Enhancing anti-thrombogenicity of biodegradable polyurethanes through drug molecule incorporation

Cancan Xu ab, Aneetta E. Kuriakose ab, Danh Truong ab, Primana Punnakitikashem ab, Kytai T. Nguyen ab and Yi Hong *ab
aDepartment of Bioengineering, University of Texas at Arlington, Arlington, TX 76019, USA. E-mail: yihong@uta.edu; Fax: +1-817-272-2251; Tel: +1-817-272-0562
bJoint Biomedical Engineering Program, University of Texas Southwestern Medical Center, Dallas, TX 75390, USA

Received 16th June 2018 , Accepted 7th August 2018

First published on 7th August 2018

Sufficient and sustained anti-thrombogenicity is essential for blood-contacting materials because blood coagulation and thrombosis caused by platelet adhesion and activation on material surfaces may lead to functional failure and even fatal outcomes. Covalently conjugating anti-thrombogenic moieties into a polymer, instead of surface modification or blending, can maintain the anti-thrombogenicity of the polymer at a high level over time. In this study, a series of randomly crosslinked, elastic, biodegradable polyurethanes (PU-DPA) were synthesized through a one-pot and one-step method from polycaprolactone (PCL) diol, hexamethylene diisocyanate (HDI) and an anti-thrombogenic drug, dipyridamole (DPA). The mechanical properties, hydrophilicity, in vitro degradation, and anti-thrombogenicity of the resultant PU-DPA polymers could be tuned by altering the incorporated DPA amount. The surface and bulk hydrophilicity of the polyurethanes decreased with increasing hydrophobic DPA amounts. All PU-DPA polymers exhibited strong mechanical properties and good elasticity. The degradation rates of PU-DPAs decreased with increasing DPA content in both PBS and lipase/PBS solutions. Covalently incorporating DPA into the polyurethane significantly reduced the platelet deposition compared to that of the polyurethane without DPA, and the polyurethane remained anti-thrombogenicity after degradation. The PU-DPA films also supported the growth of human umbilical vein endothelial cells. The attractive mechanical properties, blood compatibility, and cell compatibility of this anti-thrombogenic biodegradable polyurethane indicate that it has a great potential to be utilized for blood-contacting devices and cardiovascular tissue repair and regeneration.


Achieving anti-thrombogenicity of biomedical materials remains a challenge for blood-contacting devices and implants. Platelet adhesion and activation on material surfaces of blood-contacting medical devices and implants, such as stents, vascular grafts, artificial organ implants, and hemodialysis equipment, cause thrombotic and thromboembolic complications, which may need further intervention or induce implant failure and fatal outcomes.1,2 For example, large-diameter synthetic vascular grafts (e.g., EXCLUDER®, Endurant®, and AneuRx®) used for abdominal aortic aneurysm (AAA) treatment have limb thrombotic occlusion risks of up to 7.2%, which appears to be low, but they usually require further intervention to resolve ischemic symptoms.3–6 Applications of small-diameter synthetic vascular grafts (inner diameter < 6 mm) have been greatly limited because of their high rates of thrombotic occlusion, which may cause acute failure.7,8 Besides, the thrombotic and thromboembolic compilations during artificial heart implantation or cardiopulmonary bypass can sometimes be life-threatening.9,10 Therefore, it is essential to develop new blood-contacting materials with sufficient and sustained anti-thrombogenicity.

Biodegradable polyurethanes (PUs) have been extensively studied for blood-contacting devices, especially for vascular grafts and stent coatings.8,11–13 Surface modifications on biodegradable polyurethanes by anti-thrombogenic moieties (e.g., phosphorylcholine, sulfobetaine, anti-thrombogenic drugs, and heparin) have shown improved anti-thrombogenicity.11,14–16 However, such modifications may not have sustained anti-thrombogenicity due to the loss of surface-grafted moieties with degradation. These anti-thrombogenic molecules, such as an anti-thrombogenic drug, can also be blended with biodegradable PU to achieve anti-thrombogenic PUs. However, the anti-thrombogenicity of the drug eluting polyurethane scaffold can be gradually weakened and even lost due to the continuous and complete release of anti-thrombogenic drugs from the polyurethane scaffold with time via polymer degradation.8,17 To maintain the anti-thrombogenicity of the polymer, by covalently conjugating the anti-thrombogenic moieties into polyurethane, we can achieve an anti-thrombogenic polymer, and its thrombo-resistance can be sustained with degradation. Sulfobetaine was incorporated into the polyurethane backbone and showed markedly reduced thrombogenicity before and after degradation.17 In our previous study, dipyridamole (DPA), a clinically used anti-thrombogenic drug, was blended with a biodegradable polyurethane and then electrospun into fibrous scaffold, which has shown reduced thrombogenicity.18 Thus, it is interesting to investigate if covalently incorporating DPA into polyurethane can enhance the polymer anti-thrombogenicity without DPA release while providing long-term anti-thrombogenicity.

In this proof-of-concept study, biodegradable and elastic polyurethanes with DPA (PU-DPA) were synthesized through a one-pot and one-step method of combining polycaprolactone diol (PCL), hexamethylene diisocyanate (HDI) and DPA with four hydroxyl groups (Fig. 1A). The obtained polyurethanes exhibited a randomly crosslinked structure, and the DPA content could be tuned by varying the molar ratio of DPA to PCL. The chemical structure, mechanical properties and in vitro degradation profile of the synthesized PU-DPA polymers were characterized. The blood contact behavior of the PU-DPA films was assessed by hemolysis and platelet deposition using human whole blood. The in vitro cell compatibility of the polyurethane films was investigated by culturing human umbilical vein endothelial cells on PU-DPA films.

image file: c8tb01582a-f1.tif
Fig. 1 PU-DPA synthesis. (A) Synthesis scheme of PU-DPA from polycaprolactone (PCL) diol, hexamethylene diisocyanate (HDI) and an anti-thrombogenic drug, dipyridamole (DPA). (B) Digital images to show the color change with PU-DPA films after DPA incorporation.

Experimental methods


PCL (Mn = 2000, Sigma) was dried overnight to remove residual water in a vacuum oven at 60 °C prior to synthesis. HDI (Sigma) were dehydrated using vacuum distillation before use. Stannous octotate (Sn(Oct)2, Sigma) was dried using 4 Å molecular sieves. Dipyridamole (DPA, Sigma), dichloromethane (DCM, Sigma), 1,1,1,3,3,3-hexafluoroisopropanol (HFIP, Oakwood Product), ethanol (Decon labs, Inc.) hexamethyldisilazane (HMDS, Sigma), glutaraldehyde (Sigma), osmium tetroxide (Electron Microscopy Science), lipase from Thermomyces lanuginosus (≥100[thin space (1/6-em)]000 U g−1, Sigma) were used as received.

Synthesis of polyurethanes incorporating DPA

PU-DPAs were prepared using PCL, HDI and DPA through a one-step/one-pot synthesis method (Fig. 1A). Briefly, PCL was dissolved in DCM in a 20 ml scintillation vial, followed by addition of DPA with different ratios of DPA to PCL. After vigorous shaking to create a homogenous solution, HDI was added followed by Sn(Oct)2 catalyst (2 drops). The mixed solution was poured into a covered Teflon® dish and then placed in an oven at 70 °C for 45 minutes. The dish was then placed in a vacuum oven at 65 °C and allowed to dry overnight. The formed films were removed from the cup and immersed in 100% ethanol for 24 hours to remove unreacted monomers such as unreacted free DPA. The films were then placed in the vacuum oven at 65 °C and dried overnight. The molar ratios of hydroxyl groups in PCL to those in DPA were set as 100[thin space (1/6-em)]:[thin space (1/6-em)]0, 90[thin space (1/6-em)]:[thin space (1/6-em)]10, 80[thin space (1/6-em)]:[thin space (1/6-em)]20 and 70[thin space (1/6-em)]:[thin space (1/6-em)]30, which are referred to as PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0), PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10), PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30), respectively. The theoretical mass fractions of DPA in the PU-DPA polymers were 1.4 wt%, 3.1 wt% and 5.2 wt% for PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10), PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20), and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30), respectively.

Polyurethane characterization

Fourier transform infrared (FT-IR) spectra of the PU-DPA polymers were obtained using a Thermo Nicolet 6700 FT-IR spectrometer (Thermo Fisher Scientific, Waltham, MA). Water contact angle was measured by the sessile drop method using a KSV CAM 101 – Optical Contact Angle and Surface Tension Meter (KSV, Helsinki, Finland). Water absorption of PU-DPA polymers (n = 3) was measured by immersing a weighed film (W0) in a phosphate buffer solution (PBS, Sigma) at 37 °C for 48 hours and then by weighing the wet film (W1) after removing the surface water. The water absorption was calculated as (W1W0)/W0 × 100%. To measure the swelling ratio of PU-DPA films in HFIP, the films (n = 3) were weighed (Wd) and then immersed in HFIP at room temperature for 24 or 48 hours. At each predetermined time point, the films were weighed as Ws. The swelling ratio in HFIP was calculated as (WsWd)/Wd × 100%.

Mechanical tensile testing was carried on the polymer films by using 2 × 20 mm strips cut from the films. All testing (n = 6) was done using an MTS Insight mechanical tester (MTS System, Minneapolis, MN) fitted with a 500 N load cell (Model 56932701, MTS System, Minneapolis, MN). The testing was done at room temperature with a crosshead speed of 10 mm min−1. Cyclic stretching was conducted by stretching the PU-DPA films (2 × 20 mm, n = 3) to maximum strain of 30% and releasing back to the initial length for 10 cycles at a constant rate of 10 mm min−1.19

In vitro hydrolytic and enzymatic degradation

Polymer degradation was evaluated by weight loss of the films under hydrolytic and enzymatic conditions. Polymer films were cut to approximately 30 mg pieces and each piece (W0) was weighed precisely. For hydrolytic conditions, the weighed polymer films were immersed in 10 ml of phosphate buffer solution (PBS) at 37 °C. At the prescribed time points, samples were removed and rinsed thrice with deionized water and dried in vacuum oven at 65 °C for three days followed by weighing (W1). For enzymatic conditions, the weighed polymer films were immersed in 2 ml 100 U lipase/PBS solution in a 20 ml glass vial. The glass vials were maintained at 37 °C, and fresh lipase/PBS was replaced every two days. At each time point, the samples were removed, rinsed thrice with deionized water, and dried in a vacuum oven at 65 °C for three days followed by weighing (W1). The mass remaining was calculated by W1/W0 × 100%. Three samples were used for each polymer at each time point.

Hemolysis and platelet deposition of PU-DPA films

The hemolysis assay on PU-DPA materials was processed as previously described.20 Human blood was collected from healthy individuals into acid citrate dextrose (ACD) anticoagulant tubes using venipuncture and was handled following methods approved by the Institutional Review Board at the University of Texas at Arlington. Following this, the blood was diluted to 2% (v/v) using saline solution (0.9% NaCl (w/v)), and 200 μl of it was transferred on to polymeric disks having 6 mm diameter and incubated for 2 hours at 37 °C. The samples were then centrifuged at 1000g for 10 minutes, and the supernatant was collected to measure the absorbance at 545 nm on a UV/vis spectrophotometer (Infinite M200 plate reader, Tecan, Durham, NC). The blood diluted using deionized (DI) water and saline served as positive and negative controls, respectively. The percentage of hemolysis was calculated using the equation below:
image file: c8tb01582a-t1.tif

Anti-thrombogenicity of PU-DPA films was evaluated by platelet deposition. Briefly, the human whole blood was centrifuged at 120g for 12 minutes to obtain platelet-rich plasma (PRP), which was incubated with PU-DPA samples before and after 14 days of enzymatic degradation and glass (6 mm in diameter) for 1 hour at 37 °C. The samples were then rinsed 3 times with PBS, and the deposited platelets were then lysed with 1% Triton X-100 for 30 minutes at 37 °C. The amounts of lactate dehydrogenase (LDH) released by the deposited platelets were quantified using LDH assays following manufacturer's instructions (Clonetech, Mountain View, CA). In another parallel study, to observe the platelet deposition and aggregation, the PRP-treated samples were fixed in 2% glutaraldehyde solution overnight, stained with 1% osmium tetroxide for 2 hours, and dehydrated in a series of ethanol solutions with concentrations of 50%, 70%, 95% and 100% (v/v) for 10 minutes each step. The samples were further dried using a series of ethanol–HMDS solutions of ratios 2[thin space (1/6-em)]:[thin space (1/6-em)]1, 1[thin space (1/6-em)]:[thin space (1/6-em)]1 and 1[thin space (1/6-em)]:[thin space (1/6-em)]2 for 15 minutes at each step; then, they were sputter-coated with silver and imaged on a SEM (Hitachi S-3000N).

In vitro endothelial cell culture

Human umbilical vein endothelial cells (HUVECs, ATCC, Manassas, VA) were seeded at a density of 10[thin space (1/6-em)]000 cells per cm2 onto the surface of polymer disks (6 mm in diameter) and cultured for predetermined time points using Vasculife basal medium supplemented with growth factors (Vasculife VEGF lifefactors kit, Lifeline Cell Technologies, Frederick, MD). The cell culture medium was fully refreshed on alternate days. The tissue-cultured polystyrene (TCPS) served as a control. At each time point, the cellular viabilities (n = 4) were measured using MTS assays (Promega, WI) following manufacturer's instructions, and the cellular viability was expressed as a percentage relative to that of the control TCPS at day 1. The absorbance for the cells cultured on the control TCPS at day 1 was set as 100%. In addition, the morphology of HUVECs on the polymer film surface was also observed under SEM after fixation and dehydration, as described above.

Statistical analysis

Unless specified, all results were expressed as mean ± standard deviation. To determine the blood contact evaluation of PU-DPA films, we performed a statistical analysis using one-way ANOVA followed by post hoc Tukey testing. Two-way ANOVA was utilized followed by post hoc Sidak testing to observe the difference in cellular growth after 1 and 3 days of culture. All these data analyses were conducted in GraphPad Prism (GraphPad Software Inc., CA), and a statistically significant difference was considered when p < 0.05.

Results and discussion

Characterization of PU-DPA films

The synthesized polyurethane with DPA incorporation appeared yellow and transparent due to the intrinsic yellow color of DPA, whereas PU without DPA was colorless and transparent (Fig. 1B). The chemical structures of PU-DPA polymers were verified by FTIR spectra (Fig. 2). The characteristic peaks of polyurethanes were located at 3350 cm−1 (N–H stretching of urethane groups), 2930 cm−1 and 2870 cm−1 (symmetric and asymmetric C–H stretching), and 1730 cm−1 (C[double bond, length as m-dash]O stretching of urethane groups).21 The specific peak for DPA was located at 1530 cm−1 corresponding to the aromatic C–C symmetric stretching,22 and its intensity increased with increasing DPA amount in the polymer.
image file: c8tb01582a-f2.tif
Fig. 2 FT-IR spectra of PU-DPA polymers to verify their chemical structures.

The surface and bulk hydrophilicities of the PU-DPA films were characterized by water contact angle and water absorption, respectively (Table 1). The water contact angle increased with increasing DPA content. PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) without DPA exhibited the smallest contact angle at 89° ± 3°, whereas PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) with the highest DPA amount exhibited the largest contact angle at 99° ± 4° (p < 0.05). The addition of DPA into polyurethane also reduced polymer water absorption. PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) had the highest water absorption at 9 ± 1%, whereas PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) had the lowest water absorption at 3 ± 1% (p < 0.05). The increasing water contact angle and decreasing water absorption with increased DPA amount in the polyurethane indicated the increased surface and bulk hydrophobicities of the PU-DPA films, which were both due to the hydrophobic nature of DPA.23 The swelling ratio of PU-DPA films in HFIP (Fig. S1 in the ESI) decreased significantly with increased DPA amounts, showing the increased crosslinking degree of the PU-DPA polymer network, which is relevant for the polymer mechanical properties.

Table 1 Polymer film characterizationa
Samples Contact angle (deg) Water absorption (%) Tensile strength (MPa) Initial modulus (MPa) Breaking strain (%)
a a, b, c and d denote statistical differences between groups for each characterization.
PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) 89 ± 3a 9 ± 1a 15.3 ± 0.3a 3.1 ± 0.7a 1157 ± 109a
PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10) 95 ± 4a,b 7 ± 1b 15.0 ± 2.6a 3.6 ± 0.4a,b 1009 ± 64a
PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) 97 ± 5b 5 ± 1c 25.4 ± 1.2b 3.6 ± 0.1b 1107 ± 47a
PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) 99 ± 4b 3 ± 1d 20.3 ± 1.5c 4.4 ± 0.1c 829 ± 55b

Mechanical properties of PU-DPA films

The typical stress–strain curves of PU-DPA films are shown in Fig. 3A. In Table 1, it can be seen that PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) had the highest tensile strength (25.4 ± 1.2 MPa) compared to PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) (15.3 ± 0.3 MPa), PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10) (15.0 ± 2.6 MPa) and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) (20.3 ± 1.5 MPa). The initial moduli of the PU-DPA films increased with rising DPA contents, ranging from 3.1 ± 0.7 MPa [PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0)] to 4.4 ± 0.1 MPa [PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30)]. A higher DPA content may lead to higher crosslinking degree, which is associated with stronger and stiffer materials.24 The other reason is that the increasing DPA/PCL ratio can result in an increase of the hard segment in the crosslinked PU, which can also induce the initial modulus increase. No significant difference in the breaking strain was observed among PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) (1157 ± 109%), PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10) (1009 ± 64%), and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) (1107 ± 47%) (p > 0.05). However, the PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) film had the lowest strain at 829 ± 55% (p < 0.05) because of its highest crosslinking degree. It is notable that the tensile strength of PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) (20.3 ± 1.5 MPa) was lower than that of PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) (25.4 ± 1.2 MPa). The polymer mechanical strength is relevant to various factors such as microphase separation and crystallization, which the DPA content may interfere with. On the other hand, the shorter reaction time (45 min) and lower reaction temperature (70 °C) for PU-DPA synthesis may result in a relatively lower degree of crosslinking, which can lead to a relatively lower initial polymer modulus (<5 MPa) compared to those of previously reported crosslinked PCL-based polyurethanes [(7.2–33.8 MPa, reaction taken place at 90 °C for 6 h)25 and (40.92 MPa, reaction taken place at 85 °C for 21 h)].26
image file: c8tb01582a-f3.tif
Fig. 3 Mechanical properties of PU-DPA. (A) Typical stress–strain curves of PU-DPA films. (B) Images to show that the PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) film is highly stretchable. (C) Cyclic stretching curves of PU-DPA polymers (n = 3) at 30% strain to exhibit their high elasticity.

The high elasticity of the PU-DPA polymers was verified by cyclic stretching of PU-DPA films at a maximum strain of 30% (Fig. 3B). All groups had a large hysteresis loop in the first cycle, followed by much smaller hysteresis loops in the next nine cycles. All groups showed small irreversible deformations (<5%) at 30% strain, indicating the good elasticity of the PU-DPA films. This may be due to the nature of the microphase separation of polyurethane and the crosslinked network structure.

In vitro degradation of PU-DPA films

The PU-DPA hydrolytic and enzymatic degradation kinetics in PBS and lipase/PBS solutions are shown in Fig. 4A and B, respectively. For hydrolytic degradation (Fig. 4A), all the polymers demonstrated slow degradation rates within 8 weeks. PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) had a significantly higher mass loss (2.8 ± 1.8%) compared to PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10) (1.5 ± 0.5%), PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) (1.4 ± 1.0%), and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) (1.4 ± 0.3%) (p < 0.05) after 8 weeks of hydrolysis. The hydrolytic degradation of all PU-DPA polymers was slow because PCL is a slow degrading polymer,27 and our results were consistent with those of previous studies.28–31 The addition of DPA into polyurethane further reduced its hydrolysis rate because the hydrophobic characteristics of DPA and the formed crosslinked networks could enhance the stability of the polymer structure and make it resistant to both hydrolytic and enzymatic degradation.32 The enzymes existing in the body and the host response (e.g., macrophage accumulation) could accelerate polymer degradation.33 The enzymatic polymer degradation in lipase/PBS solution presented similar trends as that in PBS; however, these enzymatic degradation rates were correspondingly higher than those in PBS (Fig. 4B). In the lipase/PBS solution, PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) had the highest degradation rate (21.8 ± 10.4% mass remaining at day 14), whereas PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) had the lowest degradation rate (98.6 ± 0.7% mass remaining at day 14). SEM images further confirmed that the PU-DPA films were enzymatically degraded by observing changes in their surface morphologies (Fig. 4C). The surfaces of all solid polymer films were smooth and homogeneous prior to degradation. After 14 days of enzymatic degradation, the film surfaces showed roughness and even cracks. Hydrolase lipase is primarily used to catalyze ester bond break-down with hydrolysis.34 The higher DPA content was associated with lower PCL content, indicating that less ester bonds existed. In addition, the crosslinking network could resist water penetration into the polymer, which could further inversely affect its enzymatic degradation.32
image file: c8tb01582a-f4.tif
Fig. 4 Mass remaining for PU-DPA films (n = 3) in (A) PBS up to 8 weeks and (B) 100 U ml−1 lipase/PBS up to 14 days at 37 °C. * Represents significantly different groups (p < 0.05). (C) Scanning electron micrographs of the surface morphologies of PU-DPA films before and after enzymatic degradation for 14 days.

Hemolysis and platelet deposition of PU-DPA films

The blood contact behaviors of the PU-DPA materials were assessed through hemolysis, platelet adhesion, and platelet aggregation. A hemolysis assay is one of the indispensable initial tests to determine the adverse effects of polymeric biomaterials on red blood cells (RBCs) in vivo. If the material is hemo-incompatible, it disrupts RBCs, leading to the release of its intercellular contents including hemoglobin into the plasma, which can eventually result in critical consequences such as anemia, jaundice, acute renal failure, and death.35,36 Based on the ISO standard practice for assessment of hemolytic properties of materials,37 a biomaterial is classified as nonhemolytic, slightly hemolytic, or hemolytic when the percentage of hemoglobin released after incubating the whole blood with biomaterial is 0–2%, 2–5, or >5%, respectively. The results obtained for hemolysis of ACD blood with PU-DPA materials are shown in Fig. 5A. All of the polymers were observed to cause less than 2% of hemolysis with statistical differences (p > 0.05) although there existed a trend of reduced hemolysis with the increase in DPA amount in the polyurethane. This indicated that PU-DPA materials are non-hemolytic and may be suitable for the desired vascular applications.
image file: c8tb01582a-f5.tif
Fig. 5 Hemocompatibility assessment of PU-DPA films. (A) Percentage of red blood cell lysis after 2 hour exposure of whole blood with PU-DPA films (n = 4). DI water- and saline-diluted blood served as positive and negative controls, respectively. Glass is used for a comparison for blood testing studies (* denotes p < 0.05 for significant difference with respect to glass, PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0), PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10), PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30)), (B) amount of LDH released by triton-treated platelets after adhering on to PU-DPA films in 1 hour incubation at 37 °C (n = 5, $ and # denote p < 0.05 for significant difference with respect to PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) and glass, respectively); (C) SEM images to show platelet deposition and morphology on PU-DPA films.

Platelet adhesion and aggregation onto a biomaterial implant determine its anti-thrombogenic potential. In response to the proteins adsorbed onto the foreign surface, platelets interact with them via their integrin receptors and become activated. These activated platelets can further activate many types of coagulation factors, promoting thrombosis and aggregation. Therefore, the prevention of platelet adhesion to the biomaterial surface is crucial to improve its blood compatibility. As shown in Fig. 5B, increasing amounts of DPA incorporated into the polyurethanes resulted in lower amounts of platelet deposition. However, substantially reduced amount of platelet deposition on the material surface was observed only for PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) compared to that for PU films without DPA. Furthermore, in agreement with LDH assays that quantified the platelet adhesion, SEM images showed similar trends in platelet deposition on the PU films. These images in Fig. 5C revealed that the platelets deposited onto PU films with DPA have a round morphology, which indicated inactivated platelets, whereas those on PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) mostly have an irregular morphology with fully spread pseudopodia, and this represented the activated state of platelets.38 To investigate the retention of anti-thrombogenicity for the enzyme-degraded PU-DPA films, human platelet-rich plasma was contacted with PU-DPA films after 14 days of enzymatic degradation (Fig. 6). Significantly reduced platelet deposition was found on the enzyme-degraded PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) film compared with that on degraded PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) and glass (Fig. 6A). SEM images quantitatively confirmed the LDH results (Fig. 6B), in which less platelet deposition was found on the enzyme-degraded PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) films than those on degraded PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) and glass. These results showed that covalent incorporation of DPA into the polyurethane can reduce platelet deposition, decrease the propensity of the polymeric surface to activate adherent platelets and maintain anti-thrombogenicity over a long term with degradation, thereby improving their anti-thrombogenicity with endurance.

image file: c8tb01582a-f6.tif
Fig. 6 Blood platelet deposition on enzymatically degraded PU-DPA films. (A) Amount of LDH released by triton-treated platelets deposited on PU-DPA films after 1 hour incubation at 37 °C (n = 5, $ and # denote p < 0.05 for significant difference with respect to PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) and glass, respectively); (B) SEM images to show platelet deposition and morphology on enzymatically degraded PU-DPA films (14 days).

The mechanism of action of conjugated DPA on the anti-platelet activity is still not clear. We envisage that DPA molecules incorporated into polymeric films might be recognized by the anchoring molecules on the platelet's surface, leading to an increase in intracellular levels of cyclic adenosine monophosphate (cAMP) and cyclic guanine monophosphate (cGMP) within platelets.39,40 It is also plausible that DPA grafting changes the physical properties of the polymeric surface, which in turn affects the interaction of plasma proteins with the material and subsequent platelet responses. For instance, Wu et al.41 analyzed the effects of surface properties of various types of PUs and observed lower platelet adhesion on hydrophilic PU than that on its hydrophobic counterparts due to ultralow adsorption of fibrinogen from plasma to the material surface. Interestingly, the hydrophobic PUs with relatively high fibrinogen adsorption also exhibited lower platelet adhesion. The studies concluded that surface chemistry dictates the fibrinogen conformation adsorbed onto material, thereby exposing the platelet binding domains affecting their adhesion and activation.41,42 As previously noted, PU-DPA exhibited increasing hydrophobicity with increasing contents of DPA, with PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) having the highest contact angle of 99° ± 4° compared to PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) with 89° ± 3°. Therefore, the effect of the surface characteristics of PU-DPA samples on the observed platelet response cannot be eliminated. Hence, further studies will be required to investigate the mechanism of how PU-DPA materials enhance its anti-thrombogenic characteristics.

In vitro HUVEC viability on PU-DPA films

Endothelial cell layer formation on materials is a gold standard to offer long-term anti-thrombosis for implanted blood-contacting medical devices. The ability of PU-DPA films to support endothelial cell growth was exhibited in Fig. 7. Cellular viability was determined using MTS assays after 1 and 3 days of culture on PU-DPA films (Fig. 7A). The cells grown on TCPS served as the control. Significantly lower percentage of endothelial cells adhered and proliferated on the polymeric films compared to that for TCPS for both the time points (p < 0.0001). This reduction in the cell performance on the PU-DPA samples was due to the hydrophobic nature, which can be improved by numerous surface modification strategies including physical modification by plasma treatment, etching, gamma irradiation, and X-ray treatment,43–45 chemical modification by grafting hydrophilic materials,46–49 and bioactive modification by immobilization of endothelial cell-adhesive peptides on to the polymeric surface.50–52 However, cellular viabilities on PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) were significantly higher than those for PU-DPA (90[thin space (1/6-em)]:[thin space (1/6-em)]10) and PU-DPA (100[thin space (1/6-em)]:[thin space (1/6-em)]0) at day 3 (p < 0.05), which was also qualitatively verified by the electron micrographs of spread HUVECs seeded on PU-DPA films (Fig. 7B). In addition, HUVEC proliferation was observed for both PU-DPA (80[thin space (1/6-em)]:[thin space (1/6-em)]20) and PU-DPA (70[thin space (1/6-em)]:[thin space (1/6-em)]30) as their cell viabilities at day 3 were significantly higher than those at day 1 (p < 0.05). In our previous study, it was observed that DPA releasing biodegradable elastic PU (BPU) scaffolds improved human aortic endothelial cell growth over 7 days of culture compared to that for BPU scaffolds alone.20 Aldenhoff et al.53 observed that a DPA-immobilized PU-based vascular graft supported the formation of endothelial-like cell lining in a sheep carotid artery model. These reports and our results imply that DPA involvement may support endothelial cell growth. It would make the PU-DPA polymer attractive as a good blood-contacting material candidate because it has excellent anti-thrombogenicity without adverse effects on endothelialization.
image file: c8tb01582a-f7.tif
Fig. 7 In vitro HUVEC growth on PU-DPA films. (A) Cell viability measured by MTS assay to show HUVEC growth on the surfaces of PU-DPA films up to 3 days (n = 4). TCPS surface served as the control. (B) SEM images show HUVEC morphology on surfaces of PU-DPA films at day 3 (* denotes p < 0.05 for significant difference within the groups after 1 and 3 days of culture).

Some limitations exist in this study. First, the polymer mechanical properties cannot only be tuned by altering the feeding ratio of DPA to PCL, but also by altering the reaction time and temperature during the PU-DPA synthesis. In this study, 45 min of reaction time at 70 °C might be short, as discussed above. However, low-initial modulus polymers are also important for biomedical engineering applications. They can be utilized as porous scaffolds to repair soft and elastic human tissues with low initial moduli.54,55 Second, the introduction of DPA significantly reduces the biodegradation rate of polyurethane due to the increased hydrophobicity and network structure. To accelerate the biodegradability of PU-DPA, polyester diols with higher molecular weights containing more lipase-susceptible ester bonds can be used as the soft segment in the polyurethane structure54 instead of the PCL (Mn = 2000) diol. Additionally, after processing the PU-DPA polymer into 3D porous scaffold, its biodegradability may be enhanced because of the porous structure and greatly increased surface area, which can allow higher water penetration into the polymer network and more interaction between water and polymer chains.56,57 Third, material characterizations are focused on 2D films, which may provide strong evidence for anti-thrombogenic coating applications; however, this is not sufficient for 3D porous scaffold use. In the future, we will further evaluate the mechanical properties and biological functions of PU-DPA based 3D scaffolds in vitro and in vivo.


A family of biodegradable, elastic and crosslinked polyurethanes with an anti-thrombogenic drug was synthesized using a simple one step/one pot method. The crosslinked polyurethanes had good elasticity, and their mechanical and degradation properties could be tuned by altering the feeding ratio of DPA to PCL segments. The resultant anti-thrombogenic polyurethanes significantly reduced platelet deposition on polymer surfaces before and after enzymatic degradation and also supported the growth of HUVECs. These promising characteristics of the anti-thrombogenic and biodegradable elastomer show its great potential to be used as a blood-contacting biomaterial.

Conflicts of interest

There are no conflicts to declare.


We greatly appreciate the partial financial support from CAREER award 1554835 (YH) from the National Science Foundation, and R21HD090680 (YH), R01HL118498 (KTN), and T32 HL134613 (KTN, and AEK is a fellow of NIH T32) from the National Institutes of Health in the United States.


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Electronic supplementary information (ESI) available: Swelling ratio of PU-DPA films in HFIP. See DOI: 10.1039/c8tb01582a

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