive properties to modified LBL films, the catechols were


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conjugated to chitosan backbones as shown in Figure S3 (A) (Supporting Information). The 1H-NMR was shown in Figure S3 (B) (Supporting Information), and the value of the catechol proton peaks which appeared from 6.5 to 7.0 ppm indicated successful preparation, and the degree of catechol conjugation was about 31.4%. Besides, the phenylboronic acid modified micellar molecule was also successfully prepared, using the method as shown in Supporting Information (Figure S4 (A) and (B)), the degree of 2-HMPBA substitution was about 16% and the value of cholesterol substitution was about 6%. The structure of catechol-conjugated chitosan (CS-C) and phenylboronic acid modified micellar molecule was seen in Figure 1 (A) and (B). As in this job, the intermolecular interactions in the so-called sandwiched micelles embedded LBL coatings are mainly donated by the polyelectropytes. Different from traditional LBL coatings, whose main force are electrostatic interactions, the two basic components of CS-C and MIC in the micelles embedded LBL coating might do contribution in enhancing the coating stability and maintaining long-term performance of payloads. To better validate the possible interactions between each component, prior to LBLcoating construction, CS-C and MIC were mixed under pH 5.0 and pH 7.0. As seen in Figure 1 (C), the solution of CS-C and MIC were transparent at pH 5.0, onced mixed together, a faint haze like sulotion was observed, suggesting week intermolecular interactions between CS-C and MIC, mainly are electrostatic interactions. However, when the pH value went up to 7.0, the CS-C turned faint yellow, and after mixing with MIC, the mixture was more epinephelos, with no observation of the words behind. The faint yellow occurred because of the slight oxidation of catechols in chitosan to quinones,44 and the epinephelos solution was formed due to boric acid ester formation between catechols and phenylboronic acid.41 Thus, more interaction force between CS-C and MIC make the complex denser and induced epinephelos solution formation. Based on this, the corresponding LBL coating based on CS-C and MIC would also form similar interactions on the substrate-luqid interface and thus do help to acieve a tunable coating construction process. 3.2. Drug-loading into micelles



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The TEM results of MIC, MIC-AC, MIC-Rapa was shown in Figure 1 (D) and the particle size distribution was shown in Figure 1 (E). The particle size of MIC was around 200 nm, and after drug loading, the size was bigger due to the additional contribution in the hydrophobic core of micelles which resulted in the increasing of particle size. The PDI value of MIC was 0.125, indicating that the particle size distribution was more concentrated than the drug loaded micelles. Moreover, the Zetapotential value of these micelles were under -20 mV which derived from the carboxyl (-COOH) of HA that were the hydrophilic ends of MIC, which also strongly implied that it could be applied as an anionic polyelectrolyte in constructing LBL coatings. The averaged particle size, PDI, zeta-potential, drug loading and encapsulation ratio of these micelles were exhibited in Table. 1. The drug loading of MIC-AC or MIC-Rapa was close to 10%, and the encapsulation ratio was more than 40%. These results indicated that such drug loaded micelles could well be applied into the LBL coatings and functionalized as a drug eluting model. Table. 1 The data of micelles. Samples

Averaged

PDI

Zeta Potential

Drug loading

Encapsulation

IDs

diameters(nm)

(mV)

efficiency (%)

efficiency (%)

MIC

194.1±11.25

0.125±0.02

-24.3±2.53

0

0

MIC-AC

240.3±15.46

0.201±0.03

-26.3±3.04

8.12±1.06

44.15±3.74

MIC-Rapa

216.3±23.85

0.345±0.04

-31.2±4.50

9.54±1.54

52.07±5.41

3.3. The sandwiched micelles embedded LBL coating construction The surface morphology of the coating was shown in Figure 2 (A) and the crosssectional view of the LBL coatings was displayed in Figure 2 (B). Obviously seen on LBL30, the porosity was maintained and the basic micellar structure with multispherical particles near 200 nm in diameter was performed. The porosity became less with the addition of drug, which might be ascribed to the effect of drug in the coating, which needs further investigation. And there were a lot of nanoparticles on the coating of PDA, which might be derived from the oxidation and self-aggregation of the


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polydopamine components.49 In the cross-sectional view, the porous structure was seen from the bottom to the top of the LBL coating, indicating the stable construction process in each LBL cycle. As seen in Figure 2 (C) and (D), with the increasing of coating cycles, the thickness present approximately linear growth behavior, which also had good agreement with the increasement of the absorbance at 280 nm (Figure 2 (E) and (F)), indicating the step by step enrichment of catechols or phenylboronic acids. Unlike traditional layer-by-layer assembly, the relatively stable coating construction process might thanks to the introduction of interactions among each LBL components, rather than electrostatic interactions, including covalent bonding (boric acid ester bonding formation between boric acid and catechols at neutral pH,50 abundant π−π stacking interactions and hydrogen bond formation between any adjacent catechol-modified polyelectrolytes.51 Such step-by-step steadily coating growth might make potential contribution in tunable functionality introduction or drug loading. Interestingly, after immersion in PBS solution for 60 days, the coating became denser and the porosity also decreased Figure 3 (A and B), which might be ascribed to the potential of further oxidation and crosslinking of catechols at pH 7.4. As catechols are oxidized to corresponding quinone or semiquinone forms quickly that can reactive with many functional groups, including thiol, amine, and quinone itself.52, 53 Due to the sustained exposing to PBS solution, possible interactions would happen during the immersion period and the as-formed bonding would possibly make the coating denser. Those abundant intermolecular interactions might be helpful in supporting a so-called ‘selfhealing’ property, which presented potential in maintaining long-term administration of pay-loads. Thus, similar investigation was also done. As seen in Figure 3 (C), after immersion in PBS for just 1 h, the size of the artificial scratches on LBL30 were more ambiguous and the gap nearly got full healing. This phenomenon primarily validated the self-healing ability of such sandwiched LBL coating when applied in physiological environment. 3.4 AFM of the coating and water contact angle results The roughness of the LBL coatings with different cycles and different drug loading were shown in Figure 4 (A) and (B). It is obviously that, the surface roughness increased



with the increase of coating cycles. Even if the particle size of drug loaded micelles is a little bit bigger than that of non-drug loading micelles, there were no significant contribution in the surface roughness change. The LBL coating still maintained the relatively smooth surface at the cycle of 10, but present much rougher surface after the coating of 30 cycles. The water contact angle (WCA) results were also shown in Figure 4 (C), the WCA of the Si and the PDA was about 40°, and there was no significant difference between the micelles or drug-loaded micelles embedded LBL coatings at the same cycle. However, the coating became more hydrophilic after 30 cycles, indicating more hydrophilic components deposition, due to the contributions of hydrophilic polyelectrolytes. Besides, we speculated that the increased roughness also make sense in the contribution of hydrophilicity increasement at the 30 cycles deposition. 3.5 Drug release from the coatings The release profile is known to be of significance on the overall performance of a DES. Sustained drug release and therapeutic benefit both are important consideration in DES design.35 Considering surface modification technologies and drugs, alone or in combination, will contribute toward desired performance. Herein, though burst release also existed in the micelles embedded coatings (around 30% released drug after 1 day, Figure 5 (A), (B) and Table 2), with the abundant intermolecular interactions among micelles and polyelectrolytes, also combing the self-healing properties, the drugs loaded in the LBL coatings are considered to present sustained release property, there were still drug released from the coatings after 3 months. These drug release profiles are similar to that of commercial used DES.54 Besides, with the half drug loading amount into the LBL30@AC/Rapa (where only half density of the drug loaded micelles are embedded into the coating), the coatings also present similar release profile, with the only difference in the released amount of drug (Figure 5 (C)). More critically, such sandwiched micelles embedded LBL coating would also tune drug loading amount via simple coating cycles regulation. We took AC as a model drug and found that the total drug loading amount presented a good linear increasement along with the increased coating cycles (Figure 5 (D)). As such, we then speculate that due to the modular assemble of drug loaded micelles and catechol enhanced polyelectrolytes, this coating


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would contribute more in regulating the degree of inflammation, growth of intimal hyperplasia, re-endothelialization process and also anti-coagulation properties. Nowadays, sequential release of drug at a special stage is also full of challenge, because it is not easy to handle diverse drug present different release profile at a single coating. Due to the modular micelles embedding, we also tested the space-time release of different drugs. The drugs release of the LBL10-10-10@Rapa /N/ AC coating (coatings constructed with 10 cycles of Rapa loaded micelles in the bottom and micelles with no drug loaded micelles in the middle 10 cycles and 10 cycles of AC loaded micelles in the top) was shown in Figure 5 (E). The results indicated that the AC had a burst release ratio near 30% and was released preferentially and firstly reached at steady phase of drug release at 40th day, however, as the Rapa was loaded in the bottom, it had a burst release ratio of only 11.26% and reached the steady release until the 100th day (Table 2). So, it was obviously detected that it can realize the successively release of double drugs by this way, which was a significant treatment for the different symptom appeared at different time, like the anti-inflammation and anti-hyperplasia effect. Table 2 The burst release ratio of the drugs after one day.

LBL30@AC

LBL30@Rapa

LBL30@AC/Rapa

LBL10-10-10@Rapa /N/ AC

AC

28.50%

—

32.87%

33.97%

Rapa

—

30.58%

26.44%

11.26%

3.6 Hemolysis ratio of the coating. The results of hemolysis assay were displayed in Table 3. The hemolysis ratio was increased slightly with the addition of the drug, but were all below the SS, and were much less than the medical standard (5%), indicating the safety of using such coatings as blood contacting materials. Table. 3 Hemolysis ratio of the coating. Ratio（%）

Substrate

PDA

3 Cycles

10 Cycles

30 Cycles

SS

1.03±0.08

0.41±0.08

—

—

—

LBL

—

—

0.58±0.08

0.52±0.03

0.56±0.01



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LBL@AC

—

—

0.38±0.01

0.34±0.09

0.66±0.03

LBL@Rapa

—

—

0.67±0.06

0.55±0.02

0.85±0.08

LBL@AC/Rapa

—

—

0.62±0.07

0.41±0.02

0.72±0.04

3.7 Platelet activation and adhesion result. The results of the activation and adhesion of platelet on samples were shown in Figure 6 (A).

The number of adhered platelets on PDA coating and 316L SS substrate was

more than that on LBL coatings (with or without drug) with coating cycles more than 10 (Figure 6 (B)). Also the size of the platelets on LBL coatings are more round, indicating less activation of platelets. Owing to the introducing of anti-coagulation biomolecule heparin, the LBL coating would possibly perform as a potential candidate for blood-contacting application. The more heparin introduced, the less adhered and activated platelets on the coating surface. Taking LBL10 coating as an example, the total released amount of heparin after The heparin also presented sustained release from the coating (Figure S5, Supporting Information) after 2 weeks, sugguesting potential longterm anticoagulation property in real use. 3.8 The cell viability of endothelial cells. Forming a functional endothelial cell (EC) monolayer is considered of significance in addressing normal vascular functions, because of the vital biological function of endothelial cells.47 A material surface that do no harm to endothelial cell proliferation or could promote endothelialization are promising for vascular materials. The ECs behavior of adhesion and proliferation on different coatings at 4 h, 1 day and 3 days were shown in Figure S6 (A), Figure S6 (B) (Supporting Information) and Figure 7 (A). SS is clinically used and showed relatively good EC adhesion and proliferation. As a famous coating in recent decade, PDA is also demonstrated to be beneficial for cell attachment and proliferation.14 There were no obviously difference between the cells adhered on LBL layers at the cycles of 3, as the modification on substrates is not fully sufficient and the loaded drug is also negligible, and many of the cells exhibited typical cobblestone-like morphology. However, with the increase of LBL coating cycles, less cells adhered on the coating surface. The phenomenon showed agreement with those


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published results. On the aspect of LBL coating components, more introduced components of heparin might do harm to EC proliferation (cycles of 30). As reported by Beamish, heparin could inhibite SMC proliferation in a dose-dependent manner (6 μg/mL to 3.2 mg/mL)55 and Ameer also constructed a surface with heparin density of 47.2 ± 14.6 ng/mm2 that can selective inhibit smooth muscle cell proliferation while promote endothelialization.56 Besides, the difference in surface hydrophilicity and stiffness might also affect the cell behavior, with further investigation needed in the near future. On the aspect of loaded drugs, more drug released from the coating and enriched in the culture medium will suppress ECs adhesion and proliferation, especially on the 30 cycles coatings. At the same dose, compared to AC, Rapa had a stronger inhibitory effect on ECs adhesion and proliferation, and that’s why delayed healing of endothelialization on current DES.3 Though suppressed effect observed on drug loaded LBL coatings, the coatings with cycles of 10 still present acceptable cell viability compared to SS and PDA, which are demonstrated to be biocompatible substrates. The cell viability was done using CCK-8 assay. As shown in Figure 7 (B), the value was obviously decreased when the addition of coating cycles, implying the anti-cell proliferation effect with the increasement of the loaded Rapa. 3.9 The proliferation of smooth muscle cells. The SMCs behavior of adhesion, proliferation effected on the different coatings for 4h, 1 day and 3days were shown in Figure S7 (A), Figure S7 (B) (Supporting Information) and Figure 8 (A). There were more cells adhered on the coating SS and PDA after cultured for 4 h than other simples, and with the addition of the layer number, the cells were more less and round shrinkage. At 10 cycles, there were few cells observed, and when the cycles reached to 30, no cells were observed on the surface even after cultured for 3 days. This also could be ascribed to be affected by two aspects, one for LBL coating and another for loaded drugs. Interestingly, at the cycles of 10, LBL10@Rapa and LBL10@AC/Rapa present effective anti-SMCs proliferation effect, yet those coatings still showed affinity to ECs proliferation. The in vitro selective effect of drug loaded LBL coatings in cycles of 10 on ECs and SMCs suggested that such coating



might supply a potential function in directing cell fate to mediate the neointima formation and endothelialization process. The proliferation activity of adherent cells was further studied by CCK-8 assay in Figure 8 (B), which also showed the agreement with the results above. Due to the selective effect on the coating cycles of 10, they were also selected to be applied in the in vivo experiment to further validate the real performance on neointima formation and re-endothelialization. 3.10 in vitro and in vivo inflammatory response Macrophages can produce inﬂammatory cytokines quickly and reach a peak concentrations within 24 h, once they adhere onto a foreign biomaterials surface.47 The morphologies of adhered macrophages were shown in Figure 9 and also the statistical values for the macrophage number was showed in Figure S8 . Pseudopod extending and irregular shape of the macrophages were seen on SS, PDA and LBL coatings without drug loading, indicating activating macrophage potential to a certain extent. Surfaces containing amine groups, like chitosan would activate immunocytes and inflammatory cells and thus in all LBL coatings (without drug loading), the macrophages on the surfaces are more activated.57 Owing to the loading drug of AC and Rapa, the adhered number are far more less on the drug loaded LBL coatings, especially on the coating cycles of 10 and 30. Among them, the best anti-macrophage adhesion and activation effect was shown on LBL30@AC coating, as there were the most anti-inflammation drug AC encapsulation. In vivo tissue response was also done in this work and the subcutaneously implantation results were shown in Figure 10,. The coatings with cycles of 10 were selected to carry on the in vivo experiment. As accepted, with more infiltration of inflammatory cells, granulation tissue development was much stronger. Usually, a more serious tissue response is associated with a thicker fibrous encapsulation. Obviously, more wild tissue response was seen on drug-loaded coatings, representing enhanced anti-inflammation effect with the increased coating cycles. After implanting for 7 days, the capsule on SS sample is the thickest (156.61±15.40 µm), with many infiltrated inflammatory cells and few newborn blood vessel like circles, which implied a severe tissue response. And the immunological experiment was showed in Figure S9 and the green fluorescence (anti-


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CD68 antibody) was represented for the macrophages. The thickness of the fibrous capsule and the inflammatory reaction were indicated that the histocompatibility of the subcutaneous implanted samples. With the increasing of loaded drugs, the fibrous capsules were thinner and the inflammatory reaction were much milder, which showed good agreement with in vitro results. Besides, with time extension, after 14 days implantation, they fibrous capsules have fewer blood capillary on the surfaces excepted that on SS, indicating mature fibrous formation, which implied that there was a good biocompatibility on the drug loaded LBL coatings. 3.11 In vivo animal artery implantation In order to investigate the cell–material interactions of the drug loaded LBL coatings and validate the effect of such coating on re-endothelialization, SS wire and SS wire covered with different coatings were implanted into the aortaventralis of a SD rat. After 30 days implantation, the cross-sections of the blood vessel were stained with HE, and antibodies specific for ECs (anti-CD31 antibody) and for SMCs (anti-α-SMA antibody), as seen in Figure 11. According to the HE stained results, the lager ring was the vascular wall, and the smaller ring was the site where the SS wire bulk was implanted. Obviously, the thickness of neointima was much thicker on SS (173.5±20.65 μm) and PDA (113 ±9.86 μm) covered surfaces, while much thinner on LBL coatings (with or without drug loading). Detailed information was seen on the anti-α-SMA antibody staining and anti-CD31 antibody staining results. The thinner of the small ring, the stronger antiSMCs hyperplasia effect. There was the thinnest small ring on the LBL10@Rapa coating, where the amount of loaded anti-proliferation drug is the highest. Importantly, though thinnest at LBL10@Rapa coatings (14.8±3.04 μm), the CD31 staining results indicated that the neointima did not have confluent endothelial layer, indicating delayed endothelialization.3 As the staining of the outer layer of the neointima on LBL10@Rapa was weaker than the surrounding native blood vessels. Once the drug of Rapa decreased, although the thickness of neointima increased, the re-endothelialization process was better realized. When the loaded drugs were consisted of half Rapa and half AC (LBL10@AC/Rapa), the neointima was thicker (28.01 ± 5.46 μm), yet the CD31 staining of the neointima strongly indicated similar endothelial cell coverage to that on



native blood vessels. The LBL10 coating without drug also had a certain significant effect on anti-SMCs proliferation compared with SS and PDA, owing to the effect of heparin loading. Overall, such micelles embedded LBL coating would be a potential platform for neglect the adverse effect of larger amount of anti-proliferation drug, and maintaining proper endothelialization. Besides, the thickness of neointima could be adjusted via the loaded components and amounts, which is important for the implantation and safe performance of bioresorbable vascular stents, as they would suffer a degradation in the newly formed blood vessel structure. Too thinner neointima was not safe facing the fracture of stents strut and reversely too thicker would possibly cause excessive healing, which is also known as restenosis. Based on our coating design, a tunable drug loading, selective cell fate for re-endothelialization and mild tissue response was achieved, current results effectively demonstrated the potential of such coating to be applied for vascular grafts. The further job of designing and optimizing similar coatings on PLLA stents and implanting into the aortaventralis of New Zealand white rabbit is also under investigation (data not shown). 4. CONCLUSION In this job, we put forward a coating model mainly consisted of sandwiched micelles embedded LBL coatings. The components include catechol modified chitosan and phenylboronic acid modified micelles. Based on such design, there were abundant intermolecular interactions inside the coating components (e.g. boric acid ester bond formation, weak intermolecular cross-linking, π-π interactions and H-bonding), giving the coating with self-healing, tunable drug loading, sustained drug release properties. Along with this, mild tissue response and selective cell fate for re-endothelialization and mediating neointima formation were also achieved. With the systematic in vitro and in vivo study, after drug loading, the better administration of drug could be realized, and such sandwiched micelles embedded LBL coatings presented potential as a coating platform for vascular implants. 

ASSOCIATED CONTENT

Supporting Information


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The Supporting Information is available free of charge on the ACS Publications website at DOI: The methods for characterization of the sandwiched micelles embedded LBL coatings including the synthesis and analysis of catechol-conjugated chitosan (CS-C) (M1); the synthesis and analysis (1H-NMR) of MIC (M2); the drug loading capacity, encapsulation efficiency and the detection methods (M3); the heparin release from the coating (M4); the in vivo subcutaneous implantation studiesand the implantation methords (M5 and Figure S1); the process of the intravascular implantation of wires (M6 and Figure S2). And Figures S3- S9 presented the corresponding results of the 1HNMR of CS-C (Figure S3), MIC (Figure S4), heparin release test (Figure S5); EC cultured for 4 h, 1 day and 3 days (Figure S6) and SMC cultured for 4 h, 1 day and 3 days (Figure S7); the quantification of adhered macrophages (Figure S8) and the immunological study for 7 days and 14 days (Figure S9). 

AUTHOR INFORMATION

Corresponding Authors Email: [email protected], Fax:86-28-85470537, Tel: 86-28-85470537 (Dr. Rifang Luo); Email: [email protected], Tel:86-28-85410280 (Prof. Yunbing Wang, Director of National Engineering Research Center for Biomaterials). Notes The authors declare no competing financial interest. 

ACKNOWLEDGMENTS

We would like to thank those fundings listed as follows. The National Key Research and Development Program (2017YFB0702503, 2016YFC1102200), China Postdoctoral Science Foundation (2018T110976, 2017M612967), the 111 Project (The Program of Introducing Talents of Discipline to Universities (B16033)) and Sichuan Science and Technology Major Project (2018SZDZX0011). 

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Polymeric Membrane Bearing Nanocontainers of Atorvastatin Complexes. Nanomedicine 2017, 12, 2651-2674. (44) Clifford, A.; Pang, X.; Zhitomirsky, I. Biomimetically Modified Chitosan for Electrophoretic Deposition of Composites. Colloids Surf.,A 2018, 544, 28-34. (45) Kim, J.-K.; Mai, Y.-W.; Kennedy, B., Surface Analysis of Carbon Fibres Modified with PVAL Coating and the Composite Interfaces. J. Mat. Sci. 1992, 27, 6811-6816. (46) Luo, R.; Wang, X.; Deng, J.; Zhang, H.; Maitz, M. F.; Yang, L.; Wang, J.; Huang, N.; Wang, Y. Dopamine-Assisted Deposition of Poly (ethylene imine) for Efficient Heparinization. Colloids Surf., B 2016, 144, 90-98. (47) Luo, R.; Zhang, J.; Zhuang, W.; Deng, L.; Li, L.; Yu, H.; Wang, J.; Huang, N.; Wang, Y. Multifunctional Coatings that Mimic the Endothelium: Surface Bound Active Heparin Nanoparticles with in situ Generation of Nitric Oxide from Nitrosothiols. J. Mater. Chem. B. 2018, 6, 5582-5595. (48) Zhang, H.; Xie, L.; Shen, X.; Shang, T.; Luo, R.; Li, X.; You, T.; Wang, J.; Huang, N.; Wang, Y. Catechol/Polyethyleneimine Conversion Coating with Enhanced Corrosion Protection of Magnesium Alloys: Potential Applications for Vascular Implants. J. Mater. Chem. B. 2018, 6, 69366949. (49) Poinard, B.; Neo, S.; Yeo, E.; Heng, H.; Neoh, K. G.; Kah, J. Polydopamine Nanoparticles Enhance Drug Release for Combined Photodynamic and Photothermal Therapy. ACS. Appl. Mater. Interfaces 2018, 10, 21125-21136. (50) Deng, C. C.; Brooks, W. L.; Abboud, K. A.; Sumerlin, B. S. Boronic Acid-Based Hydrogels undergo Self-Healing at Neutral and Acidic pH. ACS Macro Lett. 2015, 4, 220-224. (51) Li, L.; Yan, B.; Yang, J.; Chen, L.; Zeng, H. Novel Mussel-Inspired Injectable Self-Healing Hydrogel with Anti-Biofouling Property. Adv. Mater. 2015, 27, 1294-1299. (52) Sánchez-Cortés, S.; Francioso, O.; Garcı́A-Ramos, J. V.; Ciavatta, C.; Gessa, C. Catechol Polymerization in the Presence of Silver Surface. Colloids Surf., A. 2001, 176, 177-184. (53) Pan, F.; Jia, H.; Qiao, S.; Jiang, Z.; Wang, J.; Wang, B.; Zhong, Y. Bioinspired Fabrication of High Performance Composite Membranes with Ultrathin Defect-Free Skin Layer. J. Membr. Sci. 2009, 341, 279-285. (54) Venkatraman, S.; Boey, F. Release Profiles in Drug-Eluting Stents: Issues and Uncertainties.


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Figure Captions

Scheme 1. The schematic draw of the sandwiched layer-by-layer (LBL) coating: the LBL coating construction process (A), and the diagram of embedded micelles (B), and the cross-sectional view of the components inside the coating (C) and possible existing interactions among the coating (D).

Figure 1. The structure of catechol-conjugated chitosan (CS-C) (A), phenylboronic acid modified micellular molecule (B) and the interactions between CS-C and MIC under pH 5.0 and pH 7.0 (C), respectively; the TEM of MIC, MIC-AC, MIC-Rapa (D) and the particle size of MIC, MIC-AC, MIC-Rapa (E).

Figure 2. The surface morphology (A), cross-section view of different coatings (B), The crosssection view of LBL coatings on different cycle numbers (C) and the thickness of the coatings with different cycle numbers by SEM (D), the UV-vis full spectrum absorbance value of the coatings with various layer (E), and the absorbance of each coating cycle numbers at 280nm (F).

Figure 3. The surface morphology (A), cross-sectional view (B) of the coatings after immersion for 60 days. And Optical microscopy (OM) observation of the sandwiched LBL30 coating before and after immersion in PBS for 1h (C).

Figure 4. The morphology (A) and the roughness of the coatings investigated by AFM (B) and the water contact angel (WCA) of different coatings (C).

Figure 5. The drug release behavior from micelles embedded LBL coatings: the AC release profile of LBL30@AC (A), Rapa release profile of LBL30@Rapa (B), co-release profile of LBL30@AC/Rapa (C), the AC drug loading into the micelles embedded coatings with different coating cycles (LBL1@AC, LBL3@AC, LBL5@AC, LBL10@AC, LBL20@AC, LBL30@AC) (D) and The release profile of AC and Rapa in the coating of LBL10-10-10@Rapa /N/ AC (E).

Figure 6. Morphology (up) and number (down) of adhered platelets on different samples detected


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by SEM (A) and the number of the adhesion of platelet on samples (B).

Figure 7. ECs cultured on different sample surfaces after 3 days (A) and cell viability of ECs on different samples at cultured time of day 1 and day 3 (B). The values P < 0.05 were considered to have significant difference (*P < 0.05, **P < 0.01, ***P < 0.001).

Figure 8. SMCs adhered on different samples after cultured for 3 days (A) and cell viability test of SMCs cultured on different samples at 1 day and 3 days (B). The values P < 0.05 were considered to have significant difference (*P < 0.05, **P < 0.01, ***P < 0.001).

Figure 9. Macrophage evaluation on different samples cultured for 1 day.

Figure 10. HE staining of subcutaneously implanted samples for 7 days (A) and 14 days (B).

Figure 11. Neointima on the wire implanted site with staining of HE, anti-α-SMA antibody and antiCD31 antibody. (Red arrows indicate endothelia layer of neointima and blue ones indicate the endothelia layer of native blood vessel).



For Table of Contents Only 123x66mm (300 x 300 DPI)


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Scheme 1. The schematic draw of the sandwiched layer-by-layer (LBL) coating: the LBL coating construction process (A), and the diagram of embedded micelles (B), and the cross-sectional view of the components inside the coating (C) and possible existing interactions among the coating (D). 112x62mm (300 x 300 DPI)



Figure 1. The structure of catechol-conjugated chitosan (CS-C) (A), phenylboronic acid modified micellular molecule (B) and the interactions between CS-C and MIC under pH 5.0 and pH 7.0 (C), respectively; the TEM of MIC, MIC-AC, MIC-Rapa (D) and the particle size of MIC, MIC-AC, MIC-Rapa (E). 100x117mm (300 x 300 DPI)


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Figure 2. The surface morphology (A), cross-section view of different coatings (B), The cross-section view of LBL coatings on different cycle numbers (C) and the thickness of the coatings with different cycle numbers by SEM (D), the UV-vis full spectrum absorbance value of the coatings with various layer (E), and the absorbance of each coating cycle numbers at 280nm (F). 139x118mm (300 x 300 DPI)



Figure 3. The surface morphology (A), cross-sectional view (B) of the coatings after immersion for 60 days. And Optical microscopy (OM) observation of the sandwiched LBL30 coating before and after immersion in PBS for 1h (C). 156x107mm (300 x 300 DPI)


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Figure 4. The morphology (A) and the roughness of the coatings investigated by AFM (B) and the water contact angel (WCA) of different coatings (C). 71x66mm (300 x 300 DPI)



Figure 5. The drug release behavior from micelles embedded LBL coatings: the AC release profile of LBL30@AC (A), Rapa release profile of LBL30@Rapa (B), co-release profile of LBL30@AC/Rapa (C), the AC drug loading into the micelles embedded coatings with different coating cycles (LBL1@AC, LBL3@AC, LBL5@AC, LBL10@AC, LBL20@AC, LBL30@AC) (D) and The release profile of AC and Rapa in the coating of LBL10-10-10@Rapa /N/ AC (E). 108x117mm (300 x 300 DPI)


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Figure 6. Morphology (up) and number (down) of adhered platelets on different samples detected by SEM (A) and the number of the adhesion of platelet on samples (B). 132x118mm (300 x 300 DPI)



Figure 7. ECs cultured on different sample surfaces after 3 days (A) and cell viability of ECs on different samples at cultured time of day 1 and day 3 (B). The values P < 0.05 were considered to have significant difference (*P < 0.05, **P < 0.01, ***P < 0.001). 128x116mm (300 x 300 DPI)


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Figure 8. SMCs adhered on different samples after cultured for 3 days (A) and cell viability test of SMCs cultured on different samples at 1 day and 3 days (B). The values P < 0.05 were considered to have significant difference (*P < 0.05, **P < 0.01, ***P < 0.001). 127x115mm (300 x 300 DPI)



Figure 9. Macrophage evaluation on different samples cultured for 1 day. 112x64mm (300 x 300 DPI)


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Figure 10. HE staining of subcutaneously implanted samples for 7 days (A) and 14 days (B). 68x64mm (300 x 300 DPI)



Figure 11. Neointima on the wire implanted site with staining of HE, anti-α-SMA antibody and anti-CD31 antibody. (Red arrows indicate endothelia layer of neointima and blue ones indicate the endothelia layer of native blood vessel). 91x52mm (300 x 300 DPI)


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Micelle-Embedded Layer-by-Layer Coating with Catechol and

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