Strategies for Improving Endothelial Cell Adhesion to Blood-Contacting Medical Devices

Abstract

The endothelium is a critical mediator of homeostasis on blood-contacting surfaces in the body, serving as a selective barrier to regulate processes such as clotting, immune cell adhesion, and cellular response to fluid shear stress. Implantable cardiovascular devices, including stents, vascular grafts, heart valves, and left ventricular assist devices, are in direct contact with circulating blood and carry a high risk for platelet activation and thrombosis without a stable endothelial cell (EC) monolayer. Development of a healthy endothelium on the blood-contacting surface of these devices would help ameliorate risks associated with thrombus formation and eliminate the need for long-term antiplatelet or anticoagulation therapy. Although ECs have been seeded onto or recruited to these blood-contacting surfaces, most ECs are lost upon exposure to shear stress due to circulating blood. Many investigators have attempted to generate a stable EC monolayer by improving EC adhesion using surface modifications, material coatings, nanofiber topology, and modifications to the cells. Despite some success with enhanced EC retention in vitro and in animal models, no studies to date have proven efficacious for routinely creating a stable endothelium in the clinical setting. This review summarizes past and present techniques directed at improving the adhesion of ECs to blood-contacting devices.

Impact statement

Clinical success of blood-contacting devices such as vascular grafts, stents, and heart valves has remained limited by postimplantation problems, including thrombosis and loss of patency. Without a stable endothelial cell (EC) monolayer, blood-contacting devices are at risk for platelet activation and thrombosis. Methods to improve EC adhesion on these devices have not translated to long-term in vivo success, as many ECs are lost after exposure to circulating blood. In this study, we summarize methods to improve EC adhesion and retention. Successful endothelialization of blood-contacting devices may improve patient outcomes after device implantation and limit the need for long-term antiplatelet or anticoagulation therapy.

Introduction: The Need for Endothelialization

A critical problem that persists among implantable medical devices for cardiovascular applications is the risk of platelet activation and thrombus formation. This is due to the nonspecific adsorption of serum proteins (e.g., fibrinogen, vitronectin) to the surface of implanted biomaterials. Once adsorbed, these proteins bind to circulating platelets, resulting in platelet adhesion, activation, and aggregation, which in turn lead to thrombosis.¹

As a consequence of the nonspecific nature of serum protein adsorption, all biomaterials exhibit at least some degree of thrombogenicity, which requires careful management in the clinical setting. For example, pharmacological agents (e.g., anticoagulation therapy and antiplatelet therapy) are routinely administered to inhibit thrombosis in patients implanted with devices, including coronary stents, flow diverters, mechanical heart valves, and left ventricular assist devices. These agents have the serious side effect of bleeding, with severe bleeding occurring in ∼5% of patients.²

Consequently, bleeding risk must be carefully weighed against thrombosis risk, which persists at a rate of ∼2% of patients.³ Worse yet, small-diameter synthetic grafts, including coronary artery bypass grafts, have very limited clinical utility due to 1-year patency rates of only 60%.^4,5 Loss of graft patency is typically a result of thrombosis and neointimal hyperplasia, both of which are inhibited by endothelial cells (ECs).

All blood-contacting surfaces in the body are lined with an endothelium. The endothelium provides a biologically active antithrombogenic surface by virtue of nitric oxide, prostacyclin, thrombomodulin, and heparan sulfate expression, all of which inhibit thrombosis.⁶ Therefore endothelialization of blood-contacting devices provides a solution for mimicking the natural state in the body and may offer positive long-term outcomes in patients, without the need for anticoagulation or antiplatelet therapy.⁷ Indeed, early clinical studies demonstrated improved patency rates of synthetic vascular grafts seeded with ECs.^8–10

The literature describes countless efforts to endothelialize biomaterials for implantation, yet few have demonstrated long-term in vivo success due to high rates of cell loss upon exposure to physiological blood flow. In a seminal study, Rosenman et al. found that 40% of ECs seeded on a vascular graft were lost within 1 h and more than 80% of cells were lost within 24 h of in vivo implantation.¹¹ The exact mechanism by which cells detach remains unclear. It is speculated that a lack of mature focal adhesion sites, which anchor the cell to the subendothelial extracellular matrix (ECM), can result in EC detachment upon exposure to blood flow.¹²

This review discusses the most common methods tested to improve EC adhesion and retention. Different approaches to augment EC adhesion have been applied to biomaterials and ECs to try and recapitulate the native EC-ECM interactions (Fig. 1). Despite some success in vitro and in vivo, there are limited clinical studies investigating endothelialization of synthetic vascular grafts in humans.¹³

FIG. 1.

Schematic of native and engineered adhesion for improving EC retention on blood-contacting devices. Left: attachment of ECs onto ECM shown in a native biological milieu in the presence of fluid flow and shear stress. Right: attachment of ECs onto engineered substrate or biomaterial. Adhesion-promoting strategies are represented, including genetically engineered cells, ECM coating, and surface modifications. Cell adhesion is achieved by CAMs, which are located on the surface of ECs and mediate both cell-ECM and cell-cell adhesion. Integrins are one of the superfamilies of CAMs and are especially important for cell adhesion to a substrate. CAMs, cell adhesion molecules; EC, endothelial cell; ECM, extracellular matrix. Color images are available online.

EC Function and Adhesion to the ECM

The endothelium is a continuous monolayer of ECs that lines blood-contacting surfaces in the body. Individual ECs form tight cell-cell junctions to create a semipermeable membrane at the interface between circulating blood and underlying tissue.¹⁴ Under normal conditions, ECs promote a nonthrombogenic blood-tissue interface by release of chemical substances, including nitric oxide and prostacyclin. These substances inhibit aggregation of platelets and cause vasodilation.⁶ Upon exposure to stimuli such as histamine or thrombin, ECs switch to an activated state. Activated ECs generate a prothrombotic, proliferative, and vasoconstrictive environment, which is a compensatory mechanism to minimize hemorrhage.⁶

In addition to cell-cell junctions, ECs also adhere to the underlying ECM basement membrane, which comprises laminin, collagen, fibronectin, nidogen, glycosaminoglycans, and proteoglycans.¹⁵ The ECM is an important regulator of EC functions, including adhesion and migration. ECs adhere to the ECM through basal surface integrin molecules, which are αβ heterodimeric transmembrane receptors that interact with specific ligands in the ECM. For example, the α₅β₁ integrin is specific for fibronectin and α_vβ₃ is specific for both vitronectin and fibronectin.¹⁶

Integrins cluster together to form focal contacts. These structures contain linker proteins, including α-actinin, talin, and vinculin, which bind integrins with the intracellular actin filament cytoskeleton.¹² This, in turn, leads to the formation of larger, elongated structures known as focal adhesions.¹⁷ Focal adhesions are critical sites for modulation of EC adhesion, signaling, and actin cytoskeleton structure.¹⁸

In vitro EC adhesion begins with initial attachment due to nonspecific electrostatic interactions with the substrate. This is followed by cell spreading and flattening as the cell adheres to the substrate through specific integrin binding and reorganizes its cytoskeletal structure. Cellular adhesion strength increases as focal adhesions form and mature, leading to stable connections with the underlying substrate.¹⁹

Changes in the molecular environment or applied forces, including shear stress, can cause cells to modify focal adhesions.¹⁸ Many strategies for improving EC adhesion strength to biomaterial substrates are based on exploiting integrin-ECM interaction and the mechanisms that lead to the formation of mature focal adhesions.

Biomaterial Coatings

Some of the earliest attempts at improving EC adhesion involved adding a coating onto biomaterials. The goal is to provide the cues and structures involved in native cell adhesion to encourage EC adhesion to a biomaterial, such as coating with ECM components, including fibronectin or laminin, to provide the ligands for EC-integrin interactions.¹⁵ Numerous types of coatings and biomaterials have been explored and we have summarized the in vitro studies reporting improved EC retention (Table 1). These studies and others described in this review depict how factors such as biomaterial, modification method, EC type, and flow conditions can dramatically affect the EC retention reported. Therefore, it is difficult to generalize the effectiveness of each biomaterial coating beyond the specific conditions tested in a particular study.

Table 1.

Summary of Biomaterial Coating In Vitro Studies to Improve Endothelial Cell Adhesion and Retention

Biomaterial	Coating	EC source and seeding method	Perfusion method	EC adhesion or retention	References
PTFE vascular graft	Fn	CJVECs seeded onto grafts for 1 h	Ex vivo canine carotid, external jugular arterial venous shunts for 2 h	EC retention (%; mean ± SD) Uncoated 7.60 ± 4.83 (primary culture ECs) 6.83 ± 14.03 (secondary culture ECs) Fn-coated 43.9 ± 21.04 (primary culture ECs) 30.3 ± 14.03 (secondary culture ECs)	Seeger and Klingman²²
PTFE vascular graft	Fn or Co	BAECs radiolabeled with indium 111 seeded on grafts and grown to confluence	Perfusion system with low (25 mL/min) or high (200 mL/min) flow for 1 h	Average EC retention (%; mean) as measured by remaining adherent radioactivity Low flow: 93 High flow: 89	Sentissi et al.²³
ePTFE vascular graft	Pre-clot (whole blood) or Fn	AHSVECs seeded for 90 min	Perfusion system with peristaltic pump and flow rates of 25, 100, 200, and 300 mL/min for 2 h	EC retention (%; mean ± SD) Preclot 25 mL/min: 82.4 ± 6.8 100 mL/min:79.9 ± 8.2 200 mL/min:75.4 ± 9.5 300 mL/min: 58.3 ± 15.5 Fn 25 mL/min:57.8 ± 9.9 100 mL/min:55.2 ± 13.3 200 mL/min: 55.4 ± 12.9 300 mL/min: 56.5 ± 15.2	Vohra et al.²⁴
ePTFE vascular graft	Fn	Autologous ovine carotid artery ECs labeled with 35S-methione cultured for 48 h before flow	Ex vivo perfusion in an externalized carotid-jugular shunt with blood flows of 150 mL/min and mean wall shear stress of 25 dyn/cm² for 3–24 h	EC retention (%; mean) 3 h: 81 24 h: 97	James et al.²⁵
ePTFE vascular graft	FG containing FGF-1 and heparin, Fn	CJVECs radiolabeled with indium-111 seeded overnight	Perfusion system with peristaltic pump, 1 h of 120/80 mmHg pressure and 60 pulsations/min	EC retention (%; mean ± SD) as measured by remaining adherent radioactivity FG grafts: 96 ± 5 Fn grafts: 84 ± 3	Gosselin et al.²⁶
ePTFE vascular graft	RGD peptide or Fn	AHSVECs seeded for 45 min	Perfusion system with peristaltic pump using a shear stress of 2.9 dyn/cm² for 1 h	EC retention (%; mean ± SD) Uncoated: 13.9 ± 7.9 Fn coated: 45.5 ± 2.1 RGD coated: 62.9 ± 7.5 Polylysine/glutaradialdehyde coated: 6.6 ± 1.5	Walluscheck et al.²⁷
PTFE vascular graft	Fn or naturally produced ECM	BAECs seeded for 16 h	Perfusion system using shear stress of 15 dyn/cm² for 1 h	Average EC retention (%; mean) PTFE: 20 Fn-PTFE: 54 Fn+ECM-PTFE: 85	Schneider et al.²⁸
ePTFE vascular graft	Fibroblastic matrix coated	CJVECs seeded on graft in monolayer for 24 h before ex vivo flow	Ex vivo femoro-atrial circuit in each dog; flow for 1 h	Percent of surface covered by ECs (mean ± SD) Before Uncoated: 76.90 ± 8.34 Coated: 71.65 ± 6.23 After Uncoated: 28.19 ± 9.36 Coated: 34.96 ± 10.10 EC retention (%) Coated: 48.8	Buján et al.³⁶
ePTFE vascular graft	RGD-cross-linked fibrin gel	AHSVECs seeded and grown for 9 days	Perfusion with peristaltic pump using average shear stress of 12 dyn/cm² for 24 h	EC retention (%) Non-RGD-enriched fibrin gel: 72.8 4 mg RGD-enriched fibrin gel: 86.7 8 mg RGD-enriched fibrin gel: 58.6 16 mg RGD-enriched fibrin gel: 57.0	Meinhart et al.⁴³
ePTFE vascular graft	RGD FSP or Fn	HPAECs seeded overnight	Perfusion with peristaltic pump with shear stress of 8 dyn/cm² for 24 h	EC retention (%; mean ± SD) Fn: 36 ± 11 RGD FSP: 59 ± 14	Dudash et al.²⁹
ePTFE vascular graft	ECM or CD34 antibody coating	HUVECs	Perfusion with peristaltic pump and shear stress of 8 dyn/cm² for 24 h	EC adhesion (cells/cm²) Before shear stress ePTFE: 4.1 × 10³ ± 0.5 × 10³ ECM/CD34-ePTFE: 2.0 × 10⁴ ± 5.0 × 10³ Reported the number of HUVECs on ECM/CD34-ePTFE grafts was significantly higher than ePTFE grafts, but did not provide values	Chen et al.³⁷
ePTFE or polyester elastomer vascular grafts	Fn	HUVECs labeled with indium 111 were seeded an average 18 h on the grafts before perfusion	Perfusion with peristaltic pump with shear stress of 15 dyn/cm² for 1 h	EC retention (%; mean ± SD) Fn-coated Hytrel: 92.1 ± 10.2 Fn-coated ePTFE: 61.6 ± 11.4 Hytrel: 39.7 ± 24.1 ePTFE: 25.8 ± 2.3	Kesler et al.³⁰
ePTFE or PET	Fn-PTFE Ge-PET	AHSVECs seeded for 90 min	Perfusion system with peristaltic pump using flow rates of 200 or 300 mL/min for 2 h	Average EC retention (%; mean) Fn-PTFE: 200 mL/min: 55.4 ± 12.9 300 mL/min: 56.5 ± 15.2 Gelatin-PET: 200 mL/min: 69.0 ± 6.0 300 mL/min: 66.5 ± 5.5	Vohra et al.³¹
ePTFE, PET (Ge-PET), Co-PET	Fn	HAECs harvested from saphenous veins seeded for 1 h	Perfusion system with peristaltic pump using flow rate of 200 mL/min for 24 h	EC retention (%; mean ± SD) 1 h: Fn-ePTFE: 40 ± 22 Fn-Ge-PET: 75 ± 14 Fn-Co-PET: 90 ± 3 24 h: Fn-ePTFE: 16 ± 10 Fn-Ge-PET: 25 ± 5 Fn-Co-PET: 65 ± 4	Carr et al.³²
Polyester elastomer vascular graft	Fn	CJVECs radiolabeled with indium-111 and seeded/cultured until confluence was achieved on lumen	Perfusion system with peristaltic pump; low flow condition used 120/80 mmHg and mean flow of 90 mL/min with shear rate of 240 s⁻¹; high flow condition used 200/100 mmHg with mean flow of 195 mL/min and shear rate of 520 s⁻¹ for 120 min	EC retention (%; mean) as measured by remaining adherent radioactivity Low flow: 89.7 High flow: 91.0	Greisler et al.³³
PET impregnated with collagen or polyester elastomer vascular graft	Fn, Co, or heparin	HUVECs cultured 18 h after seeding	Perfusion with peristaltic pump using 15 dyn/cm² shear stress for 1 h	EC retention (%; mean ± SD) PET with Co grafts Co: 48.8 ± 10.5 Co+heparin: 31.3 ± 11.4 Co+Fn: 49.7 ± 7.5 Co+Fn+heparin: 55.2 ± 11.1 Polyester elastomer grafts Fn+heparin: 76.3 ± 6.8 Fn: 85.2 ± 9.4	Dalsing et al.³⁴
Self-expanding nitinol endovascular stent	PLL or Fn	HUVECs	Parallel plate flow chamber using shear stress of 24 dyn/cm² for 24 h	Average EC retention (%; approximated from graph) Control uncoated: 20 PLL coated: 90 Fn coated: 90	Wang et al.³⁵
PCL-coated 316L stainless steel	Anti-CD34 antibody and mussel adhesive polypeptide	HUVECs seeded for 48 h with an enzymatic treatment incubation	No perfusion. Used wash after enzymatic treatment to determine retention	EC retention rate (%; mean approximated from graph) Uncoated: 65 Polypeptide coated: 85 anti-CD34 coated: 75 Polypeptide+anti-CD34 coated: 90	Yin et al.³⁸
Decellularized porcine carotid arteries	Basic fibroblast growth factor (bFGF)	HMECs seeded on heparin-coated grafts for 3 h	Perfusion with peristaltic pump for 6 h with 100 mL/min flow and 15 mmHg pressure or static control	Perfusion graft: 141 ± 10 cells/mm² Static graft: 234 ± 18 cells/mm² Average EC retention (%; mean) Perfusion: 60	Conklin et al.³⁹
Porcine aorta dECM	Gelatin as a carrier to deliver ECs	RAECs in gelatin perfused over dECM	Cells were delivered and perfused into the coronary arteries of decellularized adult rat whole hearts using 1 mL/min flow rate	EC retention (% of cells perfused) No gelatin: 58 ± 0 1% gelatin: 86 ± 2	Tang-Quan et al.⁴⁰
Decellularized porcine heart valve	Fucoidan and VEGF PEM film	HUVECs cultured overnight	Radial flow chamber system with peristaltic pump to generate 0.1–20 Pa (1–200 dyn/cm²) shear stress	Detachment efficiency (%; estimated from graph) 5 Pa Uncoated: 98 Fucoidan/VEGF PEM coated: 70 10 Pa Uncoated: 100 Fucoidan/VEGF PEM coated: 100	Marinval et al.⁴¹

AHSVECs, adult human saphenous vein endothelial cells; BAEC, bovine aortic endothelial cell; bFGF, basic fibroblast growth factor; CJVEC, canine jugular vein endothelial cell; Co, collagen; Co-PET, collagen-impregnated polyethylene terephthalate; dECM, decellularized extracellular matrix; EC, endothelial cell; ECM, extracellular matrix; ePTFE, expanded polytetrafluoroethylene; FG, fibrin glue; FGF-1, fibroblast growth factor 1; Fn, fibronectin; FSP, fluorosurfactant; Ge-PET, gelatin-impregnated polyethylene terephthalate; HAEC, human adult endothelial cell; HMEC, human microvascular endothelial cell; HPAEC, human pulmonary artery endothelial cell; HUVEC, human umbilical vein endothelial cell; PCL, polycaprolactone; PEM, polyelectrolyte multilayer; PET, polyethylene terephthalate; PLL, poly-l-lysine; PTFE, polytetrafluoroethylene; RAEC, rat aortic endothelial cell; RGD, arginine-glycine-aspartic acid (Arg-Gly-Asp); SD, standard deviation; VEGF, vascular endothelial growth factor.

Characterization of fibronectins by Yamada and Olden in 1978 prompted investigators to study fibronectin coatings on synthetic graft materials.²⁰ Fibronectin is a glycoprotein in the ECM shown to promote cell attachment and spreading on biomaterials through integrin α₅β₁ interactions.²¹ Coating expanded polytetrafluoroethylene (ePTFE) grafts with fibronectin has led to enhanced EC retention under numerous perfusion conditions.^22–32 The greatest improvement in EC retention was reported by Sentissi et al. with an average EC retention of ∼93% after 25 mL/min of flow for 1 h.²³ Although early retention may have been favorable in this study, Carr et al. reported a mean EC retention of only 16% ± 10% after 24 h of perfusion.³²

Other vascular graft materials such as polyester elastomer^30,33 and polyethylene terephthalate (PET)^31,32,34 coated with fibronectin have also exhibited greater EC retention under varying flow conditions compared to uncoated surfaces.^30–34 For example, fibronectin-coated collagen-impregnated PET had an EC retention of 90% ± 3% and 65% ± 4% after 1 and 24 h of flow, respectively. This retention was even greater than the value reported on fibronectin-coated ePTFE. Nitinol endovascular stents coated with fibronectin also showed enhanced EC retention compared to uncoated controls with an average 90% retention after 24 h of perfusion with 24 dyn/cm².³⁵ Other types of coatings on biomaterials include whole blood pre-clot,²⁴ fibrin glue,²⁶ biological matrices,^28,36 antibodies,³⁷ heparin,³⁴ adhesive peptides,³⁸ and growth factors.^39–41

A seminal article published by Ruoslahti and Pierschbacher sparked new perspectives on EC adhesion based on the favorable interaction between an EC integrin and its ligand in fibronectin.⁴² The arginine-glycine-aspartic acid (Arg-Gly-Asp; RGD) peptide sequence is a crucial structure in fibronectin recognized by ECs. Peptides that contain this RGD sequence promote the adhesion of cells without the need for the entire fibronectin glycoprotein, making it a more efficient coating. ePTFE grafts coated with RGD peptide show superior EC retention under flow conditions compared to fibronectin-coated or uncoated grafts.^27,29,43

Following successful outcomes in vitro, animal models have been used to examine EC retention on coated biomaterials in vivo. Peptide and protein coatings have proven effective at enhancing EC coverage on vascular grafts in dog,^28,44,45 sheep,^46–48 rat,^49–51 and pig⁵² models. ePTFE grafts coated with heptamaltose (M7) peptide to block nonspecific protein adsorption, RGD, or EC-selective CRRETAWAC (cRRE) peptide exhibited greater EC retention and enhanced patency in a porcine carotid artery model after 1 month compared to uncoated grafts.⁵² The RGD-coated grafts had the greatest patency rate of 67% and all grafts that remained patent showed luminal EC coverage.

However, some grafts still showed thrombosis early after implantation and the mechanism remained unclear. Bastijanic et al. speculated it may have been due to disruption of the confluent EC monolayer from surgical handling during implantation, pointing out a potential drawback of preseeding cells. Coating ePTFE grafts with perlecan led to significantly greater EC coverage on grafts implanted in an ovine carotid artery model for 6 weeks.⁴⁶ While there was animal-to-animal variability resulting in 14–73% EC coverage, all coated grafts had significantly greater EC coverage than the corresponding uncoated control grafts. Enhanced EC retention is particularly exciting in large animal models as rodents tend to endothelialize more easily.⁵³

Heart valves,^54,55 flow diverters,⁵⁶ and stents^57–59 have also benefited from peptide and protein coatings examined in sheep,^54,55 dog,⁵⁶ pig,^57,59 and rabbit⁵⁸ studies. For example, decellularized sheep aortic valve conduits were coated with fibronectin or stromal cell-derived factor 1 alpha (SDF-1α) and implanted as a homograft pulmonary root replacement in sheep.⁵⁵ After 5 months, endothelialization on coated valves was ∼45% and 40% for the leaflet and wall, respectively, which was significantly greater than uncoated controls.

EC retention on p64 flow-diverter stents (FDSs) with different antithrombogenic hydrophilic coatings was investigated by Moreno et al.⁵⁶ After 28 days in a canine model, histological analysis showed near-complete endothelialization for the 81.3% of FDSs that remained patent. However, there was no significant difference in the EC loss between coatings. The study authors speculate that a longer implantation period or larger sample size could distinguish the effects of these individual coatings.

While biomaterial coatings have improved EC retention, the approach has not yet proven fully successful. This is primarily because denuded regions are highly susceptible to thrombosis due to exposure to circulating blood elements and platelet activation.⁶⁰ Furthermore, peptide coatings can actually negatively impact EC retention if the coated peptide detaches from its biomaterial substrate following adhesion with the cell surface through integrin receptors.⁶¹

Surface Modifications

Modification of the biomaterial surface is another method for improving EC retention and we have summarized the in vitro studies that used surface modification to augment EC adhesion (Table 2). While coatings serve as a mediating layer between the cells and biomaterial, surface modification is a more permanent alteration to improve biomaterial biocompatibility.

Table 2.

Summary of Biomaterial Surface Modification In Vitro Studies to Improve Endothelial Cell Adhesion and Retention

Biomaterial	Modification method	EC source and seeding	Perfusion	EC adhesion or retention	References
PU films	Micro-textured patterns etched by silicon wafer (95 μm wide and 32 μm deep channels)	BAECs seeded 24 h before flow	Parallel plate flow chamber with peristaltic pump. Glass slides with the PU films seeded with ECs were subjected to 60 dyn/cm² with laminar flow for 1 h	Cell densities (cells/mm²) Unpatterned PU: static 2206 ± 37 and flow 1269 ± 219 Patterned PU total area: static 2113 ± 292 and flow 1951 ± 464 Patterned PU plateau area: static 2104 ± 298 and flow 1490 ± 328 Patterned PU channel area: static 2121 ± 287 and flow 2345 ± 320 EC retention (%) Unpatterned PU: 58 Patterned PU: 92	Daxini et al.⁶⁶
Collagen Type I Films	Nanopatterning; linear nanogrooves with widths of 332.5, 500, and 650 nm	Human MVECs cultured 48 h before shear stress	Parallel plate flow chamber with shear stress of 12 dyn/cm² for 1 h	EC retention (%; mean ± SD) Unpatterned: 35 ± 10 332.5 nm nanopatterned: 75 ± 4 600 nm nanopatterned: 91 ± 5	Zorlutuna et al.⁶⁷
PDMS stamps coated with Fn	Microcontact printing with parallel or perpendicular patterns	HUVECs cultured for 24 h before shear stress	Parallel plate flow chamber with steady laminar shear stress of 15 dyn/cm² for 48 h	EC Retention (%; mean ± SD) Unpatterned 39.35 ± 6.18 Parallel: 73.92 ± 10.61 Perpendicular: 26.53 ± 4.03	Gong et al.⁶⁸
ePTFE vascular patch	Denucleation (air removal) of graft (D). Treatment with surfactant TDMAC. Followed by Fn or RGD (R) polymer soaking	MVECs seeded for 1 h	Centrifuged to expose grafts to 0.003 dyn/cell centripetal force	EC coverage of graft (mm²) D/TDMAC:20.7 ± 3.4 D/TDMAC/Fn: 20.7 ± 2.0 D/TDMAC/R: 8.4 ± 1.8 Fn: 3.6 ± 0.9 EC retention (%; mean estimated from graph) Fn: 20 D/Fn: 38 D/TDMAC: 78 D/TDMAC/R: 38 D/TDMAC/Fn: 80	Wigod and Klitzmen⁷⁰
ePTFE graft	Peptide modification with PFSP backbone and RGD, RGE (negative control for RGD), or Fn peptides	HPAECs seeded for 2 h	No perfusion; static culture	EC population ( × 10⁴ cells/cm²) RGE FSP polymer: ∼0.25 at 20, 48, and 120 h RGD FSP polymer: ∼0.6, 1, and 2.5 at 20, 48, and 120 h, respectively Fn: ∼0.75, 1, and 1.5 at 20, 48, and 120 h, respectively	Larsen et al.⁷¹
ePTFE	FSP and EC-selective peptide sequence Cys-Arg-Arg-Glu-Thr-Ala-Trp-Ala-Cys (CRRETAWAC) modification	HPAECs cultured for 18–22 h	Rotating disk system using 0–57.4 dyn/cm² shear stress for 4 h	EC retention (%; mean approximated from graph) Up to 38.2 dyn/cm² Fn control: 80 CRRETAWAC FSP: 99 RGD FSP: 99	Larsen et al.⁷²
PU graft	RGD	HUVECs seeded for 3.5 h; additional experiment with PU+gelatin+RGD seeded for a week before 3-h perfusion	Perfusion with shear stress of 20 dyn/cm² for 1 or 3 h	EC retention (%) PU: 1 h—38 ± 4 and 3 h—8 ± 2 PU+gelatin: 1 h—107 ± 19 and 3 h—43 ± 14 PU+gelatin+RGD: 1 h—119 ± 10 and 3 h—63 ± 5 PU+RGD: 1 h—59 ± 7 and 3 h—0.2 ± 0.1 EC retention (%) PU+gelatin+RGD seeded a week: 3 h—98 ± 1	Hsu et al.⁷³
PCU graft	Gly-Arg-Gly-Asp-Ser-Pro (GRGDSP) grafting by UV irradiation	HUVECs cultured for 24 h before flow	Parallel plate flow chamber with laminar flow to supply 2 Pa (20 dyn/cm²) shear stress for 2 h	EC retention (%) PCU: 1 PCU-GRGDSP: 92	Li et al.⁷⁴
PVA vascular graft	RFGD to modify the luminal surface with NH₃	HUVECs seeded for 16 days	No perfusion; in vitro cell adhesion study	EC quantification (cells/cm²; mean ± SE) 2 μm gratings without plasma: 7680 ± 2325 2 μm gratings with NH₃ plasma: 90,758 ± 18,323 1.8 μm concave lenses without plasma: 10,728 ± 3862 1.8 μm concave lenses with NH₃ plasma: 41,066 ± 38,492	Pohan et al.⁷⁵
PLA film	Immobilization of NMCS complexed with gelatin using UV photo-initiated surface grafting polymerization	HUVECs grown on PLA films for 7 days	Parallel plate flow chamber with average shear stress of 5 dyn/cm² for 24 h	EC retention (%) PLA: 54.77 Gel/NMCS-PLA: 87.10	Zhu et al.⁷⁶
Decellularized rat aorta ECM vascular graft	Covalent linking of CBP to heparin (CBP-heparin). CBPi is the inactive form of CBP	HUVECs seeded on the lumen for up to 28 days	No perfusion; static culture	No. of ECs/mg ECM (approximated from graph) PBS: 2,500 Heparin: 2,000 CBPi-heparin: 2,000 CBP-heparin: 8,000	Jiang et al.⁷⁷
Pyrolytic carbon prosthetic valve surface	Fn, collagen type I (Co), Ln, purified adhesion domain fragments GFOGER (GFOGER) and Fn_7–10 coating (Fn_7–10); microfabricated trenches	PAVECs	Parallel plate shear bioreactor system using shear stress from 0 to 600 dyn/cm² for 50 min	Relative to maximum EC adhesion at 1 h (%; mean approximated from graph) Co: 85 Fn: 80 Ln: 78 GFOGER: 70 Fn_7–10: 70 Control: 25 EC retention (visually with immunofluorescence) after 600 dyn/cm² SS for 48 h In between channels: 0 Within channels: confluent monolayer	Frendl et al.⁶⁹
PLLA	Pulsed plasma deposition of thin films of PVAA (PLLA-PVAA) and conjugation of Fn and VEGF	PAECs or PAECs transfected with KDR grown on the films for 5 days	Parallel plate flow chamber with a shear stress of 15 dyn/cm² for 30 min	EC retention (%; mean ± SD; approximated from graph of total DNA concentration) PAECs on PLLA: 25 PAECs on PLLA-PVAA-(EDC)-Fn-VEGF: 65 KDR-PAECs on PLLA: 50 KDR-PAECs on PLLA-PVAA-(EDC)-Fn-VEGF: 65	Xu et al.⁷⁸
Titanium (Ti)	Apatite (Ap), laminin-apatite composite (L40Ap) and AlbAp composite layers	HUVECs cultured for 1 week	Perfusion system with centrifugal blood pump using 0.7 Pa (7 dyn/cm²) shear stress for 1–2 h	EC retention (%; mean ± SD) 1 h SS Ti: 15.47 ± 29.74 Ap: 29.11 ± 24.16 AlbAp: 9.21 ± 6.67 L40Ap: 87.02 ± 8.76 2 h SS Ti: 5.88 ± 12.55 Ap: 6.10 ± 7.32 AlbAp: 0.31 ± 0.58 L40Ap: 56.04 ± 19.18	He et al.⁷⁹
PMP gas-exchange membranes	Pulse vacuum cathodic arc plasma deposition to coat with titanium dioxide (TiO₂)	HUVECs cultured for 24 h before flow	Parallel plate flow chamber with laminar shear stress. Primed with 0.5 dyn/cm² for 1 h before exposing to 30 dyn/cm² for 24 h	EC monolayer remained intact after fluid flow	Pflaum et al.⁸⁰
Hybrid membrane ventricular assist device	Hyperelastic RTV 4420 silicone with patterned membrane and honeycomb hexagonal walls	Ovine SVECs seeded for 3 days to endothelialize membrane	Cyclic actuation at 2 or 4 Hz with an average output of 2.5 or 4.0 L/min, respectively	Average normalized cell density approximated from graph Static: 1 2 Hz: 0.8 4 Hz: 0.7 No significant difference between groups and EC density was comparable to control	Ferrari et al.⁸¹

AlbAp, albumin-apatite; Ap, apatite; CBP, collagen binding peptide; KDR, kinase-insert domain-containing receptor; Ln, laminin; MVEC, microvascular endothelial cell; NMCS, N-maleic acyl-chitosan; PAEC, pig aortic endothelial cell; PAVEC, porcine aortic valve endothelial cell; PBS, phosphate buffered saline; PCU, poly(carbonate urethane); PDMS, polydimethylsiloxane; PFSP, peptide fluorosurfactant polymer; PLA, poly(lactic acid); PLLA, poly (l-lactic acid); PMP, poly(4-methyl-1-pentene); PU, polyurethane; PVA, polyvinyl alcohol; PVAA, poly (vinylacetic acid); SE, standard error; SVEC, saphenous vein endothelial cell; TDMAC, tridodecylmethylamimonium chloride; Ti, Titanium; UV, ultraviolet.

Many methods can be used to either alter the chemical or physical structure of the biomaterial or add onto the existing surface to achieve a new composition.^62,63 For example, plasma treatment is a method that allows chemical groups, including -OH or -NH₂, to be covalently attached to a biomaterial to create a surface more conductive to cell attachment.⁶⁴ Plasma treatment can also be used to alter the wettability of the material, which impacts cell adhesion. Studies have shown polymer surfaces with intermediate contact angles of 40–70° allow for optimal cell adhesion.⁶⁵

Micropatterning and nanopatterning have been explored to encourage EC retention. Both Polyurethane (PU)⁶⁶ and collagen type I films⁶⁷ modified with channels, showed improved EC retention under exposure to shear stress for 1 h. Similarly, parallel micropatterns on polydimethylsiloxane (PDMS) improved EC retention after laminar shear stress for 48 h.⁶⁸ On pyrolytic carbon valves, microfabricated trenches allowed for retention of a confluent EC monolayer within the channels after exposure up to 600 dyn/cm² shear stress for 50 min.⁶⁹ However, there were no ECs retained in between channels.

Another common surface modification is the incorporation of peptide or protein sequences into the biomaterial. Vascular grafts made from ePTFE,^70–72 PU,⁷³ poly(carbonate urethane) (PCU)⁷⁴ polyvinyl alcohol (PVA),⁷⁵ poly(lactic acid) (PLA),⁷⁶ and decellularized vessels⁷⁷ have all benefited from modification with peptides. Like with RGD coating, RGD incorporation has also led to superior EC retention under flow.^70,71,73,75 Incorporation of RGD into PU grafts with gelatin coating led to 63% ± 5% EC retention under perfusion with 20 dyn/cm² shear stress for 3 h compared to 8% ± 2% for PU or 43% ± 14% for PU+gelatin, respectively.⁷³

Deposition of thin films is another approach to encourage stronger EC adhesion.^78–80 Applying pulse vacuum cathodic arc plasma on poly(4-methyl-1-pentene) (PMP) gas-exchange membranes for extracorporeal membrane oxygenation allowed the EC monolayer to be retained after 24 h of 30 dyn/cm² shear stress.⁸⁰ A different method used a honeycomb hexagonal wall topography on a novel hybrid membrane for ventricular assist devices, which allowed endothelialized membranes to maintain cell coverage under cyclic actuation up to 4 Hz.⁸¹

Peptide functionalization on vascular grafts has also been examined in vivo using small^82–85 and large animal^86,87 models. In a rabbit carotid artery model, polycaprolactone (PCL) grafts modified with RGD exhibited an endothelialization rate of 51.1% ± 6.4% at 4 weeks compared to only 11.5% ± 3.2% in control grafts.⁸³ Similarly, PVA grafts functionalized with fucoidan exhibited 80% patency and EC coverage after 4 weeks in a rabbit carotid artery model compared to 75% and 40% for ePTFE and PVA controls, respectively.⁸⁵

Modification of ePTFE grafts with human anti-CD34 antibodies helped capture endothelial progenitor cells after implantation in a porcine carotid artery model. By 4 weeks, the EC coverage on modified grafts was 88% ± 5%, which was significantly greater than 32% ± 8% on bare grafts.⁸⁶ Interestingly, none of these studies preseeded the grafts with ECs.

Some surface modification approaches have proven advantageous to specifically target ECs using peptides or ligands immobilized to the biomaterial surface in vitro.⁸⁸ Furthermore, surface modifications such as plasma treatment are a relatively simple, scalable, reproducible method for modifying the surface of biomaterials.⁸² While improved EC retention has been reported in animal models, none of the studies noted formation of a fully confluent endothelial monolayer after implantation. It is possible a combination of surface modification with another approach may be required to achieve full endothelialization. In addition, surface modifications intended to promote EC adhesion will also tend to promote platelet adhesion at denuded regions.

Nanofiber Topography

Nanofiber scaffolds have emerged recently in the tissue engineering field as a platform for improving cell adhesion. The electrospinning technique results in biomaterials with desirable features, including excellent mechanical properties, augmentable fiber diameter and alignment, and fibrous structure that mimics natural ECM.^89,90 While nanofiber scaffolds offer a native-like ECM environment to enhance adhesion and retention of ECs, modification of nanofibers is typically necessary as synthetic polymers generally suffer from poor cell adhesion.⁸⁹ The use of nanofibers to develop scaffolds and conduits for seeding ECs has gained interest in recent years and we have summarized in vitro studies that used nanofiber topography to improve EC retention (Table 3).

Table 3.

Summary of In Vitro Studies Using Nanofiber Topography to Improve Endothelial Cell Adhesion and Retention

Biomaterial	Modification method	EC source and seeding	Perfusion	EC adhesion or retention	Reference
Polycaprolactone (PCL)/collagen type I nanofibers electrospun into flat sheet	Nanofiber alignment	HUVECs seeded for 3 days	Parallel plate flow chamber with shear stress of 20 or 40 dyn/cm² for 1 h	EC retention (%) 20 dyn/cm² Fully aligned: 95 Semialigned: 80 Random: 60 40 dyn/cm² Fully aligned: 62 Semialigned: 46 Random: 27	Whited and Rylander⁹¹

Whited and Rylander showed that nanofiber alignment influences EC adhesion.⁹¹ Confluent monolayers of ECs seeded on randomly oriented or aligned nanofiber scaffolds were exposed to 20–40 dyn/cm² shear stress for 1 h. Cell retention was lowest on randomly aligned scaffolds compared to scaffolds with aligned nanofibers, with ∼95% cell retention on aligned scaffolds and 60% cell retention on randomly oriented scaffolds.

Most studies have used animal models to show improved endothelialization with nanofiber topography. Vascular grafts are commonly fabricated using nanofiber scaffolds and modification of nanofibers has led to improved endothelialization in vivo.^92–97 For example, modification of high molecular weight poly (l-lactic acid) (PLLA) nanofibers with covalent attachment of heparin and vascular endothelial growth factor (VEGF) improved EC coverage in a rat common carotid artery model.⁹⁴ After 1 month, 88% of VEGF-modified grafts and 64% of heparin-modified grafts were patent compared to 55% for untreated grafts. The EC coverage after 2 weeks was 82% on VEGF-modified grafts compared to 39% on untreated and 40% on heparin-modified grafts.

Another study used a similar approach with VEGF and heparin, but used co-axial electrospinning to create a poly(L-lactide-co-epsilon-caprolactone) [P(LLA-CL)]/collagen/elastin graft.⁹⁷ Human aortic endothelial cells (HAoECs) were cultured on the grafts and implanted in an infrarenal aortic replacement model in rabbits for 28 days. The patency reported at 28 days for electrospun grafts was 86.6% compared to 40.0% for ePTFE. Combining natural proteins with a synthetic polymer using co-axial electrospinning allows for improved biocompatibility and endothelialization, while maintaining sufficient mechanical properties to withstand physiologic blood pressures.

Nanofiber scaffolds have shown great promise for vascular graft applications and more long-term studies in large animal models should be used to confirm the successful outcomes reported in rodent and rabbit models. It should also be noted that there are numerous factors that may contribute to enhanced EC adhesion to nanofiber scaffolds, including porosity, nanofiber alignment, and biochemical cues.⁹⁸ Therefore, researchers can explore many parameters in fabricating nanofiber scaffolds to better promote EC retention.

Shear Stress and Preconditioning

Fluid shear stress, or the frictional force along the vessel wall created by the flow of blood, is a vital consideration in EC retention. ECs are continuously exposed to pulsatile shear stress due to blood flow, and shear stress modulates the expression of genes for numerous proteins, including vasoactive substances, adhesion molecules, and coagulation factors.⁹⁹ Shear stress preconditioning, in which ECs are subjected to flow before implantation, is a method to improve EC adhesion by stimulating formation and strengthening of focal contacts.¹⁰⁰ Studies have shown that exposing ECs to laminar shear stress causes the cells to align in the direction of flow and flatten, which reduces shear stress gradients.¹⁰¹

Furthermore, shear stress preconditioning leads to lower expression of proproliferative molecules and promotes adhesion of ECs in vitro.¹⁰² Laminar shear stress has also been shown to upregulate integrin expression in ECs, which is beneficial for cell adhesion.¹⁰³ Many in vitro studies report the benefit of shear stress preconditioning on EC retention under flow (Table 4).

Table 4.

Summary of Shear Stress Preconditioning In Vitro Studies to Improve Endothelial Cell Adhesion and Retention

Biomaterial	EC source and seeding	Shear stress regimen	EC adhesion or retention	References
ePTFE vascular graft coated with Fn	Autologous ovine carotid artery ECs labeled with 35S-methione cultured for 48 h before flow or exposed immediately to flow	Grafts were exposed to nonpulsatile flow with stepwise increases in the flow rate from 33 to 330 L/min, which corresponded to 3.3–16 dyn/cm² shear stress. After 2 min at 33 mL/min, each subsequent step occurred for 10 min for a total of 5 steps leading up to 330 mL/min. Total of 50-min flow.	EC retention (%; mean ± SD) 0-min culture: 96.3 ± 1.8 48-h culture: 98.7 ± 0.2	James et al.²⁵
PU vascular graft	BAECs	Cultured ECs in vitro for 6 days with or without continuous laminar shear stress. Three days at 1–2 dyn/cm² and then 3 days at ∼25 dyn/cm². Then exposed the grafts to 25 dyn/cm² for 25 s to examine cells lost.	Average EC loss (No. of cells; mean ± SEM) Static grafts: 1.35 × 10⁶ ± 0.44 × 10⁶ Preconditioned: 1.05 × 10⁴ ± 0.16 × 10⁴	Ott and Ballermann¹⁰⁴
PU vascular graft and porcine valves (Medtronic Freestyle)	SVECs cultured for × days	Adaptation phase using flow of 0.9 ± 0.3 L/min (systolic pressure: 40–50 mmHg). High flow was 3.2 ± 0.6 L/min for over 4 h. Vascular graft groups were: (1) high flow immediately (2) adaptation phase of 15-min (3) adaptation phase of 30 min. Valve groups were as follows: (1) high flow immediately (2) adaptation phase of 30 min.	Approximate EC coverage (%; mean) Vascular graft group 1: 60 Vascular group 2: 82 Vascular group 3: 100 Valve group 1: 10 on the ventricular side of leaflets, 35 on aortic side of leaflets, and 75 on free wall Valve group 2: 90	Gulbins et al.¹⁰⁵
Bioartificial tissue from neonatal human dermal fibroblasts entrapped in tubular fibrin gels	HBOECs or HUVECs seeded for 24 h	Perfusion system with peristaltic pump. Incrementally exposed to 1, 5, 10, 15, and 25 dyn/cm² for 20-min	EC coverage (%) after 25 dyn/cm² SS exposure (%; mean ± SEM) Type I collagen coated tissue culture plate: 82.6 ± 1.2 Bioartificial tissue: 94.4 ± 0.6	Ahmann et al.¹⁰⁶
Decellularized human umbilical vein scaffolds	HUVECs	Perfusion with peristaltic pump. Seeded grafts were grown for 5 days under 0.007 Pa (0.07 dyn/cm²). Media viscosity was increased to 3 mpA × s and luminal flow rate was linearly increased over 48 h from 0.03 to 1.45 Pa to achieve a mean WSS of 0.734 Pa (7.34 dyn/cm²).	Cell density (cells/mm²) Day 5: 738 ± 178 Day 6: 702 ± 95 Day 7: 619 ± 172 EC coverage (%) Day 5: ∼77 Day 6: 97.3 ± 3.6	Uzarski et al.¹⁰⁷
Silk fibroin nanofibrous scaffold	HUVECs seeded to confluence before shear stress experiments	Parallel plate flow chamber. Step 1 flow (4 dyn/8 h): instantaneously 1.2–4 dyn/cm², then increased stepwise by 4 dyn/cm² until 16 dyn/cm² over 8 h. Step 2 flow (10%/h): increasing stepwise by 10% relative to the SS of previous hour, sustained for 1 h until reaching 15.7 dyn/cm². Step 3 flow (15%/h): increasing stepwise by 15% relative to SS of previous hour, sustained for 1 h until reaching 14.9 dyn/cm². Step 4 flow (modified): increased stepwise by 10% relative to SS of previous hour, sustained for 1 h. After 6.8 dyn/cm² preconditioning, sustained period of 2 h until SS of 14.9 dyn/cm² was achieved.	Average EC retention (%; mean ± SD) 4 dyn/8 h over 24 h:38.58 ± 7.55 10%/h over 24 h: 45.82 ± 5.98 15%/h over 16 h: 2.04 ± 0.22 Modified over 32 h: 99.31 ± 4.97	Liu et al.¹⁰⁸
Decellularized porcine pulmonary valves	Differentiated ECs from porcine peripheral blood mononuclear cells	Computer-assisted pulsatile bioreactor system that controlled pressure and flow was used for preconditioning. After static seeding, pulsatile flow was applied. The rate was gradually increased through 10 steps up to 1, 2, or 3 L/min with an incremental rate increase of 0.1, 0.2, or 0.3 L/min for each, respectively. Preconditioning occurred for 7 days (each step 16.8 h).	EC coverage after 7 days of preconditioning (%; mean estimated from graph) 1 L/min: 80 2 L/min: 90 3 L/min: 2	Lee et al.¹⁰⁹
Silicone hollow fibers modified with collagen (siHF-Col patches) for membrane oxygenators	HUVECs seeded for 48 h before flow	Cells were preconditioned with 12 dyn/cm² shear stress in a parallel plate flow chamber for 24 h. The flow rate started at 10 mL/min and was increased by 10 mL/min every 15 min. After preconditioning, the patches were exposed to 30 dyn/cm² shear stress for 1 h.	EC retention (%) No preconditioning: 72.8 With preconditioning: 96.7	Salehi-Nik et al.¹¹⁰

HBOEC, human blood outgrowth endothelial cell; SEM, standard error of the mean; WSS, wall shear stress.

Shear stress preconditioning is commonly used to improve EC retention on vascular grafts.^25,104–108 In an early study by Ott and Ballermann, bovine aortic endothelial cells (BAECs) were seeded onto PU vascular grafts and cultured for 6 days with or without continuous laminar shear stress, which gradually increased from 1–25 dyn/cm².¹⁰⁴ The study found that grafts preconditioned with shear stress lost significantly fewer cells during a short pulse of 25 dyn/cm² shear stress (1.05 × 10⁴ vs. 1.35 × 10⁶ cells lost for preconditioned and non-preconditioned, respectively).

Gulbins et al. showed both PU vascular grafts and porcine valves exhibited greater EC coverage when the devices underwent a flow adaptation phase of 30 min with 0.9 ± 0.3 L/min before exposure to 3.2 ± 0.6 L/min for 4 h.¹⁰⁵ The approximate EC coverage was close to 100% on the grafts and 90% on the valves. In a related study, decellularized porcine pulmonary valves benefited from a gradual 10-step increase of 0.2 L/min up to 2.0 L/min over 7 days.¹⁰⁹ The EC coverage was ∼90% after 1 week. Silicone hollow fibers modified with collagen for membrane oxygenators have also showed enhanced EC retention with shear stress preconditioning.¹¹⁰

Liu et al. reported shear stress preconditioning can regulate EC tolerance to shear stress.¹⁰⁸ Human umbilical vein endothelial cells (HUVECs) cultured on silk fibroin nanofiber scaffolds that were flow conditioned with a 10%/h increase showed a 55% increase in retention when plated on fibronectin-coated plates compared to HUVECs in static culture. The results revealed that a gradual increase in shear stress with appropriate time-step and amplification improved EC retention, and the mechanism is mediated partly through integrin β1 and focal adhesion kinase (FAK) expression.

Animal studies validate shear stress preconditioning for improving EC retention. Pretreatment of ECs with 25 dyn/cm² shear stress led to retention of confluent EC monolayers after implantation using a rat model.¹⁰² In a porcine femoral artery model, pre-endothelialized ePTFE grafts conditioned with a steady flow increase followed by pulsatile flow retained an EC lining after 6 weeks of implantation.¹¹¹

Preconditioning ECs generally enhances EC adhesion to the biomaterial substrate and allows ECs to form a continuous monolayer under perfusion conditions. However, the drawbacks of this approach include lengthy culture times with added costs and challenges associated with maintaining the complex bioreactor environment, including sterility.¹¹² These drawbacks may limit the clinical translatability of this method to endothelialize blood-contacting devices.

EC Modifications

In contrast to methods to improve the adhesive properties of the substrate onto which ECs are seeded on, other groups have focused on enhancing the adhesive properties of ECs to improve retention on biomaterials. This can include changes at the genetic level or controlling the phenotype of the EC through environmental changes. Multiple EC modification approaches have shown improved EC adhesion and retention in vitro (Table 5).

Table 5.

Summary of Endothelial Cell Modification In Vitro Studies to Improve Endothelial Cell Adhesion and Retention

Biomaterial	Modification method	EC source and seeding	Perfusion	EC adhesion or retention	References
ePTFE	Retroviral transduction of ECs to express fibulin-5 alone or fibulin-5 and VEGF	HUVECs seeded for 48 h before flow	Pulsatile flow conditions with a pressure of 120/80 mmHg and shear stress of 7–10 dyn/cm² for 2 h	Approximate EC retention (%) Nontransduced:65 Transduced for GFP: 55 Transduced for fibulin-5: 90 Transduced for fibulin-5+VEGF: 85	Preis et al.¹¹³
ePTFE	Gene transduction with fibulin-5 and VEGF	Porcine SVECs rotationally seeded for 48 h before flow	Pulsatile flow with a pressure of 120/80 mmHg and shear stress of 7–10 dyn/cm² for 2 h	EC retention (%) Nontransduced cells: 80, 69, and 76 for each experiment Transduced cells: 95, 96, and 100 for each experiment	Tzchori et al.¹¹⁴
ePTFE	siRNA-induced suppression of SHP-1 or SHP-2	HUVECs seeded for 4 h before flow	Parallel plate flow chamber with an average shear stress of 15 dyn/cm² for 6 h	Approximate EC retention (%) scrambled siRNA-treated controls: 30 SHP-1 knockdown: 70 SHP-2 knockdown: 85	Tefft et al.¹¹⁶
PET films	ECs were grown in low-O₂ tension using a 5%O₂/5%CO₂/90%N₂ culture environment or normal O₂ conditions using a standard 5% CO₂ incubator	HUVECs grown for 3 days before flow	Parallel plate flow chamber with laminar flow using shear stresses 20 dyn/cm² at 20% O₂ for up to 24 h	EC retention (%) After 6 h Low: approximately 95 Normal: 93.3 After 24 h Low: ∼95 Normal: 77.5	Zhao et al.¹¹⁷

SHP, src homology-2 domain containing protein tyrosine phosphatase; siRNA, small interfering RNA.

Fibulin-5 is a secreted glycoprotein that helps to form and stabilize proteins of the ECM and elastic fibers.¹¹³ It also contains an RGD sequence, which promotes cellular adhesion through integrin interactions. Preis et al.¹¹³ and Tzchori et al.¹¹⁴ showed that gene transfer with vectors encoding fibulin-5 improves EC retention under flow in vitro. HUVECs exhibited 85% retention after perfusion with shear stress of 7–10 dyn/cm² for 2 h compared to 65% or 55% for nontransduced cells and GFP-transduced cells, respectively. Tzhori reported an average EC retention of 97% for fibulin-5-transduced cells compared to 75% for nontransduced cells.¹¹⁴

Small interfering RNA (siRNA) is an approach that offers modulation of the cell at the molecular level.¹¹⁵ Tefft et al. used siRNA to silence expression of src homology-2 domain containing protein tyrosine phosphatase (SHP)-1 and SHP-2, which are phosphatases known to inhibit focal contact formation.¹¹⁶ Both SHP-1 and SHP-2 silencing led to increased EC retention on ePTFE using an average 15 dyn/cm² shear stress for 6 h. Untreated and scrambled siRNA-treated cells showed significantly lower EC retention of ∼30% compared to ∼70% and 85% for the SHP-1 and SHP-2 siRNA-treated cells, respectively. Future in vivo studies need to be conducted to confirm these positive effects

Another approach used low-oxygen pretreatment of the ECs to improve retention.¹¹⁷ PET films seeded with ECs grown in low-O₂ tension exhibited stable EC retention under 20 dyn/cm² shear stress for up to 24 h, while samples seeded with ECs grown in normal O₂ tension lost ∼20% of cells.

EC modification studies have also been examined in animal models. Lahtinen et al. showed delivering plasmids for human-VEGF improved endothelialization and patency in a canine femoral artery model.¹¹⁸ The in vitro findings by Preis et al.¹¹⁹ and Tzchori et al.¹¹⁴ were also confirmed in vivo. Autologous ECs were transduced with fibulin-5 through retroviral vector, seeded onto ePTFE grafts, and implanted into sheep for 3–6 months.¹¹⁹ At 2 weeks, grafts seeded with modified ECs demonstrated 73% coverage, while grafts seeded with control ECs had 45% coverage. At explant, 8/8 grafts with modified ECs were patent, while only 3/10 control grafts were patent.

Tzchori et al. confirmed improved EC retention and patency of ePTFE hemodialysis grafts in a porcine model for 6 months.¹¹⁴ Approximately 80% of grafts seeded with transduced ECs remained patent at 16 weeks, which was significantly greater than the 29% of control grafts that remained patent.

While modifications at the genetic level improve EC retention, there is concern over the use of retroviral vectors for the transduction of human cells. Preis et al. performed full postmortem analysis on all sheep implanted with vascular grafts seeded with modified ECs expressing fibulin-5.¹¹⁹ The analysis did not reveal any pathological findings due to the use of genetically modified ECs for seeding the grafts. However, safety will need to be carefully studied before moving the approach into human trials.

Conclusions

This review highlights attempts to augment EC retention on blood-contacting devices. There have been countless studies showing different approaches to improve EC adhesion, but we chose to focus on the reports that quantified EC retention on biomaterials under flow conditions.

While progress has been made to improve cell retention, there are still no direct solutions for achieving a stable EC monolayer on the blood-contacting surface of cardiovascular devices in a clinical setting. Promising results from in vitro studies testing cell retention under shear stress and flow conditions need to be confirmed using large animal models to generate convincing preclinical data. Long-term studies in large animal models must then be translated into human trials to demonstrate efficacy (Table 6). Future work identifying strategies to maintain a stable EC monolayer on blood-contacting devices after implantation is critical for improving clinical outcomes in patients.

Table 6.

Summary of In Vivo Animal Studies Targeting Endothelial Cell Adhesion to a Blood-Contacting Surface

Modification method	Biomaterial	Adhesion target	Cell type	Animal model	Major findings	References
Protein coating	ePTFE vascular graft	Fn	Autologous canine carotid ECs radiolabeled with indium-111	Canine carotid interposition for up to 72 h	There was a rapid loss of ECs in the first 30 min of flow with an EC retention of 78.6% ± 10.0% on Fn-coated grafts compared to 70.2% ± 5.4% on uncoated grafts. After 24 h of implantation, Fn-coated grafts retained 45.7% ± 10.0% of ECs and uncoated grafts retained 17.0 ± 6.2%.	Ramalanjaona et al.⁴⁴
Protein coating	Knitted polyester vascular grafts	Collagen and dermatan sulfate	Autologous canine external jugular vein ECs	Canine carotid artery	One graft model also contained cultured SMCs. After 2 weeks, EC coverage was 92% on the lumen of grafts containing SMCs and complete coverage was observed on grafts without the SMCs added during seeding. At 12 weeks, both grafts looked similar in EC coverage.	Miwa et al.⁴⁵
Protein coating	PFTE vascular graft	Fn or naturally produced ECM	Autologous SVECs	Canine common carotid artery	Significantly more ECs were retained on Fn+ECM-coated grafts with 85% retention compared to 20% for uncoated or 54% for Fn coated after shear stress in vitro. Significant increase in EC retention after implantation and reduced incidence of thrombus formation on Fn+ECM-coated grafts.	Schneider et al.²⁸
Protein coating	ePTFE vascular graft	Perlecan coating to bind growth factors and interact with ECM proteins	Unseeded	Ovine carotid artery	EC coverage (%) varied among the five sheep. For uncoated grafts, it ranged from 9% to 64%. For perlecan coated grafts, it ranged from 14% to 73%. In all sheep, the perlecan coated grafts exhibited significantly greater EC coverage compared to uncoated control grafts.	Lord et al.⁴⁶
Protein coating	Knitted polyester graft precoated with glutaraldehyde cross-linked or non-cross-linked collagen	Fn or SDF-1α	Unseeded	Ovine carotid artery	Significantly increased EC coverage on the Fn-SDF-1α grafts with 48% ± 4% compared to 27% ± 4% in controls.	De Visscher et al.⁴⁷
Protein coating	Decellularized rat aorta	Fn	Unseeded	Rat infrarenal aorta	After 8 weeks, the lumen of Fn-coated grafts showed significantly greater EC coverage with 89.9% ± 5.45% compared to 73.6% ± 13.14% for controls. In the ascending aorta region, the Fn-coated grafts had 81.90% ± 9.34% EC coverage and controls had only 34.47% ± 10.58%.	Assmann et al.⁴⁹
Protein coating	dEAC	Matricellular protein CCN1, which interacts with α_vβ₃ integrins	Unseeded	Ovine carotid artery	After 14 weeks, the number of ECs in the middle of the graft was ∼180, 220, and 350, for PTFE, dEAC, and dEAC+CCN1, respectively. At the anastomoses, there was a 9.8-fold increase in ECs on dEAC+CCN1 compared to ePTFE. Patency was 3/3 for both dEAC and dEAC+CCN1 compared to 2/6 for ePTFE.	Böer et al.⁴⁸
Protein coating	Decellularized rat aorta vascular graft	VEGF conjugated to temperature-sensitive hydrogel (Hg)	Unseeded	Rat aorta	Significant increase in endothelium formation in the Hg-VEGF grafts with 64.8% ± 7.6% luminal surface coverage at 4 weeks compared to 40.4% ± 8.3% in uncoated controls. Nearly complete endothelial layer by 8 weeks, but was similar in both Hg-VEGF and uncoated grafts.	Iijima et al.⁵⁰
Peptide coating	ePTFE vascular graft	Heptamaltose (M7) to block nonspecific protein adsorption, RGD peptide or EC-selective CRRETAWAC peptide (cRRE)	PPAECs	Porcine carotid artery	After 1 month, one of three uncoated grafts, two of three RGD-coated, two of four M7-coated, and two of four cRRE-coated grafts remained patent. EC coverage on the lumen of all patent grafts.	Bastijanic et al.⁵²
Peptide coating	ePTFE vascular graft	SP neuropeptide-conjugated PLCL (SP-PLCL) and Hep-PLCL coatings	Unseeded	Rat subcutaneous implantation	After 2 weeks, the EC density (um²/mm²) was 6589.40 ± 424.93 in Hep/SP-PLCL, 4463.82 ± 742.51 in SP-PLCL, 960.83 ± 86.88 in Hep-PLCL, and 81.26 ± 69.89 in ePTFE. At 4 weeks, the EC density of Hep/SP-PLCL was 1.86 times greater than Hep-PLCL and 1.15 times greater than SP-PLCL.	Kim et al.⁵¹
Protein coating	Photo-oxidized bovine pericardium bioprosthetic heart valves	Fn or SDF-1α matrix coating by impregnation	Unseeded	Pulmonary position of sheep	Saw complete recellularization of the leaflets at 5 months. An EC layer was evident with eNOS immunofluorescence staining.	De Visscher et al.⁵⁴
Protein coating	Decellularized ovine aortic valve conduits	Fn or SDF-1α	Unseeded	Pulmonary valve in sheep	After 5 months, coated valves exhibited significantly greater endothelialization, which was ∼45% and 40% in the leaflet and wall, respectively, compared to ∼18% and 10% in the leaflet and wall, respectively, for uncoated controls.	Flameng et al.⁵⁵
Peptide coating	p64 FDS	PEG, glycan, or polyphosphazene antithrombogenic hydrophilic coatings	Unseeded	Canine common carotid arteries, external carotid arteries, and subclavian arteries	Near-complete endothelialization after 28 days. 81.3% of implanted FDSs remained patent. No significant difference in EC loss with any applied coating.	Moreno et al.⁵⁶
Peptide coating	Stainless steel stent	RGD peptide	EPCs infused into coronary arteries right after stent implantation	Porcine coronary artery	Four weeks after implantation, the RGD-loaded stents had significantly greater EC coverage with 83% ± 3% ECs/field compared to only 35% ± 2% and 36% ± 3% for unloaded and bare metal stents, respectively. Twelve weeks after implantation, the EC coverage was comparable for all groups.	Blindt et al.⁵⁷
Protein coating	Stainless steel stent	Human VE-cadherin	EPCs exposed to stents for 30 min before implantation	Rabbit iliac arteries	After 3 days, there were significantly more ECs on the VE-cad stents. Greater than 90% of VE-cad stent surface was covered with ECs and <10% of the bare stent surface had ECs.	Lim et al.⁵⁸
Peptide coating	Bare metal stents	Liposomes carrying an NO donor (DETA NONOate) were coated on NO-releasing stents with a LBL approach	Unseeded	Porcine anterior descending artery and coronary artery	After 28 days, there was significantly greater EC coverage (94% ± 4%) on LBL-coated stent compared to 34% ± 3% coverage on bare metal stent.	Elnaggar et al.⁵⁹
Plasma treatment	PCL vascular graft	Surface biocompatibility and lower water contact angle	Unseeded	Infrarenal aortic replacement in rat	No significant difference in EC coverage on lumen, 30.3% EC coverage for untreated grafts and 33.2% for treated grafts after 3 weeks in vivo.	de Valence et al.⁸²
Peptide functionalization	ePTFE vascular graft	Biopolymer covalently immobilized to polymer and anti-human CD34 antibody covalently coupled to the peptide	Unseeded	Porcine carotid artery	Three days postimplantation, EC coverage was significantly greater on coated grafts than uncoated, ∼95% and <5%, respectively. At 4 weeks, the coated grafts still had significantly greater EC coverage at 88% ± 5% compared to 32% ± 8% on bare grafts.	Rotmans et al.⁸⁶
Peptide functionalization	PCL vascular graft	RGD and Nap-FFGRGD molecule	Unseeded	Carotid artery in rabbit	Endothelialization rates of PCL-RGD grafts were 27.2% ± 11.5% and 51.1% ± 6.4% at 2 and 4 weeks, respectively. This was faster than rates of 1.8% ± 1.1% and 11.5% ± 3.2% at 2 and 4 weeks, respectively, in control PCL grafts.	Zheng et al.⁸³
Peptide functionalization	Decellularized rat aortic grafts	Cell adhesive sequences (RGD, REDV, YIGSR)	Unseeded	Rat aorta	After 14 days, the EC coverage of RGD/REDV-functionalized graft lumen was 25.8% ± 9.3% compared to 19.1% ± 14.1% in controls. However the difference between groups was not significant.	Aubin et al.⁸⁴
Protein functionalization	PVA vascular graft	Functionalization with fucoidan (Fg), a polysaccharide with anticoagulant and antithrombotic properties	Unseeded	Carotid artery implantation in rabbit	After 1 month of implantation, the patency rate of ePTFE, PVA, and PVA-Fg grafts was 75%, 40%, and 80%, respectively. Immunofluorescent staining showed cells positive for EC markers on the PVA-Fg graft lumens.	Yao et al.⁸⁵
Protein functionalization	PU pulmonary valve	Bromoalkylation to covalently modify the PU with cholesterol	Autologous sheep BOECs	Pulmonary valve leaflet replacement in sheep	The number of ECs covering the leaflets was ∼110 cells/1.13 mm² after 30 days and ∼160 cells/1.13 mm². Immunostaining confirmed presence of ECs.	Stachelek et al.⁸⁷
Nanofiber scaffold	P(LLA-CL) vascular graft	Heparin-bonded nanofibers	Autologous canine femoral vein ECs or unseeded	Femoral artery graft in canine (Beagles)	Non pre-endothelialized grafts had patency <15% by 24 weeks. Patency unseeded heparin-loaded grafts was 66.7% compared to 37.5% for pre-endothelialized normal [P(LLA-CL)] grafts. Significant improvement to 85% patency after 24 weeks in pre-endothelialized heparin-loaded grafts.	Wang et al.⁹²
Nanofiber scaffold	PLLA, PLCG, PCL vascular graft with chemical modification	Mucin (glycoprotein lubricant) modification with passive adsorption (MUC) or conjugation (PEG-MUC)	Unseeded	Common carotid artery in rat	After 1 month, 50% of untreated grafts, 50% of PEGylated (negative controls), 50% of MUC, and 100% of PEG-MUC grafts were patent. Continuous endothelialization was present on lumen of all patent grafts.	Janairo et al.⁹³
Nanofiber scaffold	High MW PLLA with LMW PCL to make LMW vascular graft	Covalent attachment of heparin and immobilization of VEGF	Unseeded	Common carotid artery in rat	After 1 month, 55% of untreated, 64% of heparin-modified, and 88% of VEGF-modified grafts were patent. EC coverage after 2 weeks in VEFG-modified grafts was 82% compared to only 39% in untreated and 40% in heparin-modified grafts.	Henry et al.⁹⁴
Nanofiber scaffold	Fibrin tube with PCL sheath vascular graft	Electromechanical stretching of sheath	ECFCs	Abdominal aorta of mice	Fully endothelialized luminal surface after 1 week of implantation. No significant difference on EC coverage between implanting acellular or endothelialized grafts.	Elliott et al.⁹⁵
Nanofiber scaffold	PCL/fibrin composite vascular graft	Fibrin	Unseeded	Abdominal aorta of rats	At 1 month, the EC coverage was ∼100%, 60%, and 40% for native vessel, PCL/fibrin, and PCL grafts, respectively. By 9 months, the EC coverage of PCL/fibrin grafts was ∼90% compared to 75% for PCL and 100% for native vessel.	Zhao et al.⁹⁶
Co-axial nanofiber scaffold	P(LLA-CL) collagen and elastin vascular graft	Heparin and VEGFG-165 encapsulated in the fibers	HAoECs	Infrarenal aorta of rabbit	Around 83.26% of ECs adhered on electrospun grafts after 2 h, which was significantly >46.87% for ePTFE controls. 13 out of 15 coaxial electrospun grafts remained patent, while 6 out of 15 ePTFE control grafts remained patent.	Hu et al.⁹⁷
Shear Stress preconditioning	PU vascular graft	Pretreated ECs to increasing shear stress in vitro before implantation	RAECs	Aortic interposition of rats	Pretreatment of ECs with 1 dyn/cm² shear stress for 3 days followed by 3 days with 25 dyn/cm² shear stress before implantation led to formation of a confluent EC monolayer within 24 h of implantation in vivo. Stable up to 3 months after implantation.	Dardik et al.¹⁰²
Shear Stress preconditioning	ePTFE vascular graft	Bioreactor was used to perfuse the grafts. Steady increase of flow up to 15 dyn/cm² over the first 24 h, then pulsatile flow until confluent EC layer formed	Autologous pig external jugular vein ECs	Porcine femoral artery	7/12 Endothelialized, preconditioned grafts maintained an EC lining after 6 weeks of implantation.	Büttemeyer et al.¹¹¹
EC modification	ePTFE vascular graft	pNGVL3-VEGF	Plasmid solution injected on graft	Rat abdominal aorta, rabbit abdominal aorta, and canine carotid and femoral arteries	For the rat study, VEGF transfection enhanced mid-graft endothelialization compared to controls up to 18 months with 100% completely endothelialized area compared to 75% in controls. In the rabbit study, there was no difference between VEGF and control grafts up to 43 weeks. For the canine, all carotid grafts were patent at 6 weeks. The femoral grafts with VEGF had improved patency compared to controls.	Lahtinen et al.¹¹⁸
EC modification	ePTFE vascular graft	Transduction with fibulin-5 and VEGF by retroviral vector	Autologous ECs from sheep hind limb vein	Carotid artery of sheep	At 2 weeks, the 73% EC coverage on grafts seeded with modified ECs was significantly >45% EC coverage on grafts with control ECs. At 3 months, all grafts seeded with the modified ECs were patent and only a third of control grafts remained patent. After 6 months, similar results were observed.	Preis et al.¹¹⁹
EC modification	ePTFE hemodialysis graft	Transduction with fibulin-5 and VEGF	Autologous ECs from porcine saphenous vein	Porcine carotid artery and jugular vein	Improved EC retention and patency in ePTFE grafts seeded with autologous transduced ECs compared to controls in a porcine model for 6 months. Around 80% of grafts seeded with transduced ECs were patent at 16 weeks, which was significantly >29% for control grafts.	Tzchori et al.¹¹⁴

BOEC, blood outgrowth endothelial cell; dEAC, decellularized equine carotid artery; ECFC, endothelial colony-forming cell; eNOS, endothelial nitric oxide synthase; EPCs, endothelial progenitor cells; FDS, flow-diverter stent; Hep-PLCL, heparin-conjugated PLCL; Hg, hydrogel; LBL, layer-by-layer; LMW, low molecular weight; MUC, mucin; MW, molecular weight; NO, nitric oxide; PEG, polyethylene glycol; PLCG, poly(l-lactide-co-caprolactone-co-glycolide); PLCL, poly (l-lactide-co-ɛ-caprolactone); P(LLA-CL), poly(l-lactide-co-epsilon-caprolactone); PPAEC, porcine pulmonary artery endothelial cell; SDF-1α, stromal cell derived factor 1 alpha; SMC, smooth muscle cell; SP, substance P; VE, vascular endothelial.

Footnotes

Authors' Contributions

J.T.W. and B.J.T.: conceptualization of topic. J.T.W.: article writing. A.S.: figure preparation. J.T.W., B.J.T., and A.S.: editing.

Disclosure Statement

The authors declare no conflicts of interest.

Funding Information

The authors gratefully acknowledge funding by the Advancing a Healthier Wisconsin Research and Education Program at the Medical College of Wisconsin.

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