Hyaluronic Acid Nanoparticle Composite Films ... - ACS Publications

Subscriber access provided by Universiteit Leiden / LUMC

Bio-interactions and Biocompatibility

Hyaluronic Acid Nanoparticle Composite Films Confer Favorable Time-Dependent Biofunctions for Vascular Wound Healing Ting Jiang, Zhou Xie, Feng Wu, Jiang Chen, Yuzhen Liao, Luying Liu, An-Sha Zhao, Jian Wu, Ping Yang, and Nan Huang ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/ acsbiomaterials.9b00295 • Publication Date (Web): 15 Mar 2019 Downloaded from http://pubs.acs.org on March 22, 2019

Just Accepted “Just Accepted” manuscripts have been peer-reviewed and accepted for publication. They are posted online prior to technical editing, formatting for publication and author proofing. The American Chemical Society provides “Just Accepted” as a service to the research community to expedite the dissemination of scientific material as soon as possible after acceptance. “Just Accepted” manuscripts appear in full in PDF format accompanied by an HTML abstract. “Just Accepted” manuscripts have been fully peer reviewed, but should not be considered the official version of record. They are citable by the Digital Object Identifier (DOI®). “Just Accepted” is an optional service offered to authors. Therefore, the “Just Accepted” Web site may not include all articles that will be published in the journal. After a manuscript is technically edited and formatted, it will be removed from the “Just Accepted” Web site and published as an ASAP article. Note that technical editing may introduce minor changes to the manuscript text and/or graphics which could affect content, and all legal disclaimers and ethical guidelines that apply to the journal pertain. ACS cannot be held responsible for errors or consequences arising from the use of information contained in these “Just Accepted” manuscripts.

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

Page 1 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

ACS Biomaterials Science & Engineering

Hyaluronic Acid Nanoparticle Composite Films Confer Favorable Time-Dependent Biofunctions for Vascular Wound Healing Ting Jiang1,2‡, Zhou Xie2‡, Feng Wu2, Jiang Chen2*, Yuzhen Liao2, Luying Liu2, Ansha Zhao2, Jian Wu1, Ping Yang2*, Nan Huang2 1. School of life science and engineering, Southwest jiaotong university, No. 111 of the North First Section of Second Ring Road, chengdu 610031, PR China 2. Institute of Biomaterials and Surface Engineering, Key Lab. for Advanced Technologies of Materials, Ministry of Education, Southwest Jiaotong University, No. 111 of the North First Section of Second Ring Road, Chengdu 610031, PR China

These authors contributed equally to this work

‡

*Corresponding author: [email protected] *Corresponding author: [email protected]

1

ACS Paragon Plus Environment

ACS Biomaterials Science & Engineering 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

Abstract: Vascular stent implantation is the primary treatment for coronary artery disease. Surface modification of coronary stents is a topic of interest to prevent thrombosis and restenosis and to promote endothelization. However, bioactive coatings on implants have not yet been fully developed for the time-ordered biological requirements of vascular stents. The first month after vascular stent implantation, the pathological changes in the injured vascular tissue are complex and time-ordered. Therefore, vascular stents possess time-dependent biofunctions with early phase anticoagulant and anti-inflammatory properties. In the later stage, inhibitory effects on smooth muscle cell proliferation and the promotion of endothelial cell adhesion might meet the requirements of vascular repair. We fabricated three types of hyaluronic acid nanoparticles (HA-NPs) by subjecting HA and polyetherimide to ethyl(dimethylaminopropyl) carbodiimide/N-hydroxysuccinimide coupling reaction. The HA-NPs prepared by HA with a molecular weight of 100 kD showed the best stability in a hyaluronidase environment. HA-NP composite films (HA-NCFs) were then fabricated by co-immobilizing selected HA-NPs (100 kD) and HA molecules (100 kD) through amide reaction on PDA/HD coated 316 L stainless steel surfaces. The detachment behavior of HA-NPs (100 kD) in PBS for 20 days indicated that the HA-NPs (100 kD) gradually detached from the surface. In vitro tests (anticoagulant and anti-inflammatory tests, endothelial cells and smooth muscle cells seeding, and bacterial adhesion test) indicated that the newly fabricated HA-NCFs have inhibitory effects on the adhesion of fibrinogen, platelets, macrophages, bacteria, SMCs, and ECs. As the HA-NPs detached from the surface, the HA-NCFs showed excellent gradual comprehensive biocompatibility, which promoted adhesion and proliferation of ECs while still exerting inhibitory effects on the platelets, macrophages, and SMCs. Finally, in vivo SS wire implantation test (aortic implantation in healthy Sprague-Dawley rats) showed that HA-NCFs possessed anti-inflammatory properties, inhibited the proliferation of smooth muscle cells, and promoted re-endothelialization. In particular, HA-NCFs with time-dependent biofunctions showed better anti-restenosis effects than those of surfaces modified with molecular HA, which exhibited constant biocompatibility. This study provides an important basis for the construction of HA-NP composite films with favorable time-dependent biofunctions for the time-ordered biological requirements of vascular stent. Key words: HA nanoparticles (HA-NPs), Anticoagulant, Anti-inflammatory, Endothelial Cell, Smooth Muscle Cell, Time-dependent biofunctions Introduction Percutaneous coronary intervention (PCI) is the main approach in the treatment of coronary artery disease. Compared with coronary artery bypass grafting, PCI has the advantages of low risk and short recovery time1. However, traditional materials for vascular stents, such as 316LSS stainless steel (SS), have the disadvantage of insufficient biocompatibility, which induces adverse cardiac events, such as thrombosis, inflammation, and vascular restenosis2-3. Although the emergence of drug-eluting stents (DES) to address in-stent restenosis has led to a revolution in interventional cardiology, side effects, such as inflammation, late thrombosis, and late restenosis still cause complications4-5. After PCI, the pathological changes in bare metal coronary stent implantation are time-ordered, which can be divided into three stages. The first stage occurs within one month of stent implantation in the form of acute and subacute stent thrombosis. Acute thrombosis is usually observed within a few minutes to several hours of implantation, while subacute thrombosis takes place within a few days to a 2


Page 2 of 32

Page 3 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


month, with the highest risk of occurrence during the first week after the clinical stent implantation6. While bacterial infection is rare after clinical stent implantation, its development could be fatal7. PCI treatment often induces mechanical injury of blood vessels, and in turn, the injured endothelium recruits. Subsequently, this is infiltrated by inflammatory cells. Inflammation usually takes place within several days to weeks of implantation. In this early phase, inflammatory response are accentuated in neointimal segments directly adjacent to the stent8. As a result of the thrombosis and inflammation, the intimal thickness increases considerably within 4–6 weeks in clinical stent implantation. The smooth muscle cells (SMCs) are the main cellular component of the neointimal tissue9. The second stage of pathological changes occurs within one to three months after PCI. The surface of the stent is gradually covered by surrounding tissues. However, incomplete endothelial repair gives rise to coagulation and restenosis10. In the third stage, which occurs after three months, the stent is completely covered with the surrounding tissue. However, the adverse reactions of materials and tissues may cause late coagulation and restenosis11-12. In stent design, the biological function of the stent should vary according to the sequential pathological changes. In this way, the stent could respond appropriately to the biological environment of different pathological stages after implantation. In the first month, the surface of the stent should be modified such that it meets certain requirements. First, it should prevent thrombosis, inflammation, and bacterial infection within one to two weeks6-7,

13.

In addition, it should inhibit excessive SMC

proliferation to prevent intimal hyperplasia and accelerate endothelial repair to improve fused endothelium formation on the stent surface within one month. Stents with these favorable time-dependent biofunctions could eliminate adverse reactions in the subsequent stages14. However, stents that are currently applied, do not exhibit such time-ordered activity. For instances, DESs and antiproliferative drugs have exhibited delayed endothelium regeneration that may raise the incidence of adverse cardiac events in stage 215-16. Biodegradable polymer DES17-18 has been developed to reduce the incompatibility of non-degradable scaffold. However, the inflammatory reactions and intimal hyperplasia might be caused by the degradation products of biodegradable polymers at stage 2. After a complete degradation of the polymer, the exposed material surface may also cause restenosis and late thrombosis at stage 3. The latest generation of stents is designed to be completely absorbed over time. However, compared with Everolimus-Eluting Metallic Stents, the Everolimus-Eluting Bioresorbable Scaffolds appear to possess lower efficacy and higher thrombotic risk over time19. A new type of DES, OrbusNeich’s Genous stent, has the function of capturing endothelial progenitor cell20. However, because of the lack of angiographic information to assess the late loss of the new stent, more in-depth research is needed21. Hyaluronic acid (HA) is a linear polysaccharide, which is a component of the extracellular matrix of mammalian connective tissue22. The biofunctions of HA are related to its molecular weight23. Generally, low-molecular-weight HA (LMW-HA) promotes inflammatory response and SMC proliferation, whereas high-molecular-weight HA (HMW-HA) suppresses them24. Our previous study was conducted on healthy Sprague-Dawley rats and New Zealand white rabbits. It showed that the surface modification using HA with a molecular weight of 100 kD conferred stents with anti-inflammatory, anti-proliferative, and re-endothelialization properties. However, the anti-coagulant, anti-inflammatory, and anti-proliferative abilities of the 100-kD HA-modified surface remain to be improved further25. On the contrary, HMW-HA (500–1000 kD) exhibited strong inhibitory effects on the adhesion of platelets26, macrophages27, and SMCs26. Furthermore, HMW-HA can be decomposed in vivo, resulting 3



in a gradual decrease in the inhibitory effect of the stent26. However, a major disadvantage of the surfaces modified with HMW-HA is that HMW-HA is decomposed into LMW-HA in vivo by decomposers, such as hyaluronidase (HAase) and reactive oxygen species28-29. LMW-HA, as a signal factor and a matrix component, could promote inflammation and proliferation in vivo30. Therefore, it is desirable to find a special HA structure with strong inhibitory effect on cell adhesion like HMW-HA that could gradually detach from the surface in response to sequential pathological changes, while refraining from being easily decomposed into LMW-HA in vivo. In recent years, nanoparticles of glycosaminoglycans, such as heparin31, have been shown to exhibit good inhibitory effects on cell adhesion, similar to those of HMW-HA. We speculate that HA nanoparticles (HA-NPs) have similar functions as those of HMW-HA. At the same time, HA-NPs (200–400 nm in diameter) have good stability in the blood environment and can circulate in the blood for several days32-33. Finally, a long-term controlled release of paclitaxel was achieved by preparation of paclitaxel loaded HA grafted Poly(lactic-co-glycolic acid) nanoparticles, which inhibited smooth muscle hyperplasia after stent implantation34. Polyetherimide (PEI) is a type of cationic polymer. PEI was commonly used as an efficient non-viral transfection reagent for the delivery of plasmid DNA and small RNA molecules35-36. Due to its high cationic charge density and a large number of protonable nitrogen atoms, PEI is able to form stable complexes with nucleic acids37. The cytotoxicity of PEI is known to be molecular weight-dependent. A high molecular weight of PEI is more cytotoxic than a low molecular weight38. In addition, high molecular weight of PEI may lead to a decrease of both fibroblast and osteoblast as well as adhesion and proliferation39 The cytotoxicity of PEI with lower molecular weight could be neglected40. The degree of the toxicity also depends on the cellular type41-43, including human vascular endothelial cells and murine macrophages44. Another cationic polymer, poly-L-Lysine(PLL) was also commonly used as a gene delivery. However, the effectiveness of PEI to deliver gene to the cells is higher than PLL45. The cytotoxicity of PLL is known to be molecular weight-dependent. In addition, lower molecular weight PLL also showed significantly decreased cytotoxicity 46-47. Considering the above, PEI and HA are usually assembled to form nanoparticles to carry plasmid DNA targeting the cells48-50. Accordingly, HA-NPs were prepared using HA and low molecular weight PEI. Based on this knowledge, we assumed that simultaneous immobilization of HA-NPs and 100-kD HA on the stent surface may confer favorable time-dependent biofunctions for vascular wound healing. In the early stage of stent implantation (one to two weeks), HA-NPs are supposed to inhibit side effects, such as coagulation, inflammation, and intimal hyperplasia. In the later stage (two to four weeks), as expected, HA-NPs gradually detached from the surface of the stent, and the 100-kD HA gradually exposed to promote endothelial repair, acquiring excellent comprehensive performance. Since it was first reported in 200751, polydopamine (PDA) has become an important transition coating on the surface of materials. Yang et al.52 prepared polydopamine-hexamethylendiamine (PDA-HD) transition layers that were rich in amino, and possessed good biocompatibility. Because HA contains an abundance of carboxyl groups, 100-kD HA and 100-kD HA-NPs can be immobilized onto PDA-HD-coated stent surfaces through the formation of amide bonds via ethyl(dimethylaminopropyl) carbodiimide/N-hydroxysuccinimide (EDC/NHS) coupling reaction. Accordingly, HA-NPs were prepared using HA of different molecular weights and polyetherimide (PEI) to obtain the best stability in blood environments in the presence of HAase50, 53. Subsequently, 100-kD HA-NPs and 100-kD HA were simultaneously immobilized onto the surface of PDA-HD-coated SS to prepare 100-kD HA-NP-modified PDA/HD films (HA-NCFs). We evaluated the elution rate of 100-kD HA-NPs and 4


Page 4 of 32

Page 5 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


the change in biocompatibility on the surface of the material. In addition, the in vivo biocompatibility of the HA-NCFs was examined by abdominal aortic filament implantation in healthy Sprague-Dawley rats. Materials and Methods HA-NPs Preparation First, 1 mg/ml of HA (Sangon Biotech, China) with three different molecular weights (4 kD, 100 kD, 750 kD) in water solution was activated by the EDC/NHS (Sangon Biotech, China). Then 1 ml 0.5 mg/ml Polyethylenimine, branchedPEI (Sangon Biotech, China, molecular weight = 1.5 kD) is mixed with 10 ml 1 mg/ml HA solutions for 10 min, and then the HA nanoparticles (HA-NPs) are formed by electrostatic interaction and amide reaction between the HA and the PEI. Thus we obtained the solutions of HA-NPs and the HA. The three types of HA-NPs are named based on the molecular weight of HA, which are called HA-NPs (4 kD), HA-NPs (100 kD), and HA-NPs (750 kD). Characterization of HA-NPs The shape of HA-NPs was detected by transmission electron microscope (Tecnai G2 F20 S-TWIN, FEI). The average particle size, particle dispersion index (PDI), and zeta potential of different HA-NPs were measured by dynamic light scattering (DLS) using a ZETA-SIZER, (MALVERN Nano-2S90, Malvern Ltd Malvern, UK). To detect the stability of HA-NPs in enzyme environment in vitro, the three HA-NPs were treated by hyaluronidase (HAase, Sigma, China) solutions (20 U/mL) at 37 C for 24 h, and the size and PDI were determined. Preparation of HA Nanoparticles Composited Films (HA-NCF) First, the PDA/HD coating was deposited on a mirror-polished 316L stainless steel (316LSS, diameter 8 mm, Baoji, China). This is described in detail elsewhere 52. The process of preparation is shown in Fig 1. The HA-NPs (100 kD) and HA (100 kD) solutions were prepared by the method described in Section 2.1. The HA-NPs (100 kD) and HA (100 kD) solutions were reactivated by the EDC/NHS method. Subsequently, the 316L SS coated with PDA/HD were immersed into the HA-NPs (100 kD) and HA (100 kD) solutions for incubation. The process is discussed below. After conducting reaction for 24 h, the samples were washed with ultra-pure water (three times, for 5 min) before surface analysis. The HA (100 kD) solely modified PDA/HD films (HA-F) were used as the control.

5



Figure 1. Diagram of strategy for the fabrication of HA-NCF and HA-F. Characterization of HA-NCF The amine density of HA-NCF was detected according to Hamerli et al.'s method54. The chemical properties of HA-NCF were detected by Fourier Transform Infrared Spectroscopy (NICOLET 5700 infrared spectroscopy, USA). The hydrophilicity of HA-NCF was examined using a drop shape analysis system (DSA 100, Kruss, Germany) applying the sessile drop method (5 μL droplet). Time Related Release Behavior of HA-NPs on HA-NCF in PBS The HA-NCF samples were immersed in phosphate buffer solution (PBS) (PH 7.4) solutions at 37 C for 0, 1, 3, 5, 10, 20 days,. Then the samples were thoroughly washed with ultrapure water for three times and dried by nitrogen. Optical microscope (Carl Zeiss, Germany) dark field images and scanning electron microscope (Quanta 200, FEI, Holland) pictures were taken to observe the shedding and degradation of surface particles. Time Related Biofunctions of HA-NCF Time Related Anticoagulant ability of HA-NCF The time related anticoagulant ability of HA-NCF was investigated by platelets adhesion, platelets activation and fibrinogen adhesion test. Three parallel tests were conducted per sample. Fresh citrate anticoagulated human whole blood from Chengdu Blood Center in China was used in this work. For the platelet adhesion test, the full blood sample was centrifuged at 1500 rad/min for 15 min to obtain the platelet-rich plasma (PRP). Subsequently, 70 μL of PRP was added onto the surface of HA-NCFs, which were soaked in PBS solution for 0, 1, 3, 5, 10, 20 days, in advance. After culture at 37 C for 1 hour, the samples were washed with PBS for three times, and ultra-pure water for three times. The adherent platelets were fixed with 2.5% glutaraldehyde (Sangon biotech, china) solution at 4 °C for 2 h. The morphology of the adhered platelets was observed under the fluorescence microscope (DMRX, Leica, Germany) after fixing and typical Rhodanmine staining (Sigma, USA). The platelet

surface coverage in each sample was counted using 15 random images (size: 400x) by imageJ software. The platelet activation was determined by detecting the expression of GMP14055. For this purpose,

70 μL of PRP was added onto the surface of HA-NCFs, which were soaked in PBS solution for 0, 1, 3, 5, 10, 20 days, in advance. After culture at 37 ℃ for 1 hour, the samples were washed with PBS for three times and ultra-pure water for three times. Subsequently, the samples were blocked with wt. % bovine serum albumin (BSA) (Sigma, USA) in PBS at 37 C for 30 min. Subsequently, the samples were thoroughly washed again and covered with 20 μL of mouse anti-CD62p (1:100) also called GMP-140 (MCA796GA, Serotec Co., Japan) and incubated at 37C for 1 h. After washing for three times with PBS, 20 μL of horseradish peroxidase conjugated sheep anti-mouse polyclonal antibody (second antibody) (HRP, Catalog No: 074-1806, KPL Co., South Korea) solution was added to the samples and incubated for 1 h at 37 C. Then, the samples were thoroughly washed with PBS. Next, 70 μL of chromogenic substrate 3,3,5,5-tetramethylbenzidine (TMB) (Bioss, China) solution (diluted 1:4 in PBS) was added to the sample surface. As a next step, 10 min later, 70 μL of 1 M H2SO4 was used to 6


Page 6 of 32

Page 7 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


end the reaction and the optical density at 450 nm was determined by a microplate reader. The relative amount of GMP-140 was quantified according to the calibration curve. For the fibrinogen adsorption test, the full blood was centrifuged at 3000 rad/min for 15 min to obtain the platelet-poor plasma (PPP). In addition, 70 μL of platelet-poor plasma (PPP) was added onto the surface of HA-NCFs, which were soaked in PBS solution for 0, 1, 3, 5, 10, 20 days in advance. These were incubated at 37 °C for 1 h After incubation with PPP. Subsequently, the samples were rinsed thoroughly with PBS and blocked with 1 wt% bovine serum albumin (BSA) in PBS at 37 °C for 30 min. Subsequently, the samples were thoroughly washed again and covered with 20 μL of Horseradish Peroxidase (HRP)-labeled mouse antihuman fibrinogen monoclonal antibody (primary antibody, diluted 1:200 in PBS; Sigma, St. Louis, MO) and incubated at 37 °C for 1 h. The samples were thoroughly washed with PBS, following which, 70 μL of chromogenic substrate 3,3,5,5-tetramethylbenzidine (TMB) solution (diluted 1:4 in PBS) was added to the sample surface. Subsequently, 10 min later, 70 μL of 1 M H2SO4 was used to end the reaction. The optical density at 450 nm was determined by a microplate reader. The relative amount of adsorbed Fibrinogen was quantified according to the calibration curve. The Fibrinogen adsorption was detected as reported previously56. Time Related Anti-inflammation Ability of HA-NCF Macrophage cells were isolated from the peritoneum of Sprague−Dawley (SD) rat (Dashuo Co., Ltd., Chengdu) and cultured according to the method described in

57.

Three parallel tests were

conducted for each sample. For macrophage cells’ seeding, 1 mL of 1×105 cell/ml fifth generation cell suspension was added to the surface of HA-NCFs, which were soaked in PBS solution for 0, 1, 3, 5, 10, 20 days, in advance. These were incubated at 37 °C under 5% CO2. After 2 h incubation, nonadherent cells were removed by washing with PBS, and 1 mL of fresh DMEM (sigma) was then added to the samples for 24 h incubation. Thereafter, the samples were washed three times with PBS and fixed in 4% paraformaldehyde for 12 h at room temperature. Subsequently, 50 μL of rhodamine was added to each sample surface and incubated at 37 °C for 30 min. After being rinsed three times with PBS, three parallel tests were conducted per sample. The samples were observed by fluorescence microscopy (DMRX, Leica, Germany). The number of cells in was statistically counted by 15 random images (size:

400x) from each sample. The number of cells was counted by imageJ software. Human Umbilical Vein Endothelial Cells and Human Umbilical Arterial Smooth Muscle Cells Isolation and Seeding Endothelial cells (ECs) were isolated from human umbilical vein. Briefly, to remove the residual blood, the lumen of umbilical vein was first thoroughly washed with physiological saline. Medium 199 (M199)(sigma) containing 0.1% type II collagenase (sigma) was then injected and incubated at 37 °C for 15 min. The digestion was stopped by filling M199 containing 10% FBS. Thereafter, the cell suspension was eluted and centrifuged at 1000 rpm for 5 min. The supernatant was subsequently removed, and the precipitate of ECs were resuspended in M199 containing 15% FBS and 20 μg/mL endothelial cell growth supplement (ECGS). This was then incubated at 37 °C under 5% CO2. Smooth muscle cells (SMCs) were isolated from human umbilical artery. Briefly, adventitia layer and endothelium layer that surrounded umbilical artery were first stripped away, and the residual tissue 7



was cut into small pieces. Then, the artery pieces were cultured in Dulbecco’s Modified Eagle Medium/F12 (DMEM/F12) containing 10% FBS at 37 °C under 5% CO2. SMCs were obtained by outgrowth from these tissue pieces. All the samples were sterilized by ozone before cell seeding. Primary ECs were seeded on the samples at an identical density of 1 × 104 cells/ml and incubated at 37 °C under 5% CO2 for 4 hours, 1 day and 3 days. The morphology of the adhered ECs was observed under the fluorescence microscope (DMRX, Leica, Germany) after the fixing, and typical Rhodanmine staining (Sigma, USA). The cell count (cells/mm2) was statistically counted by 15 random images (size: 40x) from each sample. The fifth generation SMCs were seeded on the samples at an identical density of 5 × 104 cells/ml and incubated at 37 °C under 5% CO2 for 1 day. The morphology of the adhered SMCs was observed under the fluorescence microscope (DMRX, Leica, Germany) after the fixing and typical Rhodanmine staining (Sigma, USA). The detail can be obtained from14. The surface coverage (%) of SMCs was statistically counted by 15 random images (size: 40x) from each sample. Antibacterial Properties of HA-NCF Bacterial seeding method was reported previously in

58.

Briefly, Pseudomonas aeruginosa

(P.aeruginosa) and Staphylococcus aureus (S.aureus) were obtained from Sichuan Provincial people's Hospital, and were cultured in incubators at 37 °C. After the emergence of multiple colonies, diluted with cell culture medium and adjusted the density to 1×106 CFU/ml, 1 ml solution was added to the well where the sample was placed, then incubated at 37 °C for 24 hours. After fixing and drying by the chemical method, adhesion and activity of bacterial on samples were observed by fluorescence microscope and cck-8 Kit, respectively. In Vivo Animal Experiment All procedures are in accordance with the animal use agreement between the China Animal Protection Commission and Southwest Jiaotong University. All ethical guidelines were followed for laboratory animals. To simulate the stent modified by HA-NFC in the blood vessel, 316LSS wires (φ = 0.1 mm, length = 12 mm) modified with HA-NCF were implanted into abdominal aortic lumen of about 300 g adult Sprague−Dawley (SD) rat (Dashuo Co., Ltd., Chengdu) for 30 days. Then, the implanted wires were extracted. The remained vessels tissues were prepared into paraffin section and were stained with hematoxylin and eosin (H&E), α-SMA (Sigma), TNF-α (Sigma) and VWF (Sigma) 59-61.

The 15 randomly selected typical optical photographs of the HE-stained cross-sectional slices

were used to calculate the hyperplastic area around the implantation site by imageJ software. Statistical Analysis At least three parallel tests were conducted per sample. The statistical signiﬁcance between the sample groups was assessed by SPSS11.5 software using one-way ANOVA and Tukey post hoc test. When the probability was p < 0.05, the difference was considered statistically signiﬁcant (p < 0.05). Results and Discussion Characterization of HA-NPs 8


Page 8 of 32

Page 9 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


PEI is a cationic polymer containing a large amount of amino groups. HA is rich in carboxyl and hydroxyl. Based on the amides reaction and electrostatic interaction between HA and PEI, it was reported that HA-PEI nanoparticles have been prepared for delivery of SiRNA into tumor cells62-63. Fig. 2A shows the possible mechanism of HA-NP formation, which is mainly based on both electrostatic forces and amide reactions between the carboxyl group of HA and the amino group of PEI. The particle dispersion index (PDI) is an important indicator of the size distribution of suspended particles. When the PDI is less than 0.3, the nanoparticles are considered to be uniformly dispersed64. As shown in Fig. 1B, the particle size and PDI of 4-kD HA-NPs were 155.5 nm and 0.071, respectively. After treatment with HAase for 24 h, the particle size and PDI were 383.5 nm and 0.078, respectively. Because the PDI was almost unchanged, the change in particle size was presumed to be caused by the swelling of nanoparticles, which might have resulted from the unique capacity to bind and retain water of HA65. The particle size/PDI of 100-kD HA-NPs was changed from 354.3 nm/0.100 to 380.9 nm/0.088 after HAase treatment. A smaller PDI value indicates a higher degree of dispersion. Meanwhile, the particle size increased, which may have been caused by slight particle swelling. The particle size/PDI of 750-kD HA decreased from 1399 nm/0.557 to 383.5 nm/0.174 after HAase treatment. The decrease in PDI suggested that the particle size distribution was more uniform, and the large particle size could be attributed to the degradation of 750-kD HA and 750-kD HA-NPs by HAase, which induced the generation of LMW-HA, and resulted in further inflammation. The above results indicated that nanoparticles assembled from LMW-HA are more difficult to decompose than HMW-HA particles in the presence of HAase. Meanwhile, 100-kD HA-NPs had the best stability in HAase environment, indicating their stability was the highest in blood environments in vivo. Furthermore, our previous research verified that for both 4-kD and greater than 500-kD HA induced inflammation in vivo, while 100-kD HA inhibited inflammation25. Therefore, we further characterized 100-kD HA-NPs, and chose them for the subsequent preparation of HA-NFCs. As shown in Fig. 2C, the solution of 100-kD HA-NPs was cloudy, signifying the formation of nanoparticles. Transmission electron microscopy showed that the diameter of 100-kD HA-NPs was approximately 300 nm, which was consistent with the measurement of particle size in Fig. 2B. The zeta potential of 100-kD HA-NPs in phosphate-buffered saline (PBS) at pH 7.4 was approximately -35 mV, indicating strong electronegativity (Fig. 2C). According to the design of our research, we expected that HA particles could be eluted from the scaffold, and these would have a targeting effect for atherosclerotic plaque. Therefore, we selected the particle size according to the research of Ga Young Lee et al.66. The size of the HA nanoparticles is acceptable in the range of 200–350 nm.

9



Figure 2. (A) Diagram of HA-NPs assembly mechanism. (B) Particle size and PDI of HA-NPs (4 kD, 100 kD, and 750 kD) before and after 20 U HAase treatment (mean ± SD, N = 3). (C) Phone photo, transmission electron microscopy image and zeta potential of HA-NPs (100 kD). Characterization of HA-NCFs As shown in Fig. 3A, the PDA/HD coating induced a yellow color on SS, and immobilization of HA-NPs changed the surface from a mirror-like to a misty appearance. Scanning electron microscopy (Fig. 3B) revealed the surface topography of the HA-NCFs, whereby HA-NPs with a diameter of approximately 300 nm were immobilized on the surface. In comparison with PDA/HD, HA-F (films immobilized with molecular HA only) and HA-NCF possessed significantly fewer amino groups (Fig. 3C). Fourier transform infrared spectroscopy (Fig. 3E) of HA-NCFs showed absorption peaks at 3400 cm-1 and 1750 cm-1, which were attributed to –COOH groups. This observation confirmed the surface immobilization of HA-NPs and molecular HA. The absorption peaks at 1640 cm-1 and 1550 cm-1 were attributed to the amide I and amide II bonds, respectively31, indicating that some of the HA-NPs and molecular HA were immobilized onto the surface through amide reactions. There is a strong electrostatic interaction and hydrogen bonding between PDA and HA molecules67 and the primary amine groups of the HD could be effectively immobilized to the carboxylic groups of HA25. Both amide reactions and electrostatic force contributed to HA-NP immobilization. The surface hydrophilicity affects protein adsorption, and is an important factor influencing cell adhesion and fate68. Because HA is hydrophilic, the HA-NCF was more hydrophilic than PDA/HD and HA-F (Fig. 3D), indicating that more HA was immobilized onto the HA-NCF. Above all, we had grafted both the HA and HA-NPs simultaneously onto the PDA/HD films successfully, which was never performed before.

10


Page 10 of 32

Page 11 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 3. (A) Visual appearance and (B) Scanning electron microscopy of SS, PDA/HD, HA-F, HA-NCF. (C) Density of amine groups for the samples (mean ± SD, N = 3, here *** represents PDA/HD with significant difference from the other three samples). (D) Contact angle measurements (mean ± SD, N = 3). (E) Fourier transform infrared spectroscopy of PDA/HD and HA-NCFs. Time-dependent Detachment Behavior of HA-NPs on HA-NCFs A few particles were observed on the surface of PDA/HD and HA-F, possibly corresponding to PDA nanoparticles formed during the polymerization of the PDA/HD coating (Fig. 4). HA-NPs fully occupied the surface of the newly fabricated HA-NCFs (HA-NCF (0 d), after immersion in PBS for 11



one to three days. The surface coverage of HA-NPs decreased from almost 100% to approximately 76%. When the immersion time was extended to five days, the surface coverage of HA-NPs decreased to approximately 42%. When the immersion time prolonged to 10–20 days, the surface coverage of HA-NPs decreased to 5–14%. As shown in Fig 4B, the number of particles remaining on the HA-NCF surface decreased significantly in 5, 10, 20 days. However, it was still higher than that on the PDA/HD and HA-F surfaces. When HA-NCFs were immersed in PBS, the number of adhered nanoparticles decreased with increasing immersion time. This indicated a rapid release of the biomolecules. In PBS, there are a large number of cations and anions remained in the PBS solution. The electrostatic assembly force of HA-NPs may be interfered by ions in PBS, leading to their degradation, which was similar to the degradation behavior of heparin/poly‑L‑Lysine nanoparticles in PBS for 28 days14. The long-term controlled release of paclitaxel was achieved by Kim T. et al. to fix paclitaxel loaded HA grafted Poly (lactic-co-glycolic acid) nanoparticles to the stent surface. This approach inhibited smooth muscle hyperplasia after stent implantation34. The release profile of HA-NPs detachment from the surface, may confer the novel time-dependent biofunction to an implanted stent. However, such a device will need to be implanted into an atherosclerotic vessel wall. It has been previously shown that the concentration of Reactive Oxygen Species (ROS)69 and activated hyaluronidase70 are elevated in atherosclerotic diseased conditions when compared to normal physiological conditions. Both ROS71 and hyaluronidase 72

could degrade HA in vivo, and might result in a faster release of HA-NPs than observed here. Further

study regarding HA release behavior under pathological conditions is required.

Figure 4. (A) Optical microscopic visualization (dark field image, the brightened dots represented HA-NPs) of HA-NCFs after immersion in PBS for various durations. (B) Surface coverage of particles on HA-NCFs after immersion in PBS for various durations. * indicates significant difference compared to HA-NCF (0 d). Blood Compatibility To evaluate the blood compatibility of the materials, SS and PDA/HD were used as positive controls whereas HA-F served as a negative control (Fig. 5A–5D). Initially, HA-NCFs exhibited strong anti-coagulant ability (at 0 d). The platelet count and spreading on the HA-NCFs (0 d) were significantly lower than those on the positive and negative controls. With increasing immersion time in 12


Page 12 of 32

Page 13 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


PBS, the HA-NCFs showed a gradual decrease in anti-coagulant ability. After soaking for a short time (0–3 d), the nanoparticle surface coverage decreased from 100% to 76%, the nanoparticle surface coverage was relatively high, the anti-coagulant effect of HA-NPs was good. As time went on (3–10 d), the nanoparticle surface coverage decreased from 76% to 14%. However, the anti-coagulant effect of HA-NPs was still good. With prolonged soaking time (10–20 d), the remaining HA-NPs gradually fell off. The nanoparticle surface coverage decreased from 14% to 5% exposing the molecular HA-modified surface and resulting in decreased anti-coagulant activity. However, after 20 days of immersion, platelet spreading on the HA-NCFs (20 d) was still lower than those on both positive and negative control surfaces (Fig. 5A). The surface coverage (Fig. 5B), platelet activation (Fig. 5C), and fibrinogen adsorption (Fig. 5D) of HA-F were lower than SS and PDA/HD, which might be due to the anticoagulant effect of 100-kD HA25. Furthermore, the platelet surface coverage (Fig. 5B), platelet activation (Fig. 5C), and fibrinogen adsorption (Fig. 5D) on HA-NCFs with different immersion times showed similar trends as that of platelet adhesion. Although the HA-NCFs showed similar platelet adhesion and fibrinogen adsorption to those on HA-F, their ability to inhibit platelet activation was still stronger than that of the negative control (P < 0.05). After the HA-NCFs were immersed for 1, 3, 5, 10 days, the nanoparticle surface coverage decreased from 100% to 14%. The corresponding anti-coagulant capabilities also decreased gradually. However, the anti-coagulant capabilities were still acceptable. After immersing HA-NCFs for 20 days, there were only 5% remaining nanoparticles, exposing the molecular HA-modified surface. The platelet surface coverage and fibrinogen adsorption were similar to HA-F, even though the platelet activation was still significantly lower than HA-F. Above all, the decrease in anti-coagulant capabilities were closely related to the detachment profiles observed in the previous section. These observations indicate that surface modification with HA-NPs result in strong anti-coagulant effect, which may be caused by the electrostatic repulsion between negatively charged HA-NPs nanoparticles and fibrinogen, which then inhibited platelet adhesion and activation73. Acute thrombosis usually occurs within a few minutes to several hours of stent implantation, whereas subacute thrombosis occurs within a few days to a month, with the highest risk of occurrence during the first week74. Liu et al. have prepared heparin/poly‑L‑Lysine nanoparticles, which provided a favorable release behavior for the healing of vascular Stent Lesions. The heparin/poly‑L‑Lysine nanoparticles on Ti surface showed favorable anticoagulant ability within 1–7 day31. The HA-NP coating exhibited excellent anti-coagulant effect within 1–10 d (a longer anticoagulant time), which met the temporal sequential requirement of anti-coagulation for the vascular stent.

13



Figure 5. (A) Rhodamine fluorescence staining of platelets on samples (B) Platelet surface coverage, (C) Relative platelet activation, and (D) Fibrinogen adsorption on different surfaces. (mean ± SD, N = 3, * indicates significant difference compared to SS and PDA/HD, # indicates significant difference compared to HA-F). Anti-inflammatory Property Macrophages play an important role in inflammatory processes after stent implantation75. Therefore, we studied the effect of HA-NCFs, which were previously immersed in PBS for 0, 1, 3 ,5, 10, 20 days, on the growth behavior of macrophages. Macrophages were cultured on HA-NCFs with SS and PDA/HD as positive controls whereas HA-F was used a negative control. HA-NCFs initially exerted a notable inhibitory effect on macrophage adhesion and spreading. After being soaked in PBS for 1 and 3 days, the nanoparticle surface coverage decreased from 100% to 76%. The surface coverage was relatively high. HA-NCFs still showed markedly better anti-inflammatory activity than the positive and negative controls. With prolonged immersion time for 5, 10 days, the nanoparticle surface coverage markedly decreased from 76% to 14%. The number of adhered macrophages increased gradually, and 14


Page 14 of 32

Page 15 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


some macrophages showed increased adhesion (Fig. 6A). After 20 days of immersion, the remaining HA-NPs gradually fell off. The nanoparticle surface coverage decreased from 14% to 5%, exposing the molecular HA-modified surface. The viability of macrophages on HA-NCFs was similar to that on HA-F. However, it was still significantly lower than that on the positive controls (Fig. 6B). During soaking, the anti-adhesion property of HA-NCFs was gradually weakened as the nanoparticle surface coverage decreased. However, macrophages on HA-NCFs showed increased adhesion when immersed for 5, 10, 20 days, which may have been influenced by the exposed PEI. Podosomes are actin-enriched adhesive structures that permit cell migration and invasion. There was a research indicated that PEI-coated nanoparticles positively modulated podosome formation triggered macrophage activation through TLR-4 signaling44, 76 The electrostatic interaction of shedding HA-NPs between HA and PEI might be disturbed by ions in PBS, leading to the degradation of HA-NPs and PEI exposing. After stent implantation, inflammatory response was observed within a few days to weeks, reaching its highest level in 1–2 days77-78. HA-NCFs showed strong anti-inflammatory ability for up to three days and relatively good anti-inflammatory effect from 5–20 days. Compared with HA-F, HA-NCFs fulfilled the sequential anti-inflammatory requirements of stent function.

Figure 6. (A) Rhodamine fluorescence staining and (B) Activity of macrophages on SS, PDA/HD, HA-F, and HA-NCF previously immersed in PBS for various durations (mean ± SD, N = 3), * indicates significant difference compared to SS and PDA/HD, # indicates significant difference compared to HA-F. Human Umbilical Vein Endothelial Cell (EC) Culture The adhesion behavior of ECs after 4 h of culture on as-fabricated HA-NCFs (0 d) was similar to that on SS and PDA/HA (Fig. 7A). On HA-NCFs pre-treated by PBS immersion, EC adhesion increased with immersion time. The number of ECs that adhered on HA-NCFs (3–10 d) after 4 h of culture was higher than that on SS and PDA/HD. HA-NCFs (20 d) showed the greatest improvement in EC adhesion, which was close to that on HA-F (Fig. 7B). The ECs cultured for 1 day and 3 days showed similar trends in terms of adhesion, spreading, and proliferation as those cultured for 4 h (Fig. 7C–7F). EC proliferation was especially improved on HA-NCFs (10–20 d) after three days of culture, with the formation of complete endothelium on sample 15



surfaces. The above results suggest that the HA-NPs on the HA-NCFs immersed for 0, 1 day suppressed EC adhesion, spreading and proliferation. However, when the immersion time increased, the HA-NPs were detached from the HA-NCF surface, gradually exposing the molecular HA-modified surface. As a result, EC adhesion, spreading, and proliferation were gradually increased. According to the research of Liu et al.14, the heparin/poly-L-lysine nanoparticles modified surface presents negative effects on ECs growth in an early phase. However, it selectively accelerates ECs growth after 10 days. After implantation, the formation of an intact endothelium on the stent surface within the first month is beneficial for inhibition of inflammation, restenosis, and late coagulation. HA-NCFs effectively promoted endothelial growth at 5–20 days.

16


Page 16 of 32

Page 17 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 7. Rhodamine fluorescence staining images of ECs cultured for (A) 4 h, (C) 1 day, and (E) 3 days on HA-NCFs immersed in PBS for various time periods. (B, D, F). Average cell count of ECs cultured for 4 h, 1 day, 3 days on each sample surface (mean ± SD, N = 3). * indicates significant difference compared to SS and PDA/HD, # indicates significant difference compared to HA-F.

17



Human Umbilical Arterial Smooth Muscle Cell (SMC) Culture After the stent implantation, the damaged vessel wall and the activation of the vascular endothelial cells would influence the vascular smooth muscle cells by changing their phenotype from healthy (contractile phenotype) to pathological (synthetic phenotype)79. The pathological hyperplasia of SMC migrated from the injured vessel wall leads to in-stent restenosis, and is the key challenge for the long-term therapy80. According to the HUVSMC isolation method, to facilitate smooth muscle migration from the umbilical artery, the umbilical artery will need to be cut to pieces. The artery pieces would then be cultured in vitro for 10–12 days. In this process, the phenotype of HUVSMC would modulate from contractile phenotype to noncontractile phenotype. Therefore, the state of SMCs used here was likely to be pathological (synthetic phenotype) before seeding81. The morphology of smooth muscle cells has been linked to their phenotype. The elongated, spindle-shaped cells were associated with a contractile state, whereas synthetic SMCs are less elongated and have a cobblestone morphology79. As shown in Fig. 8A, the morphology of SMCs was the normal spindle shape on SS and PDA/HD surface. This is typically associated with a contractile state. The irregular shape on the HA-NCF surfaces might be an indication of a migratory phenotype. The surface coverage of SMCs that adhered on HA-NCFs (0 d) after two days of culture was significantly lower than that on SS, PDA/HD and HA-F. On HA-NCFs pre-treated by PBS immersion, surface coverage of SMCs increased with immersion time. However, the surface coverage of SMCs that adhered on HA-NCFs (1–20 d) after two days of culture was still lower than that on SS, PDA/HD and HA-F. (Fig. 8A and 8B). The above results suggested that the HA-NPs on the HA-NCFs suppressed SMC spreading for 20 days, indicating that the HA-NCFs might help mitigate intimal hyperplasia. This is in line with the temporal requirements of anti-smooth muscle cell hyperplasia after stent implantation.

18


Page 18 of 32

Page 19 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 8. (A) Rhodamine fluorescence staining and (B) Surface coverage of SMCs cultured for two days on HA-NCFs immersed in PBS for various time periods (mean ± SD, N = 3). * indicates significant difference compared to SS and PDA/HD, # indicates significant difference compared to HA-F. Anti-bacterial Properties of HA-NCF Although bacterial infection is rare in the early stage of vascular stent implantation, its development could be fatal7. Therefore, the stent should have anti-bacterial properties at the early stage of implantation. To study the anti-bacterial properties of HA-NCFs, representative gram-negative 19



bacteria Pseudomonas aeruginosa (P.aeruginosa) (Fig. 9A) and gram-positive bacteria Staphylococcus aureus (S.aureus) (Fig. 9B) were selected for cell adhesion testing. Fluorescence images and Cell Counting Kit-8 (CCK-8) showed that the number and activity of P. aeruginosa and S. aureus on HA-NCFs were significantly lower than those on SS, PDA/HD, and HA-F. This indicates that HA-NCFs inhibited bacterial adhesion. In previous studies, HA was able to saturate bacterial hyaluronate lyase, which prevented the bacteria from maintaining elevated levels of tissue permeability, finally inactivating the bacteria82. Furthermore, HA had a good resistance to gram-positive bacteria S. aureus83 and gram-negative bacteria P. aeruginosa84, In addition, PEI had a permeabilization effect on bacterial outer membrane85 and inhibiting growth of bacteria by reducing the metabolic activity of biofilms86. The strong anti-microbial activity of HA-NCFs might come from the synergistic antibacterial effect of HA and PEI. The above results indicated that HA-NCFs had good resistance to gram-positive bacteria S.aureus and gram-negative bacteria P. aeruginosa conferring favorable anti-bacterial function for stent implantation.

Figure 9. Fluorescence microscopy and CCK-8 measurements of (A) P.aeruginosa and (B) S.aureus seeded on various samples (mean ± SD, N = 3)* indicates significant difference compared to SS and PDA/HD, # indicates significant difference compared to HA-F. In Vivo Animal Test of Aortic Implantation in Healthy Sprague-Dawley Rats Thirty days after aortic implantation in healthy Sprague-Dawley rats, the hematoxylin and eosin (HE) staining showed that the hyperplastic area of proliferative tissue on HA-NCFs was significantly smaller than those on SS and HA-F (Fig. 10A, 10B). This may be due to the anti-coagulant, 20


Page 20 of 32

Page 21 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


anti-inflammatory and anti-smooth muscle cell hyperplasia effects of HA-NCFs. As shown in Fig. 10C, the tissues surrounding HA-NCFs were further stained for α-smooth muscle actin (α-SMA), tumor necrosis factor-α (TNF-α), and Von Willebrand factor (VWF). It is known that α-SMA is a typical indicator of contractile smooth muscle. The transition from contractile smooth muscle to synthetic smooth muscle is a key part of hyperplasia formation and restenosis after stent implantation. As shown in Fig. 10A, positive expression of α-SMA suggested that the cells around the HA-NCFs were dominantly contractile SMCs59. However, we cannot rule out that they were myofibroblasts. This is because α-SMA was also expressed in myofibroblasts87. In addition, no positive TNF-α staining was detected in the tissues surrounding the HA-NCFs, indicating that the inflammatory response was weak60. Furthermore, strong VWF expression was observed around the HA-NCFs, signifying the formation of intact ECs61. However, it should be acknowledged that there are differences between the healthy rat models presented here, and pathological rat models presented elsewhere. In the pathological rat models, there was an oxidative stress environment around the pathological tissue having more reactive oxygen free radicals (ROS)69 and hyaluronidase70 than normal physiological condition. Both ROS71and hyaluronidase72 were able to degrade HA, which might accelerate the detachment of HA-NPs from HA-NCF, resulting in a faster loss of bioactivity. It is desirable that HA-NCFs should be further studied in pathological models. Collectively, these results indicated that HA-NCFs effectively reduced inflammation and restenosis and improved endothelial repair in vivo in healthy rat models. This serves as a useful reference for the time-dependent surface biofunctions in vascular stent design.

21



Figure 10. (A) Hyperplastic area of proliferative tissue cross-section of the artery tissues around SS, HA-F, and HA-NCF. * indicates significant difference compared to SS, # indicates significant difference compared to HA-F (mean ± SD, N = 3). (B) Typical optical photographs of the HE-stained cross-sectional slices of the artery tissues around SS, HA-F, HA-NCF implanted after 30 days. (C) α-SMA, TNF-α, and VWF immunohistochemical staining for artery tissues around HA-NCF implanted after 30 days, brown regions represent positive expression. Conclusion

22


Page 22 of 32

Page 23 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 11. Schematic diagram of the time-dependant biofunction of HA-NCFs from 0–20 days in relation to the detachment of HA-NPs. The rise of drug-eluting stents (DES) triggered a revolution in coronary artery stents. By the release of antiproliferative drugs from a stent surface, compared with bare metal stent (BMS), DES apply local pharmacotherapy to inhibit vascular SMC proliferation and thereby contribute to a reduction of 50−70% in the restenosis rate88. However, according to the research of late coronary stent thrombosis of sirolimus-eluting and paclitaxel-eluting stents, the incidence density of stent thrombosis was 1.3% per year, and late stent thrombosis occurred in 40% patients89. Thus, the late thrombosis and restenosis caused by delayed endothelialization continue to restrict the clinical application of DES. As we mentioned in the background, the pathological process of implantation site is sequential, the late thrombosis and restenosis is associated with the dynamic pathological process. To cope with this dynamic pathological process, time-dependent bioactive scaffold might be promising to solve the above problems. In this study, we found that among HA-NPs fabricated by 4-kD, 100-kD, and 750-kD HA, those prepared by 100-kD HA had the best stability in the presence of HAase. Previous studies have shown that hyaluronic acid nanoparticles, as tumor targeting drug carriers, have good stability in the blood. Therefore, the HA-NPs (100 kD) had the potential to be applied as blood contact materials, grafted HA-NPs onto the surface. In vitro experiments indicated that HA-NPs had the characteristics that detached from the surface with immersion time and a time-dependent biological activity, as shown in figure 11. The HA-NCFs showed excellent comprehensive biocompatibility, strongly promoting adhesion and proliferation of ECs while still exerting inhibitory effects on platelets, macrophages, and SMCs. Because the highest risk of thrombosis, inflammation, and hyperplasia occurs within the first week of stent implantation, the time-dependent biofunctions of HA-NCFs conferred prominent inhibitory effects immediately after implantation, which compensated for the insufficient inhibitory effect of HA-F during initial implantation. In vivo test of aortic implantation in healthy Sprague-Dawley rats demonstrated that HA-NCFs possessed good anti-inflammatory properties, inhibited SMCs proliferation, and promoted the formation of a complete endodermis in vivo. Because intact endodermis repair is the key to prevent late coagulation and restenosis, In vivo results suggested that HA-NCFs had the function of inhibiting intimal hyperplasia, restenosis and late coagulation. However, there are some limitations of this research. The release behavior of HA-NPs was under zero flow rate and normal physiological condition. In practical application, fluid shear stress of blood flow and high concentration of ROS and hyaluronidase might accelerate the shedding of HA-NPs and reduce its biological activity. Fluid shear stress generated by blood flow in the vasculature can 23



influence the phenotype of the endothelium by regulating the activity of certain flow-sensitive proteins (for example, CD54), as well as by modulating gene expression. This would finally result in altering the endothelial structure and function. Because leukocytes bind to endothelial cells via CD54, which is also known as ICAM-1 (Intercellular Adhesion Molecule 1). These transmigrate into tissues to regulate inflammation. Thus, the adhesion and proliferation of smooth muscle are also affected90-91. Therefore, fluid shear stress is also an important factor influencing the behavior of vascular cells. The results of the zero-flow cell adhesion test were different from those under fluid shear stress. The animal model used in this study is healthy SD rats. There were differences among species for pathological responses between SD rats and humans. These limitations should be noted. These results further demonstrate that HA-NCF is capable of providing a time-dependent response to the pathological processes at the vessel wall after stent implantation. Although the surface biocompatibility should be evaluated further through long-term in vivo experiments, favorable early biofunction is a precondition for the prevention of middle and late adverse cardiac events. For the biofunctional surface modification of cardiovascular stents, there were some obstacles to make the constructed functional layer to exhibit long-term and direct biological response. The importance of time-dependent surface modification is in (a) providing the time-adjusted remedy and guidance for the intravascular biological response after biomaterial implantation, (b) adjusting the coagulation, inflammation reaction, and intimal regeneration in the normal range, and (c) promoting the formation of monolayer endothelium. This study provides an important basis for the construction of novel HA-NP composite films with favorable time-dependent biofunctions to meet the sequential biological requirements of vascular stent design. The in vitro results suggested that the HA-NCFs exhibited time-dependent biofunctions. Because the highest risk of thrombosis, inflammation, and hyperplasia occur within the first week of stent implantation, the time-dependent biofunctions of HA-NCFs conferred prominent inhibitory effects immediately after implantation, which compensated for the insufficient inhibitory effect of HA-F during initial implantation. Finally, the HA-NCFs demonstrated good anti-inflammatory properties, inhibited SMCs proliferation, and promoted re-endothelialization in vivo. In particular, the time-dependent biofunctions of HA-NCFs resulted in better anti-restenosis effects than those of HA-F, which possessed constant biocompatibility. This study provides an important basis for the construction of novel HA-NP composite films with favorable time-dependent biofunctions to meet the sequential biological requirements of vascular stent design.

Acknowledgement This work was supported by the National Key R&D Program of China (2016YFC1100402) and the National Natural Science Foundation of China (no. 31700821, no. 31870958 and no. 81771988).

Reference 1.

Mohr, F. W.; Morice, M.-C.; Kappetein, A. P.; Feldman, T. E.; Ståhle, E.; Colombo, A.; Mack, M. J.;

Holmes, D. R.; Morel, M.-a.; Dyck, N. V.; Houle, V. M.; Dawkins, K. D.; Serruys, P. W., Coronary artery bypass graft surgery versus percutaneous coronary intervention in patients with three-vessel disease 24


Page 24 of 32

Page 25 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


and left main coronary disease: 5-year follow-up of the randomised, clinical SYNTAX trial. The Lancet 2013, 381 (9867), 629-638. DOI: https://doi.org/10.1016/S0140-6736(13)60141-5. 2.

Airoldi, F.; Colombo, A.; Tavano, D.; Stankovic, G.; Klugmann, S.; Paolillo, V.; Bonizzoni, E.;

Briguori, C.; Carlino, M.; Montorfano, M.; Liistro, F.; Castelli, A.; Ferrari, A.; Sgura, F.; Di Mario, C., Comparison of diamond-like carbon-coated stents versus uncoated stainless steel stents in coronary artery

disease.

The

American

Journal

of

Cardiology

2004,

93

(4),

474-477.

DOI:

https://doi.org/10.1016/j.amjcard.2003.10.048. 3.

Grewe, P. H.; Deneke, T.; Machraoui, A.; Barmeyer, J.; Müller, K. M., Acute and chronic tissue

response to coronary stent implantation: pathologic findings in human specimen. Journal of the American College of Cardiology 2000, 35 (1), 157-163. DOI: 10.1016/s0735-1097(99)00486-6. 4.

Joner, M.; Finn, A. V.; Farb, A.; Mont, E. K.; Kolodgie, F. D.; Ladich, E.; Kutys, R.; Skorija, K.; Gold,

H. K.; Virmani, R., Pathology of Drug-Eluting Stents in Humans. Delayed Healing and Late Thrombotic Risk 2006, 48 (1), 193-202. DOI: 10.1016/j.jacc.2006.03.042. 5.

Joner, M.; Finn, A. V.; Farb, A.; Mont, E. K.; Kolodgie, F. D.; Ladich, E.; Kutys, R.; Skorija, K.; Gold,

H. K.; Virmani, R., Pathology of Drug-Eluting Stents in Humans. Journal of the American College of Cardiology 2006, 48 (1), 193. 6.

Mak, K.-H.; Belli, G.; Ellis, S. G.; Moliterno, D. J., Subacute stent thrombosis: Evolving issues and

current concepts. Journal of the American College of Cardiology 1996, 27 (2), 494-503. DOI: 10.1016/0735-1097(95)00466-1. 7.

Alfonso, F.; Moreno, R.; Vergas, J., Fatal infection after rapamycin eluting coronary stent

implantation. Heart 2005, 91 (6), e51. 8.

Grewe, P. H.; Deneke, T.; Machraoui, A.; Barmeyer, J.; Müller, K.-M., Acute and chronic tissue

response to coronary stent implantation: pathologic findings in human specimen. Journal of the American College of Cardiology 2000, 35 (1), 157. DOI: 10.1016/S0735-1097(99)00486-6. 9.

Farb, A.; Sangiorgi, G.; J. Carter, A.; M. Walley, V.; Edwards, W.; Schwartz, R.; Virmani, R.,

Pathology of Acute and Chronic Coronary Stenting in Humans. 1999; Vol. 99, p 44-52. DOI: 10.1161/01.CIR.99.1.44. 10. Nakazawa, G.; Otsuka, F.; Nakano, M.; Vorpahl, M.; Yazdani, S. K.; Ladich, E.; Kolodgie, F. D.; Finn, A. V.; Virmani, R., The Pathology of Neoatherosclerosis in Human Coronary Implants. Journal of the American College of Cardiology 2011, 57 (11), 1314. DOI: 10.1016/j.jacc.2011.01.011. 11. Rossini, R.; Capodanno, D.; Lettieri, C.; Musumeci, G.; Nijaradze, T.; Romano, M.; Lortkipanidze, N.; Cicorella, N.; Biondi Zoccai, G.; Sirbu, V.; Izzo, A.; Guagliumi, G.; Valsecchi, O.; Gavazzi, A.; Angiolillo, D. J., Prevalence, Predictors, and Long-Term Prognosis of Premature Discontinuation of Oral Antiplatelet Therapy After Drug Eluting Stent Implantation. The American Journal of Cardiology 2011, 107 (2), 186-194. DOI: https://doi.org/10.1016/j.amjcard.2010.08.067. 12. Komatsu, R.; Ueda, M.; Naruko, T.; Kojima, A.; Becker, A. E., Neointimal tissue response at sites of coronary stenting in humans: macroscopic, histological, and immunohistochemical analyses. Circulation 1998, 98 (3), 224-233. DOI: 10.1161/01.cir.98.3.224. 13. Dieter, R. S., Coronary artery stent infection. Clin Cardiol 2000, 23 (11), 808-810. DOI: 10.1002/clc.4960231129. 14. Liu, T.; Zeng, Z.; Liu, Y.; Wang, J.; Maitz, M. F.; Wang, Y.; Liu, S.; Chen, J.; Huang, N., Surface modification with dopamine and heparin/poly-L-lysine nanoparticles provides a favorable release behavior for the healing of vascular stent lesions. ACS applied materials & interfaces 2014, 6 (11), 8729-8743. DOI: 10.1021/am5015309. 25



Page 26 of 32

15. Kim, J. W.; Suh, S. Y.; Choi, C. U.; Na, J. O.; Kim, E. J.; Rha, S.-W.; Park, C. G.; Seo, H. S.; Oh, D. J., Six-Month Comparison of Coronary Endothelial Dysfunction Associated With Sirolimus-Eluting Stent Versus Paclitaxel-Eluting Stent. JACC: Cardiovascular Interventions 2008, 1 (1), 65-71. DOI: 10.1016/j.jcin.2007.11.002. 16. Cassese, S.; Piccolo, R.; Galasso, G.; De Rosa, R.; Piscione, F., Twelve-month clinical outcomes of everolimus-eluting stent as compared to paclitaxel- and sirolimus-eluting stent in patients undergoing percutaneous coronary interventions. A meta-analysis of randomized clinical trials. International Journal of Cardiology 2011, 150 (1), 84-89. DOI: https://doi.org/10.1016/j.ijcard.2011.01.015. 17. Verheye, S.; Agostoni, P.; Dawkins, K. D.; Dens, J.; Rutsch, W.; Carrie, D.; Schofer, J.; Lotan, C.; Dubois, C. L.; Cohen, S. A.; Fitzgerald, P. J.; Lansky, A. J., The GENESIS (Randomized, Multicenter Study of the Pimecrolimus-Eluting and Pimecrolimus/Paclitaxel-Eluting Coronary Stent System in Patients with De Novo Lesions of the Native Coronary Arteries) Trial. JACC: Cardiovascular Interventions 2009, 2 (3), 205-214. DOI: https://doi.org/10.1016/j.jcin.2008.12.011. 18. Windecker, S.; Serruys, P. W.; Wandel, S.; Buszman, P.; Trznadel, S.; Linke, A.; Lenk, K.; Ischinger, T.; Klauss, V.; Eberli, F.; Corti, R.; Wijns, W.; Morice, M.-C.; di Mario, C.; Davies, S.; van Geuns, R.-J.; Eerdmans, P.; van Es, G.-A.; Meier, B.; Jüni, P., Biolimus-eluting stent with biodegradable polymer versus sirolimus-eluting stent with durable polymer for coronary revascularisation (LEADERS): a randomised

non-inferiority

trial.

The

Lancet

2008,

372

(9644),

1163-1173.

DOI:

https://doi.org/10.1016/S0140-6736(08)61244-1. 19. Sorrentino, S.; Giustino, G.; Mehran, R.; Kini, A. S.; Sharma, S. K.; Faggioni, M.; Farhan, S.; Vogel, B.; Indolfi, C.; Dangas, G. D., Everolimus-Eluting Bioresorbable Scaffolds Versus Everolimus-Eluting Metallic Stents. Journal of the American College of Cardiology 2017, 69 (25), 3055-3066. DOI: 10.1016/j.jacc.2017.04.011. 20. Aoki, J.; Serruys, P. W.; van Beusekom, H.; Ong, A. T. L.; McFadden, E. P.; Sianos, G.; van der Giessen, W. J.; Regar, E.; de Feyter, P. J.; Davis, H. R.; Rowland, S.; Kutryk, M. J. B., Endothelial Progenitor Cell Capture by Stents Coated With Antibody Against CD34. The HEALING-FIM (Healthy Endothelial Accelerated Lining Inhibits Neointimal Growth-First In Man) Registry 2005, 45 (10), 1574-1579. DOI: 10.1016/j.jacc.2005.01.048. 21. Melissa, C.; Edgar, T.; Hang, L. C.; Kian Keong, P.; Adrian, L.; Jimmy, L.; Ing Han, L.; Yean Teng, L.; Huay Cheem, T., Use of endothelial progenitor cell capture stent (Genous Bio-Engineered R Stent) during primary percutaneous coronary intervention in acute myocardial infarction: intermediate- to long-term clinical follow-up. American Heart Journal 2008, 155 (1), 128-132. 22. Oxlund, H.; Andreassen, T. T., The roles of hyaluronic acid, collagen and elastin in the mechanical properties of connective tissues. Journal of Anatomy 1980, 131 (Pt 4), 611-620. 23. Kogan, G.; Šoltés, L.; Stern, R.; Gemeiner, P., Hyaluronic acid: a natural biopolymer with a broad range of biomedical and industrial applications. Biotechnology Letters 2007, 29 (1), 17-25. DOI: 10.1007/s10529-006-9219-z. 24. Noble, P. W.; Liang, J.; Jiang, D., Hyaluronan as an Immune Regulator in Human Diseases. Physiological reviews 2011, 91 (1), 221-264. DOI: 10.1152/physrev.00052.2009. 25. Li, J.; Wu, F.; Zhang, K.; He, Z.; Zou, D.; Luo, X.; Fan, Y.; Yang, P.; Zhao, A.; Huang, N., Controlling Molecular Weight of Hyaluronic Acid Conjugated on Amine-rich Surface: Toward Better Multifunctional Biomaterials for Cardiovascular Implants. ACS Applied Materials & Interfaces 2017, 9 (36), 30343-30358. DOI: 10.1021/acsami.7b07444. 26. Wu, F.; Li, J.; Zhang, K.; He, Z.; Yang, P.; Zou, D.; Huang, N., Multifunctional Coating Based on 26


Page 27 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Hyaluronic Acid and Dopamine Conjugate for Potential Application on Surface Modification of Cardiovascular Implanted Devices. ACS Applied Materials & Interfaces 2016, 8 (1), 109-121. DOI: 10.1021/acsami.5b07427. 27. Rayahin, J. E.; Buhrman, J. S.; Zhang, Y.; Koh, T. J.; Gemeinhart, R. A., High and Low Molecular Weight Hyaluronic Acid Differentially Influence Macrophage Activation. ACS Biomaterials Science & Engineering 2015, 1 (7), 481-493. DOI: 10.1021/acsbiomaterials.5b00181. 28. Yoon, H. Y.; Koo, H.; Choi, K. Y.; Lee, S. J.; Kim, K.; Kwon, I. C.; Leary, J. F.; Park, K.; Yuk, S. H.; Park, J. H.; Choi, K., Tumor-targeting hyaluronic acid nanoparticles for photodynamic imaging and therapy. Biomaterials 2012, 33 (15), 3980-3989. DOI: https://doi.org/10.1016/j.biomaterials.2012.02.016. 29. Bonomini, F.; Tengattini, S.; Fabiano, A.; Bianchi, R.; Rezzani, R., Atherosclerosis and oxidative stress. Histol Histopathol 2008, 23 (3), 381-390. DOI: 10.14670/hh-23.381. 30. Raines, E. W., The extracellular matrix can regulate vascular cell migration, proliferation, and survival: relationships to vascular disease. International Journal of Experimental Pathology 2000, 81 (3), 173-182. DOI: doi:10.1046/j.1365-2613.2000.00155.x. 31. Liu, T.; Liu, Y.; Chen, Y.; Liu, S.; Maitz, M. F.; Wang, X.; Zhang, K.; Wang, J.; Wang, Y.; Chen, J.; Huang, N., Immobilization of heparin/poly-l-lysine nanoparticles on dopamine-coated surface to create a heparin density gradient for selective direction of platelet and vascular cells behavior. Acta Biomaterialia 2014, 10 (5), 1940-1954. DOI: https://doi.org/10.1016/j.actbio.2013.12.013. 32. Choi, K. Y.; Chung, H.; Min, K. H.; Yoon, H. Y.; Kim, K.; Park, J. H.; Kwon, I. C.; Jeong, S. Y., Self-assembled hyaluronic acid nanoparticles for active tumor targeting. Biomaterials 2010, 31 (1), 106-114. DOI: https://doi.org/10.1016/j.biomaterials.2009.09.030. 33. Choi, K. Y.; Min, K. H.; Na, J. H.; Choi, K.; Kim, K.; Park, J. H.; Kwon, I. C.; Jeong, S. Y., Self-assembled hyaluronic acid nanoparticles as a potential drug carrier for cancer therapy: synthesis, characterization, and in vivo biodistribution. Journal of Materials Chemistry 2009, 19 (24), 4102-4107. DOI: 10.1039/B900456D. 34. Kim, T.; Lee, H.; Jang, Y.; Gwan Park, T., Controlled Release of Paclitaxel from Heparinized Metal Stent Fabricated by Layer-by-Layer Assembly of Polylysine and Hyaluronic Acid-g-Poly(lactic-co-glycolic acid) Micelles Encapsulating Paclitaxel. 2009; Vol. 10, p 1532-9. DOI: 10.1021/bm900116r. 35. Boussif, O.; Lezoualc'h, F.; Zanta, M. A.; Mergny, M. D.; Scherman, D.; Demeneix, B.; Behr, J. P., A versatile vector for gene and oligonucleotide transfer into cells in culture and in vivo: polyethylenimine. Proceedings of the National Academy of Sciences 1995, 92 (16), 7297-7301. DOI: 10.1073/pnas.92.16.7297. 36. Urban-Klein, B.; Werth, S.; Abuharbeid, S.; Czubayko, F.; Aigner, A., RNAi-mediated gene-targeting through systemic application of polyethylenimine (PEI)-complexed siRNA in vivo. Gene Therapy 2004, 12, 461. DOI: 10.1038/sj.gt.3302425. 37. Behr, J.-P., The Proton Sponge: a Trick to Enter Cells the Viruses Did Not Exploit. CHIMIA International Journal for Chemistry 1997, 51 (1-2), 34-36. 38. Morimoto, K.; Nishikawa, M.; Kawakami, S.; Nakano, T.; Hattori, Y.; Fumoto, S.; Yamashita, F.; Hashida, M., Molecular weight-dependent gene transfection activity of unmodified and galactosylated polyethyleneimine on hepatoma cells and mouse liver. Molecular Therapy 2003, 7 (2), 254-261. DOI: https://doi.org/10.1016/S1525-0016(02)00053-9. 39. Brunot, C.; Ponsonnet, L.; Lagneau, C.; Farge, P.; Picart, C.; Grosgogeat, B., Cytotoxicity of polyethyleneimine (PEI), precursor base layer of polyelectrolyte multilayer films. Biomaterials 2007, 28 (4), 632-640. DOI: https://doi.org/10.1016/j.biomaterials.2006.09.026. 27



Page 28 of 32

40. Fischer, D.; Bieber, T.; Li, Y.; Elsässer, H.-P.; Kissel, T., A Novel Non-Viral Vector for DNA Delivery Based on Low Molecular Weight, Branched Polyethylenimine: Effect of Molecular Weight on Transfection Efficiency and Cytotoxicity. Pharmaceutical Research 1999, 16 (8), 1273-1279. DOI: 10.1023/A:1014861900478. 41. Lakard, S.; Herlem, G.; Propper, A.; Kastner, A.; Michel, G.; Vallès-Villarreal, N.; Gharbi, T.; Fahys, B., Adhesion and proliferation of cells on new polymers modified biomaterials. Bioelectrochemistry 2004, 62 (1), 19-27. DOI: https://doi.org/10.1016/j.bioelechem.2003.09.009. 42. Vleggeert-Lankamp, C. L. A. M.; Pêgo, A. P.; Lakke, E. A. J. F.; Deenen, M.; Marani, E.; Thomeer, R. T. W. M., Adhesion and proliferation of human Schwann cells on adhesive coatings. Biomaterials 2004, 25 (14), 2741-2751. DOI: https://doi.org/10.1016/j.biomaterials.2003.09.067. 43. Moghimi, S. M.; Symonds, P.; Murray, J. C.; Hunter, A. C.; Debska, G.; Szewczyk, A., A two-stage poly(ethylenimine)-mediated cytotoxicity: implications for gene transfer/therapy. Molecular Therapy 2005, 11 (6), 990-995. DOI: https://doi.org/10.1016/j.ymthe.2005.02.010. 44. Mulens-Arias, V.; Rojas, J. M.; Pérez-Yagüe, S.; Morales, M. P.; Barber, D. F., Polyethylenimine-coated SPIONs trigger macrophage activation through TLR-4 signaling and ROS production

and

modulate

podosome

dynamics.

Biomaterials

2015,

52,

494-506.

DOI:

https://doi.org/10.1016/j.biomaterials.2015.02.068. 45. Farrell, L.-L.; Pepin, J.; Kucharski, C.; Lin, X.; Xu, Z.; Uludag, H., A comparison of the effectiveness of cationic polymers poly-l-lysine (PLL) and polyethylenimine (PEI) for non-viral delivery of plasmid DNA to bone marrow stromal cells (BMSC). Eur. J. Pharm. Biopharm. 2007, 65 (3), 388-397. DOI: https://doi.org/10.1016/j.ejpb.2006.11.026. 46. Arnold, L. J.; Dagan, A.; Gutheil, J.; Kaplan, N. O., Antineoplastic activity of poly(L-lysine) with some ascites tumor cells. Proceedings of the National Academy of Sciences 1979, 76 (7), 3246-3250. DOI: 10.1073/pnas.76.7.3246. 47. Wolfert, M. A.; Seymour, L. W., Atomic force microscopic analysis of the influence of the molecular weight of poly(L)lysine on the size of polyelectrolyte complexes formed with DNA. Gene Ther. 1996, 3 (3), 269-273. 48. Jiang, G.; Park, K.; Kim, J.; Kim, K. S.; Hahn, S. K., Target Specific Intracellular Delivery of siRNA/PEI−HA Complex by Receptor Mediated Endocytosis. Molecular Pharmaceutics 2009, 6 (3), 727-737. DOI: 10.1021/mp800176t. 49. Lee, M.-Y.; Park, S.-J.; Park, K.; Kim, K. S.; Lee, H.; Hahn, S. K., Target-Specific Gene Silencing of Layer-by-Layer Assembled Gold–Cysteamine/siRNA/PEI/HA Nanocomplex. ACS Nano 2011, 5 (8), 6138-6147. DOI: 10.1021/nn2017793. 50. Jiang, G.; Park, K.; Kim, J.; Kim, K. S.; Oh, E. J.; Kang, H.; Han, S.-E.; Oh, Y.-K.; Park, T. G.; Kwang Hahn, S., Hyaluronic acid–polyethyleneimine conjugate for target specific intracellular delivery of siRNA. Biopolymers 2008, 89 (7), 635-642. DOI: doi:10.1002/bip.20978. 51. Hongcai, G.; Yimin, S.; Jiajing, Z.; Rong, X.; Hongwei, D., Mussel-Inspired Synthesis of Polydopamine-Functionalized Graphene Hydrogel as Reusable Adsorbents for Water Purification. 2016. DOI: 10.1021/am302500v.s001. 52. Yang, Y.; Qi, P.; Wen, F.; Li, X.; Xia, Q.; Maitz, M. F.; Yang, Z.; Shen, R.; Tu, Q.; Huang, N., Mussel-Inspired One-Step Adherent Coating Rich in Amine Groups for Covalent Immobilization of Heparin: Hemocompatibility, Growth Behaviors of Vascular Cells, and Tissue Response. ACS Applied Materials & Interfaces 2014, 6 (16), 14608-14620. DOI: 10.1021/am503925r. 53. Ito, T.; Yoshihara, C.; Hamada, K.; Koyama, Y., DNA/polyethyleneimine/hyaluronic acid small 28


Page 29 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


complex particles and tumor suppression in mice. Biomaterials 2010, 31 (10), 2912-2918. DOI: https://doi.org/10.1016/j.biomaterials.2009.12.032. 54. Hamerli, P.; Weigel, T.; Groth, T.; Paul, D., Surface properties of and cell adhesion onto allylamine-plasma-coated polyethylenterephtalat membranes. Biomaterials 2003, 24 (22), 3989-3999. DOI: https://doi.org/10.1016/S0142-9612(03)00312-0. 55. Liao, Y.; Li, L.; Chen, J.; Yang, P.; Zhao, A.; Sun, H.; Huang, N., Tailoring of TiO2 films by H2SO4 treatment and UV irradiation to improve anticoagulant ability and endothelial cell compatibility. Colloids

and

Surfaces

B:

Biointerfaces

2017,

155,

314-322.

DOI:

https://doi.org/10.1016/j.colsurfb.2017.04.021. 56. Chen, J.; Yang, P.; Liao, Y.; Wang, J.; Chen, H.; Sun, H.; Huang, N., Effect of the Duration of UV Irradiation on the Anticoagulant Properties of Titanium Dioxide Films. ACS Applied Materials & Interfaces 2015, 7 (7), 4423-4432. DOI: 10.1021/am509006y. 57. Li, G.; Yang, P.; Guo, X.; Huang, N.; Shen, R., An in vitro evaluation of inflammation response of titanium functionalized with heparin/fibronectin complex. Cytokine 2011, 56 (2), 208-217. DOI: https://doi.org/10.1016/j.cyto.2011.06.020. 58. Liu, X.; Chen, J.; Qu, C.; Bo, G.; Jiang, L.; Zhao, H.; Zhang, J.; Lin, Y.; Hua, Y.; Yang, P.; Huang, N.; Yang, Z., A Mussel-Inspired Facile Method to Prepare Multilayer-AgNP-Loaded Contact Lens for Early Treatment of Bacterial and Fungal Keratitis. ACS Biomaterials Science & Engineering 2018, 4 (5), 1568-1579. DOI: 10.1021/acsbiomaterials.7b00977. 59. Sharifpoor, S.; Simmons, C. A.; Labow, R. S.; Paul Santerre, J., Functional characterization of human coronary artery smooth muscle cells under cyclic mechanical strain in a degradable polyurethane

scaffold.

Biomaterials

2011,

32

(21),

4816-4829.

DOI:

https://doi.org/10.1016/j.biomaterials.2011.03.034. 60. Xiao, B.; Laroui, H.; Ayyadurai, S.; Viennois, E.; Charania, M. A.; Zhang, Y.; Merlin, D., Mannosylated bioreducible nanoparticle-mediated macrophage-specific TNF-α RNA interference for IBD

therapy.

Biomaterials

2013,

34

(30),

7471-7482.

DOI:

https://doi.org/10.1016/j.biomaterials.2013.06.008. 61. Jaffe, E. A.; Hoyer, L. W.; Nachman, R. L., Synthesis of von Willebrand Factor by Cultured Human Endothelial Cells. Proceedings of the National Academy of Sciences 1974, 71 (5), 1906-1909. DOI: 10.1073/pnas.71.5.1906. 62. Ganesh, S.; Iyer, A. K.; Morrissey, D. V.; Amiji, M. M., Hyaluronic acid based self-assembling nanosystems for CD44 target mediated siRNA delivery to solid tumors. Biomaterials 2013, 34 (13), 3489-3502. DOI: https://doi.org/10.1016/j.biomaterials.2013.01.077. 63. Han, S.-E.; Kang, H.; Shim, G. Y.; Kim, S. J.; Choi, H.-G.; Kim, J.; Hahn, S. K.; Oh, Y.-K., Cationic derivatives of biocompatible hyaluronic acids for delivery of siRNA and antisense oligonucleotides. Journal of Drug Targeting 2009, 17 (2), 123-132. DOI: 10.1080/10611860802472461. 64. Oyarzun-Ampuero, F. A.; Brea, J.; Loza, M. I.; Torres, D.; Alonso, M. J., Chitosan–hyaluronic acid nanoparticles loaded with heparin for the treatment of asthma. International Journal of Pharmaceutics 2009, 381 (2), 122-129. DOI: https://doi.org/10.1016/j.ijpharm.2009.04.009. 65. Papakonstantinou, E.; Roth, M.; Karakiulakis, G., Hyaluronic acid: A key molecule in skin aging. Dermato-Endocrinology 2012, 4 (3), 253-258. DOI: 10.4161/derm.21923. 66. Lee, G. Y.; Kim, J.-H.; Choi, K. Y.; Yoon, H. Y.; Kim, K.; Kwon, I. C.; Choi, K.; Lee, B.-H.; Park, J. H.; Kim, I.-S., Hyaluronic acid nanoparticles for active targeting atherosclerosis. Biomaterials 2015, 53, 341-348. DOI: 10.1016/j.biomaterials.2015.02.089. 29



Page 30 of 32

67. Xue, P.; Li, Q.; Li, Y.; Sun, L.; Zhang, L.; Xu, Z.; Kang, Y., Surface Modification of Poly(dimethylsiloxane) with Polydopamine and Hyaluronic Acid To Enhance Hemocompatibility for Potential Applications in Medical Implants or Devices. Acs Appl Mater Inter 2017, 9 (39), 33632-33644. DOI: 10.1021/acsami.7b10260. 68. Arima, Y.; Iwata, H., Effect of wettability and surface functional groups on protein adsorption and cell adhesion using well-defined mixed self-assembled monolayers. Biomaterials 2007, 28 (20), 3074-3082. DOI: https://doi.org/10.1016/j.biomaterials.2007.03.013. 69. Juni, R. P.; Duckers, H. J.; Vanhoutte, P. M.; Virmani, R.; Moens, A. L., Oxidative Stress and Pathological Changes After Coronary Artery Interventions. Journal of the American College of Cardiology 2013, 61 (14), 1471-1481. DOI: 10.1016/j.jacc.2012.11.068. 70. Kucur, M.; Karadag, B.; Isman, F. K.; Ataev, Y.; Duman, D.; Karadag, N.; Ongen, Z.; Vural, V. A., Plasma hyaluronidase activity as an indicator of atherosclerosis in patients with coronary artery disease. Bratisl Lek Listy 2009, 110 (1), 21-26. 71. Deguine, V.; Menasche, M.; Ferrari, P.; Fraisse, L.; Pouliquen, Y.; Robert, L., Free radical depolymerization of hyaluronan by maillard reaction products: Role in liquefaction of aging vitreous. International

Journal

of

Biological

Macromolecules

1998,

22

(1),

17-22.

DOI:

https://doi.org/10.1016/S0141-8130(97)00084-6. 72. Harada, H.; Takahashi, M., CD44-dependent intracellular and extracellular catabolism of hyaluronic acid by hyaluronidase-1 and -2. J Biol Chem 2007, 282 (8), 5597-5607. DOI: 10.1074/jbc.m608358200. 73. Mestechkina, N. M.; Shcherbukhin, V. D., Sulfated polysaccharides and their anticoagulant activity: A review. Applied Biochemistry and Microbiology 2010, 46 (3), 267-273. DOI: 10.1134/S000368381003004X. 74. Park, D.-W.; Park, S.-W.; Park, K.-H.; Lee, B.-K.; Kim, Y.-H.; Lee, C. W.; Hong, M.-K.; Kim, J.-J.; Park, S.-J., Frequency of and Risk Factors for Stent Thrombosis After Drug-Eluting Stent Implantation During Long-Term Follow-Up. The American Journal of Cardiology 2006, 98 (3), 352-356. DOI: https://doi.org/10.1016/j.amjcard.2006.02.039. 75. Wilson, G. J.; Nakazawa, G.; Schwartz, R. S.; Huibregtse, B.; Poff, B.; Herbst, T. J.; Baim, D. S.; Virmani, R., Comparison of Inflammatory Response After Implantation of Sirolimus- and Paclitaxel-Eluting Stents in Porcine Coronary Arteries. Circulation 2009, 120 (2), 141-149. DOI: doi:10.1161/CIRCULATIONAHA.107.730010. 76. Chen, H.; Li, P.; Yin, Y.; Cai, X.; Huang, Z.; Chen, J.; Dong, L.; Zhang, J., The promotion of type 1 T helper cell responses to cationic polymers in vivo via toll-like receptor-4 mediated IL-12 secretion. Biomaterials 2010, 31 (32), 8172-8180. DOI: https://doi.org/10.1016/j.biomaterials.2010.07.056. 77. Welt, F. G. P.; Rogers, C., Inflammation and Restenosis in the Stent Era. Arterioscler Thromb Vasc Biol 2002, 22 (11), 1769-1776. DOI: doi:10.1161/01.ATV.0000037100.44766.5B. 78. Li, J.-J.; Fang, C.-H.; Jiang, H.; Huang, C.-X.; Hui, R.-T.; Chen, M.-Z., Time course of inflammatory response after renal artery stenting in patients with atherosclerotic renal stenosis. Clinica Chimica Acta 2004, 350 (1), 115-121. DOI: https://doi.org/10.1016/j.cccn.2004.07.013. 79. Rensen, S. S. M.; Doevendans, P. A. F. M.; van Eys, G. J. J. M., Regulation and characteristics of vascular smooth muscle cell phenotypic diversity. Netherlands Heart Journal 2007, 15 (3), 100-108. DOI: 10.1007/BF03085963. 80. Otsuka, F.; Finn, A. V.; Yazdani, S. K.; Nakano, M.; Kolodgie, F. D.; Virmani, R., The importance of the endothelium in atherothrombosis and coronary stenting. Nature Reviews Cardiology 2012, 9, 439. 30


Page 31 of 32 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


DOI: 10.1038/nrcardio.2012.64 https://www.nature.com/articles/nrcardio.2012.64#supplementary-information. 81. Chamley-Campbell, J.; Campbell, G. R.; Ross, R., The smooth muscle cell in culture. Physiological Reviews 1979, 59 (1), 1-61. DOI: 10.1152/physrev.1979.59.1.1. 82. Carlson, G. A.; Dragoo, J. L.; Samimi, B.; Bruckner, D. A.; Bernard, G. W.; Hedrick, M.; Benhaim, P., Bacteriostatic properties of biomatrices against common orthopaedic pathogens. Biochemical and Biophysical

Research

Communications

2004,

321

(2),

472-478.

DOI:

https://doi.org/10.1016/j.bbrc.2004.06.165. 83. Ardizzoni, A.; Neglia, R. G.; Baschieri, M. C.; Cermelli, C.; Caratozzolo, M.; Righi, E.; Palmieri, B.; Blasi, E., Influence of hyaluronic acid on bacterial and fungal species, including clinically relevant opportunistic pathogens. Journal of Materials Science: Materials in Medicine 2011, 22 (10), 2329. DOI: 10.1007/s10856-011-4408-2. 84. Radaeva, I. F.; Kostina, G. A.; Il'ina, S. G.; Kostyleva, R. N., [Antimicrobial activity of hyaluronic acid]. Zh Mikrobiol Epidemiol Immunobiol 2001,

(1), 74-75.

85. Helander, I. M.; Alakomi, H.-L.; Latva-Kala, K.; Koski, P., Polyethyleneimine is an effective permeabilizer of Gram-negative bacteria. Microbiology 1997, 143 (10), 3193-3199. DOI: doi:10.1099/00221287-143-10-3193. 86. Azevedo, M. M.; Ramalho, P.; Silva, A. P.; Teixeira-Santos, R.; Pina-Vaz, C.; Rodrigues, A. G., Polyethyleneimine and polyethyleneimine-based nanoparticles: novel bacterial and yeast biofilm inhibitors.

Journal

of

Medical

Microbiology

2014,

63

(9),

1167-1173.

DOI:

doi:10.1099/jmm.0.069609-0. 87. Hinz, B.; Phan, S. H.; Thannickal, V. J.; Galli, A.; Bochaton-Piallat, M.-L.; Gabbiani, G., The Myofibroblast: One Function, Multiple Origins. The American Journal of Pathology 2007, 170 (6), 1807-1816. DOI: https://doi.org/10.2353/ajpath.2007.070112. 88. Serruys, P. W.; Kutryk, M. J. B.; Ong, A. T. L., Coronary-Artery Stents. New England Journal of Medicine 2006, 354 (5), 483-495. DOI: 10.1056/NEJMra051091. 89. Daemen, J.; Wenaweser, P.; Tsuchida, K.; Abrecht, L.; Vaina, S.; Morger, C.; Kukreja, N.; Jüni, P.; Sianos, G.; Hellige, G.; van Domburg, R. T.; Hess, O. M.; Boersma, E.; Meier, B.; Windecker, S.; Serruys, P. W., Early and late coronary stent thrombosis of sirolimus-eluting and paclitaxel-eluting stents in routine clinical practice: data from a large two-institutional cohort study. The Lancet 2007, 369 (9562), 667-678. DOI: https://doi.org/10.1016/S0140-6736(07)60314-6. 90. Morigi, M.; Zoja, C.; Figliuzzi, M.; Foppolo, M.; Micheletti, G.; Bontempelli, M.; Saronni, M.; Remuzzi, G.; Remuzzi, A., Fluid shear stress modulates surface expression of adhesion molecules by endothelial cells. Blood 1995, 85 (7), 1696-1703. 91. Dewey, J. C. F.; Bussolari, S. R.; Gimbrone, J. M. A.; Davies, P. F., The Dynamic Response of Vascular Endothelial Cells to Fluid Shear Stress. Journal of Biomechanical Engineering 1981, 103 (3), 177-185. DOI: 10.1115/1.3138276.

31



For Table of Contents Use Only Hyaluronic Acid Nanoparticle Composite Films Confer Favorable Time-Dependent Biofunctions for Vascular Wound Healing

Ting Jiang1,2‡, Zhou Xie2‡, Feng Wu2, Jiang Chen2*, Yuzhen Liao2, Luying Liu2, Ansha Zhao2, Jian Wu1, Ping Yang2*, Nan Huang2

32


Page 32 of 32

Hyaluronic Acid Nanoparticle Composite Films ... - ACS Publications

Recommend Documents