Zwitterionic Polymer-Grafted Polylactic Acid Vascular Patches Based

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Bio-interactions and Biocompatibility

Zwitterionic Polymer-Grafted Polylactic Acid Vascular Patches Based on Decellularized Scaffold for Tissue Engineering Jun Zhang, Lei He, Guo Wei, Xuefeng Jiang, Lei Fu, Yue Zhao, Luxia Zhang, Lutao Yang, Yajuan Li, Yutong Wang, Hong Mo, and Jian Shen ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/ acsbiomaterials.9b00684 • Publication Date (Web): 25 Jul 2019 Downloaded from pubs.acs.org on July 25, 2019

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ACS Biomaterials Science & Engineering

Zwitterionic Polymer-Grafted Polylactic Acid Vascular Patches Based on Decellularized Scaffold for Tissue Engineering Jun Zhang,†,§ Lei He,†,§ Guo Wei,† Xuefeng Jiang,† Lei Fu,† Yue Zhao,† Luxia Zhang,† Lutao Yang,† Yajuan Li,† Yutong Wang,‡ Hong Mo,†,* Jian Shen†,*

†Jiangsu

Collaborative Innovation Center of Biomedical Functional Materials, National

and Local Joint Engineering Research Center of Biomedical Functional Materials, Jiangsu Engineering Research Center for Biomedical Function Materials, School of Chemistry and Materials Science, Nanjing Normal University, Wenyuan Road #1, Xianlin University Town, Qixia District, Nanjing 210023, Jiangsu Province, China ‡College

of Materials Science and Engineering, Nanjing Forestry University, Longpan

Road #159, Xuanwu District, Nanjing 210037, Jiangsu Province, China

§

These authors contributed equally to this work.

*Authors to whom all correspondence should be addressed E-mail Address: [email protected] or [email protected] Tel: +86-25-83598031, Fax: +86-25-83716813

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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 More than ten million people suffer from cardiovascular diseases, and diseased blood vessels need to be treated with vascular patches. For a vascular patch, good affinity for endothelial progenitor cells is a key factor to promote the formation of endothelial tissue – endothelialization. To construct such a vascular patch with good cell affnity, in this work, we first synthesized a reactive zwitterionic organophosphate containing a phosphorylcholine

head

group:

6-(acryloyloxy)hexyl-2-(N-isopropyl-N,N-

dimethylammonio)ethyl phosphate (AHEP). We then grafted AHEP onto a polylactic acid (PLA)-coated decellularized scaffold to obtain a vascular patch. Its hydrophilicity and biocompatibility were investigated. Its in vivo performance was also examined in a pig model with B-ultrasonography, Doppler spectrum and computed tomography angiography. The vascular patch demonstrated nonhemolytic property, noncytotoxicity, long in vitro coagulation times, strong ability to resist platelet adhesion, and good affinity for endothelial progenitor cells. The vascular patch was able to maintain the long-term patency (5 months) of the surgical arteries. Hence, the zwitterionic polymer-grafted PLA vascular patch may be a promising candidate for vascular tissue engineering.

Keywords: phosphorylcholine, zwitterionic polymer, polylactic acid, decellularized scaffold, vascular patch

2


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 INTRODUCTION Cardiovascular diseases are still the main cause of death globally, and more than ten million people suffer from these life-threatening diseases.1–3 Diseased blood vessels need to be treated with vascular patches.4–8 Undoubtedly, autologous vascular patches are the optimal choice, due to their non-immunogenicity and identical mechanical property with recipient’s blood vessel. However, the use of autologous patches may be limited in terms of quality and source, due to patient’s medical complications such as disease and previous surgeries. Synthetic vascular patches, based on materials such as expanded poly(tetrafluoroethylene) and poly(ethylene terephthalate), are likely to induce postsurgical failure since they are prone to calcification, inflammatory response and thrombus formation.4,9 Allogeneic and xenogeneic vascular patches are from different donors and thus have high availability. After decellularization, the cellular components in allogeneic and xenogeneic tissues are removed while extracellular matrices (ECMs) are retained. The decellularized products are called “decellularized scaffolds” (DCSs). Hence, DCSs are non-immunogenic due to the removal of cellular antigens. DCSs also have high mechanical similarity to native tissues due to the retention of vascular microarchitecture, which can provide mechanical support for regenerating tissue after in vivo implantation. In addition, the retained ECMs in the DCS can constitute a physiological microenvironment, deliver biochemical and cell signals, and guide cell attachment and growth.10–14 Because of these characteristics, DCSs have attracted considerable attention in tissue engineering applications.15–18 Nevertheless, DCSs are porous and 3



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cannot be used as vascular patches directly. Generally, before use, a DCS needs coating with a polymeric film, such as polyurethane (PU) or polylactic acid (PLA) films.2,5 PLA is a “double green” polymer because it is derived from renewable sources and can degrade into natural products. Due to its high biocompatibility and good in vivo degradability as well as easy processability, PLA is widely used in biomedical applications, such as hard tissue repair and drug carriers.19–21 However, the limitedly hydrophilic surface of PLA may induce non-specific protein adsorption and bacterial adhesion (namely, fouling), which can cause thrombus formation and surgical-site infection and thus lead to postsurgical failure.22–25 An effective approach to enhancing surface hydrophilicity and fouling-resistance is to introduce zwitterionic polymers.26–32 Among

zwitterionic

polymers,

polymers

containing

phosphorylcholine

(2-

(trimethylammonio)ethyl phosphate, Figure 1) head groups have received a great deal of attention, because the polymers have similar headgroup to that of phospholipid phosphatidylcholine in the membranes of all eukaryotic and most other cells, and thus they are considered biocompatible.5,33–35 For example, Kobayashi et al. grafted 2(methacryloyloxy)ethyl-2-(N,N,N-trimethylammonio)ethyl phosphate (also called 2(methacryloyloxy)ethyl phosphorylcholine, MPC, Figure 1) onto polyethylene (PE) via surface-initiated polymerization to obtain a superhydrophilic material with the water contact angle of 6°, which significantly increased the pristine material’s hydrophilicity.32 Similarly, Xu et al. prepared a poly(MPC)-grafted silicone film via ozone-induced polymerization. The film demonstrates enhanced hydrophilicity and resistance to platelet adhesion.33 The enhancement in hydrophilicity and antifouling 4


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ability is caused by the tight hydration water layer which is formed via electrostatic interactions and hydrogen bond due to the physical and chemical characteristics of the phosphorylcholine head group.5,31

+

O O

N

P O O

phosphorylcholine head group: 2-(trimethylammonio)ethyl phosphate O Hydrophilic head O

Hydrophobic tail

O

O

O

+

N

P O

O O a representative phospholipid phosphatidylcholine: 3-(tetradecyloxy)-2-(6,8dioctadecyloxy)propyl-2-(N,N,N-trimethylammonio)ethyl phosphate

+

O

O O

P

N O

O

O 2-(methacryloyloxy)ethyl-2-(N,N,N-trimethylammonio)ethyl phosphate MPC

+

O

O O

P

N O

O

O

6-(acryloyloxy)hexyl-2-(N-isopripy-N,N-dimethylammoino)ethyl phosphate AHEP Figure 1. zwitterionic organophosphates containing a phosphorylcholine head group

However, too strong hydration ability severely inhibits cell attachment and proliferation on zwitterionic surfaces, hindering tissue regeneration.22 Therefore, 5



regulating the hydrophilicity-hydrophobicity of zwitterionic surfaces is of great importance in promoting cell adhesion and growth as well as tissue regeneration. A variety of methods have been used to mediate the hydrophilicity-hydrophobicity, such as by grafting different content of poly(MPC) onto surfaces, by introducing chiral molecules, and by changing spacer length between charged groups.33,36–39 For phosphatidylcholines, they usually have a hydrophilic head and a hydrophobic tail (Figure 1). For example, 3-(tetradecyloxy)propyl-2-(6,8-dioctadecyloxy)-2-(N,N,Ntrimethylammonio)ethyl phosphate is hydrophobic since they contain two long alkyl groups in the tail.40 MPC is highly hydrophilic due to the short 2-carbon linear alkyl group between ester group and phosphate group in the tail. Hence, on the surface of a vascular patch, the hydrophilicity-hydrophobicity may also be tuned by changing the chain-length of the linear alkyl group between ester group and phosphate group in zwitterionic compounds containing a phosphorylcholine head group. Such an approach has not been reported yet in the literature. In this work, we synthesized a reactive zwitterionic organophosphate containing a phosphorylcholine

head

group:

6-(acryloyloxy)hexyl-2-(N-isopropyl-N,N-

dimethylammonio)ethyl phosphate (AHEP) with

a 6-carbon linear alkyl group

between ester group and phosphate group (Figure 1). AHEP was grafted onto a PLAcoated DCS (PLA/DCS) to form a vascular patch PLA-g-poly(AHEP)/DCS (Figure 2). Its zwitterionic surface hydrophilicity and affinity for endothelial progenitor cells (EPCs) were investigated. Its in vivo performance was also examined in a pig model with B-ultrasonography, Doppler spectra and computed tomography angiography 6


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(CTA). The zwitterionic surface demonstrated the high anti-platelet adhesion ability and the good affinity for EPCs, and the patch could maintain the long-term patency (5 months) of the surgical arteries. Thus, the zwitterionic polymer-grafted patch PLA-gpoly(AHEP)/DCS may have great potential in tissue engineering applications.

Figure 2. Scheme of the construction of a zwitterionic polymer-grafted PLA vascular patch and its in vivo implantation in a pig model

 EXPERIMENTAL SECTION Materials. Carotid arteries of 6-month-old beagle dogs, EPCs, red blood cells (RBCs), whole blood, and platelet-poor plasma (PPP) were provided by Gulou hospital (Nanjing, China). Cyanmethemoglobin reagent, triethylamine (TEA), tetrahydrofuran (THF), acetonitrile, methanol, diethyl ether, N,N-dimethylformamide (DMF), hexane1,6-diol, acryloyl chloride, 2-chloro-2-oxo-1,3,2-dioxaphospholane (COP), benzoyl peroxide, acetone, and PLA was purchased from Sigma-Aldrich (Shanghai, China). Water was deionized. Phosphate-buffered saline (PBS, pH = 7.4) was Mg- and Ca-free. All biomaterials were sterilized with 75% ethanol before use. 7



Synthesis of a zwitterionic organophosphate containing a phosphorylcholine head group. In synthesis experiments, unless noted, all materials were reagent grade and used without further purification. All reactions and distillation were performed under dry N2. Before use, TEA was distilled from sodium hydroxide, THF from sodium and benzophenone, and acetonitrile from calcium hydride. Glassware was oven-dried, assembled hot, and cooled under dry N2 prior to use. Liquids were transferred by standard syringe techniques. Chromatographic separations were performed in standard column methods with silica gel (230–400 mesh). AHEP was synthesized via 3 steps (Figure 3). Step 1: Synthesis of 2-(6-hydroxy)hexyloxy-2-oxo-1,3,2-dioxaphospholane (HOP). HOP was synthesized in the reported method.41 Hexane-1,6-diol (6.49 g, 0.055 mol) and TEA (5.05 g, 0.05 mol) were added to a flask containing THF (75 mL). The mixture was stirred and cooled to -20 °C in an ice-salt bath. COP (7.13 g in 10 mL of THF) was added dropwise over 2 h and the mixture was allowed to stir for 1 h at -20 °C. The mixture was removed from the ice-salt bath and allowed to warm to room temperature. After stirred for 1 h, the mixture was allowed to stand for 2 h and filtered. The filtrate was concentrated by rotary evaporation to give a yellow liquid (HOP). HOP was used for the following reaction without further purification. 1H NMR (400 MHz, CDCl3) (Figure S1): δ = 1.44 ppm (m, 4H, PO(CH2)2CH2CH2(CH2)2OH), 1.60 (m, 2H, PO(CH2)4CH2CH2OH), 1.73 (m, 2H, POCH2CH2(CH2)4OH), 3.23 (s, 1H, PO(CH2)6OH), 3.67 (m, 2H, PO(CH2)5CH2OH), 3.83 (m, 2H, POCH2(CH2)5OH), 4.11 (m, 4H, OCH2CH2O). 8


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Step 2: Synthesis of (6-hydroxy)hexyl-2-(N,N-dimethylisopropylamino)ethyl phosphate (HEP). HEP was synthesized in the reported method.41 HOP (5.60 g, 0.025 mol), acetonitrile (50 mL) and N,N-dimethylisopropylamine (1.65 g, 0.03 mol) were added to a stainless steel cylinder (8 cm  15 cm). The cylinder was sealed and then shaken in a shaking incubator at 70 °C for 3 d. The mixture was transferred to a roundbottom flask. The solvent was removed by rotary evaporation to give a yellow crude product. The crude product was dissolved in methanol (20 mL). Diethyl ether (50 mL) was added for precipitation. The precipitate was removed by filtration. The filtrate was concentrated by rotary evaporation to give a light green liquid (HEP). HEP was used for the following reaction without further purification. 1H NMR (400 MHz, D2O) (Figure S2): δ = 1.14 ppm (d, 4H, HO(CH2)2CH2CH2(CH2)2OP, J = 6.68 Hz), 1.23 (m, 6H,

NCH(CH3)2),

1.40

(m,

2H,

HOCH2CH2(CH2)4OP),

1.51

(m,

2H,

HO(CH2)4CH2CH2OP), 2.63 (s, 3H, NCH3), 2.88 (s, 1H, NCH(CH3)2), 3.43 (m, 3H, NCH3), 3.59 (m, 2H, HOCH2(CH2)5OP), 3.63 (m, 2H, HO(CH2)5CH2OP), 3.76 (m, 2H, NCH2CH2O), 3.99 (m, 2H, NCH2CH2O). Step 3: Synthesis of AHEP. HEP (5.82 g, 0.02 mol) and TEA (3.03 g, 0.03 mol) were added to a flask containing acetonitrile (75 mL). The mixture was stirred and cooled to -20 °C in an ice-salt bath. Acryloyl chloride (2.85 g, 0.03 mol) in acetonitrile (10 mL) was added dropwise over 30 min. The mixture was removed from the ice-salt bath and allowed to warm to room temperature. After stirred for 2 h, the mixture was filtered. The filtrate was concentrated by rotary evaporation to give a yellow liquid (AHEP). AHEP was purified by column chromatography (6% EtOAc and 1% methanol 9



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in hexane). FT-IR (KBr, cm-1) (Figure S5): 729 (s, v–(CH2)6–), 1026 (s, vPO–CH2), 1202 (s, vC–O), 1265 (s, vP–O), 1637 (s, vC=C), 1730 (s, vC=O). 1H NMR (400 MHz, D2O) (Figure S3): δ = 1.22 ppm (m, 4H, O(CH2)2CH2CH2(CH2)2OP), 1.32 (m, 6H, NCH(CH3)2), 1.48 (m, 2H, OCH2CH2(CH2)4OP), 1.60 (m, 2H, O(CH2)4CH2CH2OP), 2.71 (s, 3H, NCH3), 2.93 (s, 1H, NCH(CH3)2), 3.12 (m, 3H, NCH3), 3.52 (m, 2H, OCH2(CH2)5OP), 3.82 (m, 2H, O(CH2)5CH2OP), 4.03 (m, 2H, OCH2CH2N), 4.13 (m, 2H, OCH2CH2N), 5.90 (m, 1H, H2C=CH), 6.12 (m, 1H, cis-H, H2C=CH), 6.36 (m, 1H, trans-H, H2C=CH).

13C

NMR (100 MHz, D2O) (Figure S4): δ = 13.6 ppm (NCH(CH3)2), 24.5 (O(CH2)3CH2(CH2)2OP), 24.7 (O(CH2)2CH2(CH2)3OP), 29.6 (OCH2CH2(CH2)4OP), 31.1 (O(CH2)4CH2CH2OP), 58.1 (N(CH3)2), 61.1 (OCH2CH2N), 61.7 (OCH2CH2N), 65.4 (NCH(CH3)2), 66.5 (OCH2(CH2)5OP), 66.9 (O(CH2)5CH2OP), 127.8 (H2C=CH), 132.3 (H2C=CH), 168.8 (H2C=CHC). HRMS for C16H33NO6P [M+H]+, 366.2040 (calculated), 366.2068 (measured), (Figure S6). Preparation of PLA/DCS. PLA (10.0 g) was dissolved in DMF (100 mL) to give a solution. This solution was coated onto a DCS by dip-coating three times to give a composite film PLA/DCS. The film was dried in vacuum at 60 °C for 24 h. Preparation of zwitterionic polymer-grafted polylactic acid vascular patch PLA-g-poly(AHEP)/DCS. Graft polymerization was performed according to the method reported in the literature.42,43 PLA/DCS was placed into a flask. Under dry N2, benzoyl peroxide (BPO) in acetone (5.0×10-2 M, 20 mL) was added. The mixture was refluxed for 30 min and then heated to 70 °C. AHEP in water (0.25 M, 50 mL) was added dropwise over 1 h. The mixture was stirred for 2.5 h at 70 °C to give PLA-g10


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poly(AHEP)/DCS. PLA-g-poly(AHEP)/DCS was washed alternately with CH3OH and water, sonicated, and dried in vacuum at 37 °C. Scanning electron microscopy (SEM). Surface morphologies were observed with JSM-7600F SEM (Japan). Water contact angle measurements. Water contact angles were measured with DSA100 optical contact angle tester (Germany) at 25 °C. The test solution was deionized water with a droplet size of 90 μm. In vitro coagulation time tests. The coagulation times of each patch were tested according to the method reported by our group.8 Plasma recalcification time tests. The plasma recalcification time of each patch was tested according to the method reported by our group.8 Platelet adhesion tests. The experiments of platelet adhesion to each patch were carried out according to the method reported by our group.8 Cytotoxicity evaluation. The cytotoxicity of each patch was evaluated according to the method reported by our group.8 Cell attachment and proliferation tests. EPCs were used to evaluate the attachment and proliferation on the zwitterionic surface of the patch according to the method reported by our group.8 In vivo implantation study. PLA/DCS and PLA-g-poly(AHEP)/DCS was used for in vivo implantation study according to the method reported by our group.8 Statistical analysis. Each test was performed in triplicate except additional statement. All data were expressed as mean ± standard deviation (SD). Statistical 11



significance between groups was analyzed with one way Analysis of Variance (ANOVA), followed by the Bonferroni post hoc test. The value of p less than 0.05 was considered statistically significant.

 RESULTS Synthesis and characterization of AHEP. Zwitterionic AHEP was synthesized according to the scheme presented in Figure 3. The molecular structure was characterized by FT-IR, 1H NMR,

13C

NMR and HRMS. In the FT-IR spectrum

(Figure S5), the peak at 729 cm-1 was attributed to –(CH2)6–, and those at 1026 and 1265 cm-1 to –OPO–CH2– and –OPO–, respectively. The peaks at 1637 and 1730 cm-1 were assigned to –C=C– and ester carbonyl, respectively. The 1H NMR spectrum is shown in Figure S3. The peaks at 5.90, 6.12 and 6.36 were attributed to three protons in the ethylenic bond –CH2=CH–. The 13C NMR spectrum is shown in Figure S4, the peaks at 127.8 and 132.3 were ascribed to two carbons in the ethylenic bond –CH2=CH–. These data were consistent with those reported in the literature.44–47 In addition, the HRMS showed a peak for [M+H]+ (AHEP with a proton) at 366.2068 (Figure S6), which was consistent with the calculated weight (366.2040). This indicated that zwitterionic AHEP was successfully synthesized.

12


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Cl

O P O

O

O(CH2)6OH

O HO(CH2)6OH, TEA THF

P O

N CH3CN

O

HOP

COP

O

O HO(CH2)6OPO(CH2)2N(CH3)2 O

O

H2C CHCCl TEA, CH3CN

H2C

CHCO(CH2)6OPO(CH2)2N(CH3)2

CH(CH3)2

O

HEP

O

CH(CH3)2

AHEP Figure 3. Scheme of the synthesis of AHEP

Graft ratio, molecular weight distribution, average graft length, and reaction mechanism of PLA-g-poly(AHEP). For convenience, when we tested the graft ratio and the molecular weight distribution as well as the average graft length, we did not coat the film onto a DCS. The graft ratio of PLA-g-poly(AHEP) was 17.6% when it was calculated according to the film weights (Supporting Information, section 8 “Graft Ratio”). The graft ratio was 17.3% when it was calculated according to the average molecular weights. These two data were similar. As presented in Table S2, the polydispersion index of PLA-g-poly(AHEP) (2.0) was comparable to that of PLA (2.1). The average graft length of PLA-g-poly(AHEP) was 7.0 (Supporting Information, section 10 “Average graft length”). A proposed reaction mechanism of the AHEP grafting is shown in Figure 4. BPO decomposed into two primary free radicals while heating. The primary free radical incorporated with the hydrogen of tertiary carbon atom in PLA to generate a PLA free radical and benzoic acid. At the same time, the primary free radical also reacted with AHEP to form a monomeric free radical. The monomer free radical continuously reacted with AHEP to form a long-chain free radical. Finally, the long-chain free radical combined with the PLA free radical to terminate the 13



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reaction. O CO

O

O

2

OC O

O CO

O

CH3

CH3

+

O

+ H 2C

O

n

O O

+

O

-

H2 C

CO

CH(CH3)2

+

n

O O

CHCO(CH2)6OPO(CH2)2N(CH3)2 O

O

O

O

O CO

CO

C OH

CH

O

+

(CH3)2 N (CH2)2OPO(CH2)6OC

-

O CO

+

H2 C

O

O

(H3C)2HC O CH

O

+

m-1 H2C

O

+

-

O

-

O

O

CH(CH3)2

(CH3)2 N (CH2)2OPO(CH2)6OC (H3C)2HC

O

+

CHCO(CH2)6OPO(CH2)2N(CH3)2 O

CH3

O

O

n

CH3 O

O

O O

O

H2 C

CO

+

CH

O

n

m

(CH3)2 N (CH2)2OPO(CH2)6OC (H3C)2HC

-

O

O

Figure 4. The reaction mechanism of AHEP grafting

XPS spectra and thickness of PLA/DCS and PLA-g-poly(AHEP)/DCS. The XPS spectra of PLA/DCS and PLA-g-poly(AHEP)/DCS from 127.5 to 140 eV at binding energy are shown in Figure S7. PLA/DCS did not display the P2p peak whereas PLAg-poly(AHEP)/DCS showed it at 134 eV since PLA does not contain phosphorus whereas AHEP contains phosphorus. The value of the P2p peak for AHEP was consistent with those for sodium triphosphate (133.6 eV) and phosphatidylcholine (134 eV).2,5 This indicated that AHEP was successfully grafted onto PLA/DCS. The thickness of DCS, PLA/DCS and PLA-g-poly(AHEP)/DCS was 0.95±0.08, 1.07±0.06 and 1.07±0.08 mm. Thus, the thickness of the PLA coating onto the DCS was 0.12±0.14 mm (Table S4). 14


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Thermal property. TGA curves of PLA/DCS and PLA-g-poly(AHEP)/DCS are shown in Figure S8. The decomposition temperatures of PLA/DCS and PLA-gpoly(AHEP)/DCS were 276 and 274 °C, respectively. This suggested that the graft polymerization of AHEP did not affect the thermal stability of the patch obviously. In our previous work, we modified the PU film with phosphatidylcholine-PU nanoparticles before the PU film was coated onto the DCS. The thermal stability of the PU film is not significantly affected by the surface modification with phosphatidylcholine-PU

nanoparticles,

since

the

materials’

decomposition

temperatures before and after the surface modification are 239.7 and 208.9 °C, respectively.5 Mechanical properties. The tensile strength of DCS, PLA/DCS and PLA-gpoly(AHEP)/DCS is shown in Figure S9A. After PLA was coated onto the DCS (namely, PLA/DCS), the material’s tensile strength substantially increased from 27.2 to 45.1 MPa, which might be caused by the stiffness of PLA. In contrast, the graft polymerization of AHEP did not have significant influence on the tensile strength of PLA/DCS, since the tensile strength before and after the graft polymerization were 45.1 and 44.9 MPa, respectively. The elongation at break of DCS, PLA/DCS and PLA-gpoly(AHEP)/DCS is shown in Figure S9B. After PLA coating, the materials’ elongation at break slightly increased from 434.8% to 482.3%. The graft polymerization of AHEP did not affect the elongation at break of PLA/DCS, since the elongation at break of the grafted patch PLA-g-poly(AHEP)/DCS was still 482.3%. This indicated that PLA coating and AHEP grafting did not compromise the mechanical 15



properties of the pristine material DCS. Water contact angle analysis. As shown in Figure 5, the water contact angle of PLA/DCS was 80°, suggesting that PLA was hardly hydrophilic. By contrast, the water contact angle of PLA-g-poly(AHEP)/DCS was 40°, indicating that the hydrophilicity of the patch was significantly enhanced due to the graft polymerization of zwitterionic AHEP onto the surface.

Figure 5. Water contact angles: (A) PLA/DCS, (B) PLA-g-poly(AHEP)/DCS. The results are expressed as mean ± SD, n = 3 (p < 0.05).

Hemolysis and RBC morphologies. The hemolysis rates of PLA/DCS and PLA-gpoly(AHEP)/DCS were 1.86% and 1.03%, respectively (Table S3). According to ASTM F756-17, a material is considered nonhemolytic if its hemolysis rate is less than 2%. Therefore, PLA/DCS and PLA-g-poly(AHEP)/DCS were both nonhemolytic, and the blood compatibility of the patch was further improved by introducing zwitterionic AHEP. Additionally, the morphologies of RBCs exposed to normal saline, PLA/DCS and PLA-g-poly(AHEP)/DCS are shown in Figure S10. The RBCs exposed to PLA/DCS and PLA-g-poly(AHEP)/DCS presented the same morphology as those exposed to normal saline. This suggested that both PLA/DCS and PLA-gpoly(AHEP)/DCS did not damage RBCs. 16


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Plasma recalcification time and in vitro coagulation times. The plasma recalcification times of PLA/DCS and PLA-g-poly(AHEP)/DCS were 12.1 min and 22.3 min, respectively (Figure 6A). The plasma recalcification time was substantially prolonged due to the graft polymerization of zwitterionic AHEP. Similarly, APTT, TT and PT of PLA/DCS were 80 s, 15 s and 14 s respectively, while those of PLA-gpoly(AHEP)/DCS were 90 s, 25 s and 22 s respectively (Figure 6B–D). All the in vitro coagulation times were prolonged due to the introduction of zwitterionic poly(AHEP). These findings indicated that the blood compatibility of the path was improved through the graft polymerization of AHEP. 40

25

A

B

PRT

20 15 10

20

10

5 0

APTT

30

Time (s)

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trol Con

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trol Con

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Figure 6. (A) Plasma recalcification time (PRT); the in vitro coagulation times: (B) APTT, (C) TT, (D) PT. The results are expressed as mean ± SD, n = 3 (p < 0.05).

Platelet adhesion. A large number of platelets adhered to PLA/DCS, and some were even deformed and stretched their pseudopods (Figure 7A). In contrast, few platelets adhered to PLA-g-poly(AHEP)/DCS (Figure 7B). This suggested that the zwitterionic poly(AHEP)-grafted surface possessed high resistance to platelet adhesion.

Figure 7. SEM images (×2000) of platelet adhesion: (A) PLA/DCS, (B) PLA-g-poly(AHEP)/DCS

Cytotoxicity. The viabilities of EPCs incubated in the extracts of PLA/DCS and PLA-g-poly(AHEP)/DCS are shown in Figure 8. At 6 h, the cell viability (97%) for 18


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PLA-g-poly(AHEP)/DCS was higher than that for PLA/DCS (92%). At 12 h, the cell viability (96%) for PLA-g-poly(AHEP)/DCS was also higher than that for PLA/DCS (87%). Similarly, at 24 h, the cell viability (91%) for PLA-g-poly(AHEP)/DCS was higher than that for PLA/DCS (78%). These data indicated that the introduction of zwitterionic poly(AHEP) improved the cell compatibility of the patch. Additionally, although the cell viabilities for both PLA/DCS and PLA-g-poly(AHEP)/DCS decreased with the incubation time, the cell viability for PLA/DCS dropped more than that for PLA-g-poly(AHEP)/DCS. This also indicated that PLA-g-poly(AHEP)/DCS had better cell compatibility than PLA/DCS. 6h 12h 24h

120 100 Cell viability (%)

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80 60 40 20 0

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PLA-g-poly(AHEP)/DCS

Figure 8. The cell viabilities in the extracts of PLA/DCS and PLA-g-poly(AHEP)/DCS. The results are expressed as mean ± SD, n = 3 (p < 0.05).

Cell attachment and proliferation. As shown in Figure 9, on PLA-gpoly(AHEP)/DCS, the cells increased with the incubation time. All the cells presented normal morphology, and no apoptotic cells were observed. This indicated that the zwitterionic poly(AHEP)-grafted surface favored the attachment and growth of EPCs.

19



Figure 9. Fluorescence microscopy images of cell attachment and proliferation on PLA-gpoly(AHEP)/DCS: (A) 4 h, (B) 8 h, (C) 12 h, (D) 48 h

In vivo implantation. The B ultrasound images and the Doppler spectra for PLA/DCS and PLA-g-poly(AHEP)/DCS at 2 weeks of the in vivo implantation are shown in Figure 10. For PLA/DCS, the surgical site did not display red or orange in the B ultrasound image (Figure 10A), indicating that blood flow did not circulate through the surgical artery, and this artery was blocked. In addition, almost no signals were detected in the Doppler spectrum (Figure 10B), which indicated that PLA/DCS was unable to maintain the 2-week patency of the surgical artery after the in vivo implantation. By contrast, for PLA-g-poly(AHEP)/DCS, the surgical site showed red and orange in the B ultrasound image at 2-week implantation (Figure 10C), suggesting that blood flow circulated through the surgical artery. Additionally, its Doppler spectrum at 2-week implantation exhibited regular and periodic signals (Figure 10D), which also proved that blood flow circulated in this artery normally. This indicated that 20


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PLA-g-poly(AHEP)/DCS was able to maintain the patency of the surgical artery at 2week implantation.

Figure 10. At 2-week implantation, the B-ultrasound image (A) and the Doppler spectrum (B) for PLA/DCS, the B-ultrasound image (C) and the Doppler spectrum (D) for PLA-gpoly(AHEP)/DCS.

The CTA images for PLA/DCS and PLA-g-poly(AHEP)/DCS at 5-month implantation are shown in Figure 11. For PLA/DCS, the CTA image demonstrated that the surgical artery is occluded and no blood flew through it (red arrow, Figure 11A). In contrast, for PLA-g-poly(AHEP)/DCS, the blood flow circulating in the surgical artery (the right branch) was similar to the normal artery (the left branch) (Figure 11B), and no vessel stenosis and occlusion were observed. These findings indicated that PLAg-poly(AHEP)/DCS could maintain the long-term patency (5 months) of the surgical 21



artery whereas PLA/DCS blocked the surgical artery even at 2-week implantation.

Figure 11. At 5-month implantation, CTA images: (A) PLA/DCS, (B) PLA-g-poly(AHEP)/DCS

 DISCUSSION DCSs including carotid arteries, livers, lungs and small intestines are increasingly used for regenerating tissues and organs.2,5,15–19 DCSs, in which structural components are preserved, can provide geometry support and guide tissue remodeling. For DCSs that are derived from blood vessels, since their microarchitecture is retained, they have similar mechanical properties to native blood vessels, such as adequate compliance and burst pressure, and thus they can withstand in vivo dynamic blood flow impact.2,5,17,18 In this work, the DCS was made from the carotid artery of a 6-month-old beagle dog, and PLA coating and AHEP grafting did not impair its mechanical property. Hence, the DCS was a qualified biological scaffold. In addition, the reserved ECMs in the DCS, upon in vivo implantation, constitute a 3D microenvironment with tissue-specific composition and architecture, mediating the biochemical and biophysical cues of cells.14 Since the ECM complexity cannot be readily imitated with synthetic materials, DCSs have received more and more attention in tissue engineering and preclinical 22


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studies. The surface hydrophilicity of a material has significant influence on its biomedical applications. A hydrophobic or low hydrophilic surface may easily induce non-specific protein adsorption and bacterial colonization, and lead to subsequent thrombus formation and inflammatory responses.22–25 PLA surface is hardly hydrophilic, with the water contact angle of 80°. Therefore, further hydrophilicization is necessary for PLA surface to resist fouling and prevent thrombus formation and inflammatory responses.29–31 In this work, by grafting AHEP onto the surface of PLA/DCS, the hydrophilicity of the patch was substantially improved, with the water contact angle decreasing from 80° to 40°. MPC is commonly used to prepare superhydrophilic brushes with the water contact angle of less than 10°.29,32 However, for biomaterials, the surfaces with too strong hydrophilicity or hydration ability may also inhibit the adhesion and proliferation of tissue-specific cells, and thus hinder tissue regeneration.22 AHEP contains a 6-carbon linear alkyl group between ester group and phosphate group, which is longer than the 2-carbon linear alkyl group in MPC. Hence, the patch grafted with zwitterionic poly(AHEP) had appropriate hydrophilicity and enhanced affinity for EPCs while inhibiting platelet adhesion (Figures 7 and 9). By contrast, the unmodified PLA/DCS induced the adhesion of a large number of platelets, demonstrating low antifouling ability. EPCs, which are a subset of stem cells, can home to diseased and damaged sites of blood vessels, and differentiate into mature endothelial cells, thereby forming endothelia and maintaining the patency of blood vessels. However, EPCs circulate in 23



bloodstream at relatively low concentration. Therefore, the high EPC affinity of vascular patches can contribute to the formation of endothelial tissue – endothelialization.48 To this end, molecules or substances such as sodium triphosphate and phosphatidylcholine have been introduced onto vascular patch surfaces to fish out EPCs from bloodstream.2,5 In this work, AHEP graft polymerization provided biomembrane mimicry for the patch, since the headgroup of AHEP resembles that of phospholipid in cell membranes.5,33–35 Hence, PLA-g-poly(AHEP)/DCS had good affinity for EPCs, promoted in situ endothelialization and maintained the long-term patency of the surgical artery (Figure 11B). In contrast, the unmodified PLA/DCS resulted in postsurgical failure 2 weeks after in vivo implantation (Figures 10A and 10B). In general, when a biomaterial is implanted, non-specific proteins in the host tissue adsorb onto the material surface, which may trigger inflammatory response due to the invasion of a foreign object. Macrophages view the protein-adsorbed biomaterial as a foreign object, subsequently adhere to its surface, and try to ingest it but fail since it is large. Hence, the macrophages merge together, forming a giant foreign body cell. This cell secretes inflammatory cytokines and causes fibroblast attachment.24,49–51 In this work, for PLA-g-poly(AHEP)/DCS, the xenotransplantation from a dog to a pig did not cause immune rejection or inflammatory response, which was attributed to the removal of cellular antigens in the DCS and the high antifouling ability of the zwitterionic surface. In this case, the zwitterionic surface effectively inhibited non-specific protein adsorption and bacterial adhesion, thus preventing inflammatory response. In addition, 24


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pigs are large animals and have greater genetic and physiologic similarities to humans, such as disease mechanism and biological responses. Therefore, pig models can more accurately reflect pathogenesis than mouse models, which helps accelerate the translation from preclinical studies to clinical practices.52–54

 CONCLUSIONS In the patch PLA-g-poly(AHEP)/DCS, the DCS provided geometry support for the generative tissue and transmitted biomechanical and biochemical cues due to the retention of structural components and ECMs, and the poly(AHEP)-grafted zwitterionic surface effectively prevented thrombus formation and promoted endothelialization due to the strong ability to resist platelet adhesion and the good affinity for EPCs. As a result, the patch could maintain the long-term patency (5 months) of the surgical arteries. Hence, the zwitterionic polymer-grafted PLA vascular patch may have potential applications in tissue engineering.

 ASSOCIATED CONTENT Supporting Information The Supporting Information is available free of charge on the ACS Publications website at DOI: Preparation of a decellularized scaffold (Page S-1); 1H NMR spectrum of HOP (Figure S1, Page S-2); 1H NMR spectrum of HEP (Figure S2, Page S-3); 1H NMR spectrum of AHEP (Figure S3, Page S-3); 13C NMR spectrum of AHEP (Figure S4, 25



Page S-4); FT-IR spectrum of AHEP (Figure S5, Page S-4); HRMS of AHEP (Figure S6, Page S-5); Graft ratio (Table S1, Page S-6); Molecular weight distribution (Table S2, Page S-6); Average graft length (Page S-7); X-ray photoelectron spectroscopy (XPS) (Figure S7, Page S-7); Thermogravimetric analysis (TGA) (Figure S8, Page S-8); Mechanical properties (Figure S9, Page S-9); Hemolysis assays (Table S3, Page S-9); Red blood cell morphologies (Figure S10, Page S-11); Thickness measurements of DCS, PLA/DCS and PLA-g-poly(AHEP)/DCS (Table S4, Page S11).

 AUTHOR INFORMATION Corresponding Authors *E-mail: [email protected]. Telephone: +86-25-83598031. Fax: +86-25-83716813. *E-mail: [email protected] ORCID Lei He: 0000-0003-1959-2491 Hong Mo: 0000-0002-6374-3632 Author Contributions §

J.Z. and L.H.: These authors contributed equally to this work.

Notes The authors declare no competing financial interest.

 ACKNOWLEDGEMENTS The authors are grateful for the financial support of the National Natural Science 26


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Foundation of China (21604041) and for the assistance provided by Dr. Cheng Liu and Prof. Tong Qiao from Gulou Hospital (Nanjing, China).

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Zwitterionic Polymer-Grafted Polylactic Acid Vascular Patches Based

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