Heparin-Eluting Electrospun Nanofiber Yarns for Antithrombotic

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Heparin-eluting Electrospun Nanofiber Yarns for Antithrombotic Vascular Sutures Sooneon Bae, Michael J. DiBalsi, Nicole Meilinger, Chengqi Zhang, Erica Beal, Guzeliya Korneva, Robert O. Brown, Konstantin G Kornev, and Jeoung Soo Lee ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b14888 • Publication Date (Web): 20 Feb 2018 Downloaded from http://pubs.acs.org on February 21, 2018

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ACS Applied Materials & Interfaces

Heparin-eluting Electrospun Nanofiber Yarns for Antithrombotic Vascular Sutures

Sooneon Bae†,‡,#, Michael J. DiBalsi†, #, Nicole Meilinger†, Chengqi Zhang¶, Erica Beal†, Guzeliya Korneva†, Robert O. Brown§, Konstantin G. Kornev¶, Jeoung Soo Lee†,*

†

Department of Bioengineering, Clemson University, Clemson, South Carolina 29634, USA

‡

Dental and Craniofacial Trauma Research & Tissue Regeneration Directorate, United States Army Institute of Surgical Research, JBSA Fort Sam Houston, Texas 78234, USA

§

Department of Head & Neck Surgery, Greenville Health System, Greenville, SC 29615, USA ¶

Department of Materials Science and Engineering, Clemson University, Clemson, South Carolina 29634, USA

*Corresponding author: Jeoung Soo Lee, Ph. D. Tel.: 864-656-3212 Fax: 864-656-4466 E-mail: [email protected]

KEYWORDS: electrospinning, heparin, cationic amphiphilic copolymer, nanofiber yarn, and vascular suture

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ABSTRACT The surgical connection of blood vessels, anastomosis, is a critical procedure in many reparative, transplantation, and reconstructive surgical procedures. However, effective restoration of circulation is complicated by pathological clotting (thrombosis) or progressive occlusion due to excess cell proliferation that often leads to additional surgeries and increases morbidity and mortality risk for patients. Pharmaceutical agents have been tested to prevent these complications, but many have unacceptable systemic side effects. Therefore, an alternative approach to deliver these drugs at the site of injury in a controlled manner is necessary. The objective of this study was to develop electrospun nanofibers composed of PLGA, PEO, and positively charged copolymer, poly(lactide-co-glycolide)-graft-polyethylenimine (PgP) for electrostatic binding and release of heparin for application as an anti-thrombotic microvascular suture. PgP was synthesized with different coupling ratios between PLGA and bPEI to obtain PgP1 (~1 PLGA grafted to 1 bPEI) and PgP3.7 (~3.7 PLGA grafted to 1 bPEI). Nanofiber yarns (PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7) were fabricated by electrospinning. Heparin immobilization on the positively charged nanofiber yarns was visualized using fluoresceinconjugated heparin (F-Hep) and the amount of immobilized F-Hep was higher on both PLGA/PEO/PgP3.7 and PLGA/PEO/PgP1 than yarns without PgP (PLGA/PEO). We also found that F-Hep was released from both PgP-containing yarns in a sustained manner over 20 days, while over 60% of F-Hep was released within 4 hours from PLGA/PEO. Finally, we observed that heparin-eluting nanofiber yarns with both PgP1 and PgP3.7 showed significantly longer clotting times than nanofiber yarns without PgP. The clotting time of PLGA/PEO/PgP3.7 was not significantly different than that of free heparin (0.5 µg/ml). These results show that heparin-


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eluting electrospun nanofiber yarns may offer a basis for the development of microvascular sutures with anticoagulant activity. 1. INTRODUCTION Sutures are the most widely used medical devices in surgical procedures.1,2 One of the most common applications of sutures is vascular anastomosis, the connection of two tubular structures within the vascular system. Vascular graft implantation and bypass surgery both rely upon suture-based anastomosis. Damage to the vascular endothelium at the site of anastomosis initiates the coagulation cascade, triggers a pro-inflammatory response, and activates smooth muscle cells, leading to major complications such as pathological clots (thrombosis) and progressive occlusion (stenosis).3–6 Thus, there is a need for the design of bioactive sutures capable of inhibiting thrombosis and cellular activation. Another example illustrating the need for antithrombotic activity of sutures is the transplantation of tissue from one location, the “donor site”, to the desired location, the “recipient site”, with suture-based vascular anastomosis. Such free tissue transfer procedures are frequently utilized in a variety of surgical fields including oral maxillofacial, orthopedic, plastic and reconstructive, and otolaryngology.7 Thrombotic occlusion is the main cause of graft failure in free tissue transfers, occurring in 4 percent of free flaps and up to 30 percent of digital replantation. To reduce the risk of these complications, systemic therapy with anti-coagulant drugs is frequently required.3,6,8–10 Despite all these efforts, the challenge to mitigate the risk of thrombosis post-surgery still exists and needs to be addressed.11,12 To develop a suture with improved resistance to thrombosis, we propose to take advantage of the high surface area provided by electrospun nanofibers for binding and release of anti-thrombotic drugs.13–16 Existing technologies allow electrospinning of highly aligned



nanofibers that can be subsequently bundled or twisted to form nanofiber yarns.14,16–20 Since the first report of the fabrication of electrospun nanofibers as a delivery carrier for the model drug tetracycline, many studies have investigated drug delivery from nanofibers. Drug loading can be achieved by 1) dissolving or dispersing the drugs in polymer solutions, resulting in one-phase nanofibers or 2) conjugating the drugs directly onto the polymers for a sustained or stimuliresponsive drug release or 3) using multi-axial needles to develop multi-phase nanofibers.16, 21–24 For example, Lu et al. developed cationized gelatin-coated polycaprolactone fibers by coaxial electrospinning for constructing a core–shell fibrous membrane and immobilized the heparin/VEGF complex through ionic interactions.25 The results demonstrated sustained release of VEGF by electrostatic interaction compared to simple adsorption of drugs on the surface of nanofibers. Recently, several studies have investigated electrospinning for the development of drug-loaded surgical sutures, incorporating antibiotics and anti-coagulant drugs into electrospun fibers in order to deliver them in a controlled manner.26–28 Electrospun nanofibers containing curcumin, a natural turmeric compound known for its antioxidant, anticoagulant, and antibacterial properties, are currently being investigated for soft tissue regeneration and suture applications. The most commonly used anti-coagulant for anastomotic thrombosis prevention is heparin, a negatively charged polysaccharide that binds anti-thrombin III, substantially increasing its anti-thrombotic activity.29,30 Many studies employing heparin-immobilized or heparin-coated biomaterials have been shown to reduce platelet adhesion, increase plasma recalcification time, and increase activated partial thromboplastin time (APTT).30–35 Heparin has also been shown to reduce vascular smooth muscle cell (VSMC) proliferation caused by endothelial injury to the vascular wall which leads to the occurrence of intimal hyperplasia and


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ultimately vascular sclerosis and stenosis.36,37 Several previous studies describe incorporation of anionic heparin into elelctrospun fibers as a means for loading and release of cationic growth factors.38–40 In order to create a vascular suture capable of preventing anastomotic thrombosis, we sought to create electrospun fibers incorporating a cationic polymer for electrostatic binding and local, sustained release of heparin. In this study, we report heparin-eluting electrospun nanofiber yarns for vascular suture application. Nanofibers were electrospun from a solution of the biodegradable polyester poly(lactide-co-glycolide) (PLGA) incorporating a positively charged copolymer, poly(lactideco-glycolide)-g-polyethylenimine (PgP) developed in our lab.41 We hypothesized that the positively charged PgP will provide the capacity for binding negatively charged macromolecules such as heparin via electrostatic interaction (Figure 1). The morphology, surface chemical composition, and mechanical properties of resultant yarns were characterized. Using fluoresceinconjugated heparin (F-Hep), we show increased heparin binding and prolonged release relative to control fibers not containing PgP. The antithrombotic activity of the immobilized/released heparin was shown by increasing the clotting time of bovine platelet-rich plasma in vitro.

Figure 1. Heparin-loaded electrospun nanofiber yarns. Negatively charged heparins are immobilized on the positively charged electrospun nanofiber yarns via electrostatic interactions.



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2. EXPERIMENTAL SECTION 2.1. Materials Poly (lactide-co-glycolide) (PLGA) (50:50, 50 kDa and 50:50, 100 kDa) was purchased from Durect Corporation (Cupertino, CA). Heparin sodium salt from porcine intestinal mucosa, poly(ethylene oxide) (PEO) (1,000 kDa), anhydrous dimethylformamide (DMF), N,Ndimethylacetamide (DMAC), N-hydroxysuccinimide (NHS), N,N'-dicyclocarbodiimide (DCC), 1-ethyl-3-(3-dimethylaminopropryl)carbodiimide

hydrochloride

(EDC),

and

branched

polyethylenimine (bPEI) (25 kDa) were obtained from Sigma-Aldrich (St. Louis, MO). Dialysis tubing (MWCO 50 kDa) was obtained from Spectrum Laboratories, Inc. (Rancho Dominguez, CA). Dimethyl sulfoxide (DMSO) was purchased from Thermo Fisher Scientific (Waltham, MA). PD-10 desalting columns were obtained from GE Healthcare Bio-Sciences Corporation (Piscataway, NJ). 4'-(aminomethyl) fluorescein hydrochloride was obtained from Molecular Probes (Eugene, OR).

2.2. Synthesis and characterization of poly (lactide-co-glycolide)-graft-polyethylenimine (PgP) Two PgPs with different PLGA grafting ratio by using different molar ratio of PLGA and bPEI were synthesized as described in our previous work.41 Briefly, PLGA (50:50, 50 kDa; 1 gm, 20 µmole) was dissolved in 20 ml of anhydrous DMF. NHS (3.5 mg, 30 µmole) and DCC (6.2 mg, 30 µmole) were added to the PLGA solution and the mixture was stirred for 2 hours to


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activate the terminal carboxylic acid group of PLGA. The byproduct dicyclohexylurea was then removed by filteration. To conjugate PLGA to bPEI at a grafting ratio of 1:1, the activated PLGA was added dropwise to bPEI (416.7 mg, 16.7 µmole; PLGA:bPEI =1.2:1 molar ratio) in 10 ml of DMF. To conjugate PLGA to bPEI at a grafting ratio of 4:1, the activated PLGA was added dropwise to bPEI (50 mg, 2.0 µmole; PLGA:bPEI =10:1 molar ratio) in 10 ml of DMF. The mixture was then reacted for 24 hours at room temperature under magnetic stirring at 350 rpm. After reaction overnight, the reactant solution was dialyzed against deionized water (MWCO = 50 kDa) for 3 days, centrifuged at 4,763 g for 10 min to remove unreacted PLGA precipitates, filtered through 0.45 µm membrane, and then lyophilized to obtain PgP. The structure and molecular weight of the two PgPs were determined by 1H NMR (Bruker 300 MHz, DMSO-d6). A designation was given to each PgP based on the PLGA grafting ratio. PgP1 is composed of 1 PLGA grafted to 1 bPEI and PgP3.7 is composed of approximately 3.7 PLGA grafted to 1 bPEI. The structure of PgPs was also characterized by FTIR spectrometer equipped with a Thermo-SpectraTech Foundation Series Endurance Diamond ATR (Thermo-Nicolet Magna 550). Hydrophilicity-lipophilicy balance (HLB) of each PgP was calculated by Griffin’s method as follows: HLB = (Molecular mass of hydrophilic portion / Molecular mass of whole polymer) × 20

Table 1. Compositions of electrospun nanofiber yarns Composition (w/v %)

Yarns

PLGA

PEO

PgP

PLGA/PEO

13.0

2.0

-

PLGA/PEO/PgP1*

13.0

2.0

0.4

PLGA/PEO/PgP3.7*

13.0

2.0

1.5

*Hydrophilic-lipophilic balance (HLB): PgP1 = 6.67 and PgP3.7 = 2.7



2.3. Fabrication of Nanofiber Yarns using Electrospinning The nanofiber yarns were fabricated by using the electrospinning platform as shown in Figure S1.18,42 Polymer solutions were formulated by varying the composition of PLGA, PEO, and PgP, as described in Table 1. Briefly, PLGA (50:50, 100 kDa) was completely dissolved in dimethylacetamide (DMAC) at room temperature. PEO (1,000 kDa) and PgP1 or PgP3.7 were then added and stirred at 60 ˚C overnight. The spinning solution containing PLGA and PEO without PgP was also prepared and used for comparison. After complete dissolution, the polymer solution was transferred in a 10 ml syringe wrapped in a heat sleeve and locked in a programmable syringe pump (NE-300; New Era Pump System, Farmingdale, NY). The positive electrode was applied to a 20-gauge blunt-tip needle (EXEL International Inc., Culver City, CA) and the negative electrode was applied to the custom-made rotating mandrel with four alumina bars at 12 kV using high voltage power supply (Glassman Series EH, High Bridge, NJ). With the applied voltage, polymer mixture was spun with flow rate 0.25–0.30 mL/h, collected onto four alumina bars of the custom-made rotating mandrel placed approximately 20 cm apart from the needle. The nanofiber yarns were formed by collecting the aligned fibers from the mandrel by using a custom-built twisting device composed of two brushes mounted on the ends of the holder that spin in the opposite direction. After twisting the nanofibers, the both ends of the yarn were cut and samples were stored in a desiccator until further use. These yarns are designated to PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7.

2.4. Surface Characterization of Electrospun Nanofiber Yarns


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The surface morphology of the nanofiber yarns was examined with a field emissionscanning electron microscope (FE-SEM; Hitachi S4800, Tokyo, Japan). Samples were mounted onto a SEM stage using a double-sided carbon tape, sputter-coated with palladium/platinum for 60 s, and observed at an accelerating voltage of 12.0 kV and magnifications of 250X and 1000X. To measure the diameter of the yarns, three images of each yarn were taken from the center portions and the two ends using phase contrast microscopy (Axiovert 200, Carl Zeiss, Thornwood, NY) under 10X magnification. Three measurements were obtained from each of the three images and averaged. In addition, the average diameter of nanofibers was determined from three independent yarns’ FE-SEM micrographs using an Image Analysis Measurement tool of Photoshop CS6 (Adobe, San Hose, CA). Six nanofibers from 3 nanofiber yarns’ FE-SEM micrographs were measured and average diameter of yarns expressed as the mean ± standard deviation (SD). To evaluate the surface composition of nanofiber yarns, X-ray photoelectron spectroscopy (XPS) spectra were taken on a Surface Science Instruments S-probe spectrometer in Surface Analysis Recharge Center (University of Washington). This instrument has a monochromatized Al Kα X-ray and a low energy electron flood gun for charge neutralization of non-conducting samples. The samples were fastened to the sample holder with double sided tape and run as insulators. X-ray spot size for these acquisitions was approximately 800 µm. Pressure in the analytical chamber during spectral acquisition was less than 5 x 10-9 Torr. Pass energy for survey and detail scans was 150eV. Pass energy for high resolution scans was 50 eV. The takeoff angle (the angle between the sample normal and the input axis of the energy analyzer) was ~0º (0º take-off angle ≅ 100 Å sampling depth). Service Physics Hawk Data Analysis 7 Software was used calculate surface atomic concentrations using peak areas above a linear background



and elemental sensitivity factors. The binding energy scales of the high-resolution spectra were calibrated by assigning the most intense C1s high-resolution peak a binding energy of 285 eV. One spot was analyzed from the center of each sample. The wettability of the nanofiber yarns was characterized by measuring the contact angles at different pH conditions. Two different pH solutions water (pH 7) and 0.1 N HCl (pH 2) were used as the wetting liquids in these experiments to examine the effect of the H+ on the wettability of the nanofiber yarns. The hypothesis was that if amines were present, as the pH of probe liquid is lowered, an increasing number of surface amines will be protonated, making the surface more polar and reducing the water contact angle. We performed two series of experiments: 1) the glass coverslips were spin-coated with the polymers corresponding to each yarn sample and the contact angle of droplets of water (pH 7) and 0.1 N HCl (pH 2) solutions on the spin-coated polymers was measured using Kruss® Drop Shape Analyzer (DSA 10), and 2) the contact angle on the nanofiber yarns were measured using the procedure developed earlier in Dr. Kornev’s lab.20 Briefly, the nanofiber yarns were vertically immersed into a reservoir filled with the probe liquids water (pH 7) and 0.1 N HCl (pH 2). The contact angle of nanofiber yarns were obtained by imaging the visible meniscus formed on the nanofiber yarns and fitting its profile with the theoretical one. In addition, the fractional area of pores at the surface of nanofiber yarns was measured using Cassie-Baxter equation: cos ߠ௬௔௥௡ = ሺ1 − ߶ሻ cos ߠ௙௜௟௠ + ߶ , where ߠ௬௔௥௡ is the contact angle measured on nanofiber yarns, ߠ௙௜௟௠ is the contact angle measured on films, and ߶ is the surface fraction of pores per total surface area of the yarn.20 The estimations of the fractional area of pores are summarized in Table S1. The surface tension of water (pH 7) and 0.1 N HCl (pH 2) was also evaluated using Kruss® Drop shape analyzer (DSA 10) to allow us for


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interpretation of the contact angle measurements. The obtained surface tension of HCl, 75.2 ± 0.1 mN/m, is greater than the surface tension of water, 72.8 mN/m.

2.5. Mechanical Properties of Nanofiber Yarns In order to characterize the mechanical properties of the nanofiber yarns, the samples were cut into 50 mm segments and mounted in rough-surface aluminum grips using Gator P100 sandpaper (Ali Industries, Inc., Fairborn, OH) to prevent slippage. The separation distance between top and bottom grips was set to a 30 mm. Sample cross-sectional area was calculated based upon measurements of the diameter of the yarn samples, although this is a limitation in these measurements since the yarn diameter includes both all the fiber areas as well as some open space between individual fibers. Using a 100 N load cell, the samples were strained until failure at 5.0 mm/min using a MTS Synergie 100 (MTS Systems Corp., Eden Prairie, MN) and data analyzed using TestWorks® 4 software.

2.6. Preparation of Fluorescein-conjugated Heparin (F-Hep) In order to visualize and quantify the amount of heparin immobilized to the positively charged nanofiber yarns, fluorescein was covalently conjugated to heparin using a modified method from previously described by Osmond et al.43 Briefly, 10% EDC solution in 0.1 M 4morpholinoethanesulfonic acid (MES) buffer (pH 4.7) was added to a 1% heparin solution in 0.1 M MES buffer to activate the carboxylic acid groups of heparins. A 16% 4'-(aminomethyl) fluorescein hydrochloride solution in DMF (approximately 1:10 molar ratio between fluorescein and heparin) was added to the activated heparin solution and stirred at room temperature



overnight. The fluorescein conjugated heparin (F-Hep) was purified using PD-10 desalting columns (GE Healthcare Bio-Sciences Corp., Piscataway, NJ) and lyophilized.

2.7. F-Hep Immobilization and Quantification To visualize and quantitate heparin immobilization, the fluorescein-conjugated heparin (F-Hep) was incubated with nanofiber yarns. Briefly, nanofiber yarns (0.5 mg) were washed three times with deionized water, incubated in 1 mL of 0.1% F-Hep solution in PBS buffer at room temperature for 4 hours with gentle shaking, and then washed with PBS to remove any unbound F-Hep. Electrospun nanofibers without PgP (PLGA/PEO) were used as a control. The nanofiber yarns were observed using an inverted fluorescence microscope (EVOS® FL Cell Imaging System, Bothell, WA) to visualize the bounded F-Hep. The amount of loaded F-Hep on nanofiber yarns was measured using a Synergy™ 4 multi-detection microplate reader (BioTek Instruments, Winooski, VT). The heparin-immobilized nanofiber yarns were dissolved in a microcentrifuge tube containing 200 µL of DMSO using a probe sonicator (Omni Sonic Ruptor 4000, Kennesaw, GA). The loading amount of F-Hep on the yarns was quantified by measuring fluorescence using a Synergy™ 4 multi-detection microplate reader (BioTek Instruments; excitation, 494 nm; emission, 521 nm) and calculated using a standard curve generated from serial dilutions of F-Hep solution. The results were normalized by dividing the amount of F-Hep loaded by each sample’s respective weight post heparin-immobilization and lyophilization and expressed as µg F-Hep per mg yarn. The nanofiber yarn without F-Heparin was used as a negative control.

2.8. In Vitro Release of F-Hep from the Nanofiber Yarns


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F-Hep immobilized nanofiber yarns were incubated in 1 ml of PBS buffer (pH 7.4) at 37 ˚C under the dark condition. At pre-determined time points, 1 ml of PBS (pH 7.4) was removed from each well and replaced with 1 ml of fresh PBS (pH 7.4). The concentration of F-Hep released from nanofiber yarns were measured using a Synergy™ 4 multi-detection microplate reader (BioTek Instruments; excitation, 494 nm; emission, 521 nm) and calculated from standard curves prepared by serial dilutions of F-Hep solution of known concentration as described above.

2.9. Activated Partial Thromboplastin Time (APTT) To

immobilize heparin,

nanofiber

yarns

(PLGA/PEO,

PLGA/PEO/PgP1, and

PLGA/PEO/PgP3.7) were washed three times with deionized water and were incubated in 1 mL of 0.1% heparin solution in PBS buffer at room temperature for 4 hours with gentle shaking, washed with PBS to remove any unbound heparin, and then freeze dried. Bovine platelet-rich plasma (PRP) was prepared using procedure previously described by Clemmons et al.44 Briefly, bovine blood was collected from farm healthy cows into Dry Blood Collection Bag (Jorgensen Laboratories, Inc., Loveland, CO) containing Acid-dextrose-citrate (ACD) anticoagulant (volume ratio of anticoagulant to blood = 1 : 9). The blood was centrifuged at 200 g for 15 min at room temperature and the supernatant, the platelet rich plasma (PRP) was collected. In order to determine the anti-thrombotic activity of heparin loaded on nanofiber yarns, the clotting time of bovine platelet-rich plasma (PRP) treated with heparin loaded PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7 nanofiber yarns (0.5 mg fiber/sample) was measured by activated partial



thromboplastin time (APTT) assay as previously described.45,46 The clotting time of PRP treated with heparin (0.1 and 0.5 µg/ml) was measured and used for comparison. To measure clotting time, 500 µL of PBS (without magnesium and calcium) was added to each well of a sterile 24 well plate. A pre-weighed yarn sample (~ 0.5 mg) or 10 µL of heparin (0.1 and 0.5 µg/ml) in PBS were added to the each well. 500 µL of pre-warmed (37 ˚C) bovine platelet-rich plasma (PRP) was then added to each well and the plate was shaken gently for 5 minutes at room temperature followed by 20 min of incubation at 37 ˚C. 100 µL of 0.1 M calcium chloride (CaCl2) that was prepared in PBS, filtered and warmed to 37˚C was added to each well to activate the coagulation cascade, and the time to the formation of the clots was recorded. A clot was carefully observed when the clear solution started to become cloudy, eventually leading to a gel-like phase. The average clotting time was acquired compared with controls (n = 6). The samples incubated in PBS buffer without heparin were used as a negative control.

2.10. Statistical Analysis The results were analyzed using one-way analysis of variance (ANOVA) followed by Tukey multiple pairwise comparison tests with GraphPad Prism software (GraphPad, San Diego, CA). p < 0.05 was accepted as statistically significant. All quantitative data are expressed as mean ± standard deviation (SD).

3. RESULTS 3.1. Synthesis and Characterization of PgP


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Two different PgPs (PgP1 and PgP3.7) were successfully synthesized using different molar ratios of PLGA and bPEI. The structure of PgP and the grafting ratio of PLGA to bPEI were verified by 1H-NMR (DMSO-d6) (δ = 1.47 (3H, PLGA –CH3), δ = 3.2–3.7 (PEI backbone – CH2CH2–), δ = 4.9 (2H, PLGA –CH2), δ = 5.2 (1H, PLGA –CH)). The ratio of the integrals of the PEI backbone (δ = 3.2–3.7) to the methine of PLGA (δ = 5.2) indicates that approximately 1 PLGA was grafted to 1 bPEI molecule when PgP was synthesized using the molar ratio of PLGA: bPEI (1.2:1) and designated as PgP1 (Figure 2A, Top) and molecular weight of PgP1 was about 75, 000. The ratio of the integrals of the PEI backbone (δ = 3.5) to the methine of PLGA (δ = 5.2) also demonstrated that approximately 3.7 PLGA molecules were grafted to 1 bPEI molecule when PgP was synthesized using the molar ratio of PLGA and bPEI (10:1) and designated as PgP3.7 (Figure 2A, Bottom) and molecular weight of PgP1 was about 75, 000. Hydrophilic and Lipophilic Balance (HLB) values were calculated and for PgP1 and PgP3.7 were 6.67 and 2.7, respectively. Figure 2B shows FT-IR spectra of PLGA, PgP1 and PgP3.7. The FT-IR spectrum of PgP1 and PgP3.7 shows typical amide carbonyl peaks at 1637 cm−1 along with typical ester carbonyl peak (1747 cm−1) originating from PLGA.



Figure 2. Characterization of poly(lactide-co-glycolide)-g-polyethylenimine (PgP) by 1H-NMR and ATR-FTIR. (A) 1H-NMR spectrum of PgP1 (Top) and PgP3.7 (Bottom) using DMSO-d6 as a solvent, and (B) ATR-FTIR spectra of PLGA, PgP1, and PgP3.7.

3.2. Surface Characterization of Electrospun Nanofiber Yarns The fabrication parameters used for the electrospinning of PLGA/PEO/PgP nanofibers were optimized by controlling the flow rate, voltage, needle gauge, and distance between the syringe needle and the mandrel. The compositions of PLGA/PEO nanofibers electrospun with and without PgP are summarized in Table 1. A high molecular weight PEO was used to enhance the viscoelasticity of the spinning solution, resulting in the stable formation of electrospun nanofibers.47,48 We counted the number of twists of the each yarn using hand tally counter for 1 minutes and the average number of twists of yarns was 161.7± 9.1 per minutes. FE-SEM micrographs show uniformly twisted fibrous yarns with the average diameter of 93.00 ± 5.93 µm (PLGA/PEO), 105.56 ± 18.67 µm (PLGA/PEO/PgP1), and 92.74 ± 20.11 µm (PLGA/PEO/PgP3.7), respectively (Figure 3). No significant differences in morphology and diameter of nanofibers yarns among groups were observed. The average diameters of the nanofibers were 48.0 ± 6.38 nm (PLGA/PEO), 51.9 ± 7.93 nm (PLGA/PEO/PgP1), and 43.4 ± 5.11 nm (PLGA/PEO/PgP3.7), respectively (Figure S2).


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Figure 3. FE-SEM micrographs of electrospun nanofiber yarns. PLGA/PEO (A and D), PLGA/PEO/PgP1 (B and E), and PLGA/PEO/PgP3.7 (C and F). Original magnification: 250X, scale bar: 200 µm (A-C); original magnification: 1000X, scale bar: 50 µm (D-F).

The elemental composition of the nanofiber yarn surfaces was characterized by XPS. The presence of PgP on the surface of nanofibers was confirmed by the presence of nitrogen (N 1s) peaks on both PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7, but not on PLGA/PEO (Table 2 and Figure S3). However, no significant difference in the nitrogen (N) content was identified between PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7.

Table 2. Elemental composition of electrospun nanofiber yarns as determined by XPS Yarns

C%

O%

N%

PLGA/PEO

60.2 ± 0.68

39.8 ± 0.68

0

PLGA/PEO/PgP1

62.4 ± 1.75

37.1 ± 1.85

0.5 ± 0.12

PLGA/PEO/PgP3.7

59.6 ± 1.03

40.2 ± 1.06

0.2 ± 0.04



The contact angles were assessed with two different probe solutions water (pH 7) and 0.1 N HCl (pH 2) on the spin-coated films and electrospun nanofiber yarns. The fitting procedure and the result of all measurements are shown in Figure 4 and Figure S4. The contact angles on the PLGA/PEO films were consistent with the surface tension measurements (Figure 4A): the liquid with a greater surface energy, ߪு஼௟ , i.e. an aqueous solution of 0.1 N HCl (pH 2), provides a greater contact angle. The measurements with 0.1 N HCl on PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 films support the hypothesis that the amines on the heparin-loaded films do interact with H+, leading to the increase in the surface polarity. In addition, the PLGA/PEO/PgP3.7 films show smaller contact angle relative to the PLGA/PEO/PgP1 films. The contact angles formed by menisci on electrospun nanofiber yarns in Figure 4B appear significantly lower relative to those on the corresponding films in Figure 4A. As shown in our previous publication,20 this tendency can be explained by the Cassie-Baxter theory of wetting of porous materials: making hydrophilic material porous significantly strengthens the material hydrophilicity, and vice versa, making hydrophobic material porous significantly increases the material hydrophobicity. According to the analysis of contact angles on the films, we would expect the PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 nanofiber yarns to be less wettable by water. Indeed, the results of our experiments confirm this expectation. Analyzing the results for all menisci on the nanofiber yarns, we observed that menisci make the contact angles on PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 yarns slightly greater than those on the PLGA/PEO yarns. However, this difference in the contact angles is very small, suggesting that the available pore space is large and the fraction of the liquid/liquid interface to the total surface area supporting meniscus is large.


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Figure 4. Contact angles on the surface of electrospun nanofiber yarns (A) and spin-coated films (B) at different pH conditions (pH 2 and pH 7). The contact angle of the nanofiber yarns was measured using water (pH 7) and 0.1 N HCl (pH 2) solutions. Data represent the mean ± SD. *p < 0.05 compared with PLGA/PEO at pH 7, †p < 0.05 compared with PLGA/PEO at pH 2, #p < 0.05, ‡p < 0.05.

3.3. Mechanical Properties of Electrospun Nanofiber Yarns We also investigated whether the mechanical properties of the nanofiber yarns containing PgPs were suitable for vascular suture application. The nanofiber yarns containing either PgP1 (PLGA/PEO/PgP1) or PgP3.7 (PLGA/PEO/PgP3.7) showed a lower elastic modulus than the nonPgP nanofiber yarns, while the PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 nanofiber yarns showed no significant difference in mechanical properties (Table 3). The elongation at break of nanofiber yarns was significantly elevated in the PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 nanofiber yarns, while it is not significantly elevated in the PLGA/PEO nanofiber yarns. Interestingly, the addition of PgP1 to the nanofiber yarns resulted in a significant increase in the tensile strength, while no significant differences were observed on the PLGA/PEO and PLGA/PEO/PgP3.7 nanofiber yarns. Table 3. Mechanical properties of electrospun nanofiber yarns 19 ACS Paragon Plus Environment


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Diameter (µm)

Elastic Modulus (MPa)

Elongation to Failure (%)

Tensile Strength (MPa)

93.00 ± 5.93

666.99 ± 52.85

29.0 ± 1.48

10.93 ± 3.07

PLGA/PEO/PgP1

105.56 ±18.67

487.55 ± 31.82

78.44 ± 5.71

23.34 ± 4.97

PLGA/PEO/PgP3.7

92.74 ± 20.11

453.25 ± 31.95

67.2 ± 18.4

13.4 ± 4.37

Yarns PLGA/PEO

3.4. F-Hep Loading efficiency on Nanofiber Yarns Figure 5 shows that negatively charged heparin can bind to electrospun nanofiber yarns containing a positively charged PgP via electrostatic interaction. Both PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 nanofiber yarns showed significantly higher loading amount of F-Hep (µg/mg fiber) than the PLGA/PEO nanofiber yarns (Figure 5A). The amount of F-Hep loaded on PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7 was 1.03 ± 0.06, 1.63 ± 0.21, and 3.21 ± 0.12 µg/mg yarn, respectively. Figure 5B shows representative fluorescence images of F-Hep loaded nanofiber yarns. It must be noted that PLGA/PEO/PgP3.7 had a stronger fluorescence intensity than PLGA/PEO/PgP1 and PLGA/PEO, indicating that the increased amount of F-Hep was loaded in PLGA/PEO/PgP3.7 nanofiber yarns. These results demonstrate that a higher amount of PgP leads to an increased heparin conjugation to the nanofiber yarns.

Figure 5. F-Hep immobilized nanofiber yarns. (A) The amount of F-Hep loaded on nanofiber 20 ACS Paragon Plus Environment

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yarns (n = 3 samples per group). Data represent the mean ± SD. *p < 0.05 compared with PLGA/PEO, †p < 0.05 compared with PLGA/PEO/PgP1. (B) Visualization of immobilized FHep on nanofiber yarns by fluorescent microscopy. Original magnification: 100X.

3.5. Release of F-Hep from Electrospun Nanofiber Yarns In vitro release profiles of F-Hep from nanofiber yarns were assessed after incubation in PBS buffer (pH 7.4) for 20 days. Figure 6A shows cumulative F-Hep release during the first 8 hours. We observed a burst release of F-Hep from PLGA/PEO nanofiber yarn due to a weak interaction between heparin and PLGA/PEO yarns, while F-Hep exhibited much slower, sustained release in both PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 nanofiber yarns due to a relatively strong electrostatic interaction between negatively charged heparin and positively charged yarns. The % cumulative F-Hep release from PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7 was 55.2%, 40.2, and 22.1% at 2 hour incubation, and 72.4%, 54.1%, and 27.8% released at 8 hour incubation, respectively. Moreover, PLGA/PEO/PgP3.7 nanofiber showed a controlled F-Hep release over 20 days, while F-Hep release reached a maximum within 4 days from both PLGA/PEO and PLGA/PEO/PgP1 nanofiber yarns (Figure 6B). The % release of F-Hep from PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7 was 91.8%, 70.6%, and 64.8% at 20 days, respectively. In addition, FE-SEM micrographs confirmed that the surface morphology of nanofiber yarns after the release test for 20 days was preserved (Figure 6C).



Figure 6. Cumulative F-Hep release in vitro from nanofiber yarns for the first 8 h (A) and for 20 days (B) (n = 3 samples per group). Data represent the mean ± SD. (C) FE-SEM micrographs of nanofiber yarns at 20 days post-release (original magnifications: 250X and 1000X; scale bars: 200 µm and 50 µm).

3.6. Antithrombotic Activity of Heparin-immobilized nanofiber yarns In order to determine the anti-thrombotic activity of heparin that was loaded on nanofiber yarns, the clotting time of bovine platelet-rich plasma (PRP) in the presence of PLGA/PEO, PLGA/PEO/PgP1, and PLGA/PEO/PgP3.7 nanofiber yarns was measured by activated partial thromboplastin time (APTT) assay.44 The PLGA/PEO yarns were used as controls and heparin (0.1 and 0.5 µg/ml) solutions used as positive controls. The clotting time with the 22 ACS Paragon Plus Environment

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PLGA/PEO/PgP3.7 nanofiber yarns was significantly longer than that with the PLGA/PEO/PgP1 and PLGA/PEO (Figure 7). The PLGA/PEO/PgP3.7 nanofiber yarns exhibited significantly extended clotting time compared with the PRP samples treated with heparin solutions (0.1 and 0.5 µg/ml), which correspond to the amount of immobilized heparin on electrospun nanofiber yarns. The clotting time of bovine platelet-rich plasma (PRP) treated with PLGA/PEO/PgP3.7, PLGA/PEO/PgP1, and PLGA/PEO was 231.7 ± 8.14, 187.8 ± 6.18, and 169.5 ± 7.05 sec, respectively. In addition, PRP’s clotting times treated with heparin solutions (0.1 and 0.5 µg/ml) were 186.0 ± 3.61 and 216.3 ± 12.66 sec, respectively.

Figure 7. Anti-thrombotic activity of heparin immobilized nanofiber yarns. Clotting time was measured by activated partial thromboplastin time (APTT) assay (n = 6 samples per group). Data represent the mean ± SD. *p < 0.05 compared with PBS w/o heparin (Hep), †p < 0.05 compared with heparin (0.1 µg/ml), NS: not significant.

4. DISCUSSION AND CONCLUSIONS Anastomotic thrombosis, the most frequent complication in vascular surgery, often results in stenosis of the vessel, leading to the graft failure and revision surgery or can cause an



embolism, leading to life-threatening ischemia.2,49 Numerous pharmaceutical agents have been tested to reduce thrombus formation, and ultimately prevent these high-risk complications, but the results have been disappointing. The agents either have undesirable systemic effects or they cannot achieve a controlled, timed release in the circulation. With an expectation that, a negatively charged anti-thrombotic heparin will bind to the positively charged PgP and provide a sustained and controlled release of heparin, one would like to have a suture containing both compounds to prevent thrombosis. The two different electrospun yarns, PLGA/PEO/PgP1 and PLGA/PEO/PgP3.7 with the different PgP1 and PgP3.7, illustrate the effect of the compound hydrophobicity. As shown in Table 1, the higher hydrophobicity of PgP3.7 (HLB = 2.7) than PgP1 (HLB = 6.67) allowed 1.5% PgP3.7 composition in the PLGA/PEO/PgP3.7 yarn compared to 0.4% PgP1 composition in PLGA/PEO/PgP1. These results are correlated with the amount of heparin loaded on the yarns as we found that the more PgP in the yarns can provide a more positive charge allowing to bind more heparin (Table 1 and Figure 3). In addition, the electrospun nanofiber yarns appeared highly uniform with smooth surface (Figure 3). The yarns have about 100 µm diameter which is suitable for applications as a microvascular suture. Commercial sutures that are typically used for microvascular anastomosis are USP 6-0 sized but are also still small enough for vascular applications elsewhere in the body.50 One advantage of fabricating yarns by electrospinning technology is that the diameter of the yarns can be easily adjusted by increasing or decreasing the spinning time. One limitation we observed is that the mechanical strength of the nanofiber yarns was less compared to a commercially available similar diameter suture.51–53 This deficiency can be addressed by adding some other stronger polymers into the spinning dope.19


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In our heparin loading study, fluorescein was conjugated to the heparin for visualization. Previously, Luong-Van et al. developed heparin-releasing poly (Ɛ-caprolactone) electrospun fiber meshes by dissolving heparin directly into the polymer solution and electrospinning onto flat collection plate. They reported that the loading amount of heparin was about 0.51 µg and 5.8 µg per mg fiber by using 0.05 and 0.5% (w/v) heparin in polymer solution, respectively. Interestingly, we were also able to achieve similar loading amount of F-Hep via electrostatic interaction

37

. In our F-Hep release study, incorporation of a positively charged PgP in

PLGA/PEO nanofiber yarns resulted in a slower, more sustained release of heparin compared to the yarns without PgP. The PLGA/PEO nanofiber yarns without PgP showed a burst initial release within 4 hours (Figure 6). We observed that the PLGA/PEO/PgP3.7 nanofiber yarns containing a higher amount of PgP released heparin in a much more controlled and sustained manner compared to PLGA/PEO/PgP1 nanofiber yarns. This supports the hypothesis that the release rate of negatively charged heparin was predominantly controlled by the amount of positively charged PgP in the nanofiber yarns due to electrostatic interaction. Su et al. prepared core-shell structured electrospun nanofibers loaded with heparin in the core of nanofibers and reported that the release mechanism of heparin from electrospun mats was by Fickian diffusion and erosion.37 They achieved a sustained heparin release from the nanofibers and the relative release rate of heparin from polymer was faster when the nanofibrous mat was loaded with less heparin. However, they also observed that nanofiber mats loaded with higher amounts of heparin lost their original morphology after 14 day release study. In addition, the morphology of both PLGA/PEO/PgP electrospun nanofiber yarns used in our study was not altered after subjecting them to a release study for 20 days (Figure 6). We think that the release mechanism of heparin from PgP



containing PLGA/PEO nanofiber yarns is diffusion, but dissociation of heparin from the nanofiber surfaces. We believe that this can be also an important feature for clinical applications of these yarns as the drug-loaded vascular sutures. In the antithrombotic activity test of immobilized heparin on the nanofiber yarns, we observed that all heparin immobilized onto yarns showed significantly longer clotting time than PBS without heparin. Hep-PLGA/PEO/PgP3.7 showed significantly longer clotting time than Hep-PLGA/PEO/PgP1 or Hep-PLGA/PEO (nanofiber yarns without PgP). The clotting time of Hep-PLGA/PEO/PgP3.7 was longer than that that of free heparin at concentrations (0.1 and 0.5 µg/ml). Interestingly, our anti-thrombotic activity data indicate that clotting of plasma in the presence of free heparin (0.5 µg) were required approximately 4 minutes. Based on the heparin loading study with F-Hep, approximately 1.75 µg heparin were immobilized on the PLGA/PEO/PgP3.7 (0.5 mg yarn). The amount of heparin released from yarns for 4 minutes can be 0.32 µg (HepPLGA/PEO), 0.036 µg (Hep-PLGA/PEO/PgP1), ~0.04 µg (Hep-PLGA/PEO/PgP3.7) per mg yarns, suggesting that heparin may be still bound to the yarn surface and the bound heparin may still be as active as free heparin. This antithrombotic activity of the bound heparin along with the promising sustained and controlled release profile favor the feasibility of these heparin loaded PLGA/PEO/PgP3.7 electrospun nanofiber yarns as the potential vascular sutures. In summary, we successfully synthesized the positively charged copolymers, poly (lactide-co-glycolide)-graft-polyethylenimine (PgP) with different coupling ratios between PLGA and bPEI. The PLGA/PEO/PgP nanofibers were fabricated by electrospinning and positively charged nanofiber yarns were formed by collecting the aligned fibers from the mandrel and by using a custom-built twisting device. Negatively charged heparin was


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immobilized onto positively charged nanofiber yarns allowing a slow release of heparin from yarns in a controlled manner. The heparin loaded nanofiber yarns with PgP showed significantly longer clotting times compared to the yarns without PgP. These results show that the heparinimmobilized nanofiber yarns could be utilized in vascular suture application with effective anticoagulant activity.

ASSOCIATED CONTENT Supporting Information The electrospinning platform used for fabrication of nanofiber yarns; fractional pore area and elemental composition measurements in the surface of electrospun nanofiber yarns; and average diameters of electrospun nanofibers. The Supporting Information is available free of charge on the ACS Publications website at DOI:.

AUTHOR INFORMATION Corresponding Author *E-mail: [email protected] Author Contributions #

S.B. and M.J.D. contributed equally to this work. S.B., M.J.D., G.K., K.G.K., R.O.B., and J.S.L.

designed the experiments. S.B., M.J.D., N.M., C.Z., E.B., and G.K. performed the experiments. All authors analyzed and reviewed the data, contributed to manuscript preparation, and approved the final version of the manuscript.

Notes



The authors declare no competing financial interest.

ACKNOWLEDGEMENTS We would like to thank the COBRE Bioengineering and Bioimaging Core for access to electrospinning platform and MTS Synergie 100. We also thank the Electron Microscopy Facility (EMF), Division of Research at Clemson University for access to FE-SEM. XPS analysis was conducted by Mr. Gerry Hammer at the Molecular Analysis Facility, a National Nanotechnology Coordinated Infrastructure site at the University of Washington which is supported in part by the National Science Foundation (Grant No. ECC-1542101), the University of Washington, the Molecular Engineering & Sciences Institute, the Clean Energy Institute, and the National Institutes of Health. This work was supported by the National Institute of General Medical Sciences (NIGMS) of the National Institutes of Health (NIH) (Grant No. 5P20GM103444-07) through the SCBioCRAFT COBRE Center of Regeneration and Formation of Tissues.


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GRAPHIC ABSTRACT


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Heparin-Eluting Electrospun Nanofiber Yarns for Antithrombotic

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