Blood Compatible Polymers

  • Sara AlibeikEmail author
  • Kyla N. Sask
Living reference work entry
Part of the Polymers and Polymeric Composites: A Reference Series book series (POPOC)


Medical devices made from polymeric materials come in contact with blood in a wide range of applications, including stents, artificial vascular grafts, hemodialysis membranes, catheters, and sutures, among others. In this chapter, an overview of the ongoing investigations with blood compatible polymers is provided. A summary of polymers used in blood contacting devices will be given, followed by details focusing on each of the types of polymers that are most commonly used. Furthermore, a description of the efforts made in improving the blood compatibility of these polymers will be provided, as most synthetic polymers are required to go through some level of modification in order to be used in blood contacting devices. Most modification strategies address the changes in surface properties of these polymers with the aim of controlling the interactions between blood components and the polymeric surface. Among these modification techniques, use of bioinert molecules, bioactive molecules, and a combination of the two molecules are the subject of most studies.

List of Abbreviations


Albumin-coated vascular graft


Atomic force microscopy


Activated partial thromboplastin time








Bovine serum albumin






Conjugated linoleic acid




Castor-oil-mono- hydrogenated acetates






Di(2-ethylhexyl) terephthalate


Cyclohexane 1,2-dicarboxylate


Di-iso-nonyl phthalate


Endothelial cell(s)


Extracorporeal circulation




Endothelial progenitor cells


Ethylene vinyl alcohol copolymer


Glycidyl methacrylate


Human serum albumin


Low-molecular-weight heparin


Left ventricular assist device




2-methacryloyloxethyl phosphorylcholine


Molecular weight(s)


Nitric oxide






Poly(acrylonitrile-co-maleic anhydride)






Polyethylene glycol


Poly(ethylene glycol) methacrylate


Polyethylene oxide




Polyethylene terephthalate








Poly(2-methoxyethyl acrylate)






Poly hydroxyl-ethylmethacrylate


Poly(2-methacryloyloxyethyl phosphorylcholine)


Poly(oligo(ethylene glycol) methacrylate)




Poly(propylene oxide)


Plasma recalcification time






Polytetramethylene oxide








Poly(vinyl alcohol)




Polyvinylidine fluoride




Recombinant hirudin


Surface initiated atom transfer radical polymerization




Segmented polyurethane


Segmented polyurethanes






Tri-2-ethylhexyl trimellitate


Tris-octyl tri-mellitate


Tissue plasminogen activator


Unfractionated heparin

1 Overview of Blood Compatible Polymers

Polymeric biomaterials have widely been used in blood contacting devices, with the primary use in cardiovascular applications. These include both devices in contact with blood for a short time such as catheters and hemodialysis membranes, as well as long-term blood contacting devices such as stents, vascular grafts, left ventricular assist devices, and heart valves.

A list of synthetic polymers extensively tested and used in blood contacting devices is provided in Table 1. Various modification strategies have been carried out on these polymers and a summary of some of the main references is given. These studies will be discussed throughout each polymer section of this chapter.
Table 1

Synthetic polymers used in common blood contacting devices



Modification strategies [Reference numbers]

Polyurethane (PU)

Catheters, left ventricular assist devices (LVAD), pacemaker lead insulation, artificial vascular grafts

[2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 46, 47, 48, 49, 154]


Catheters, blood pumps, heart valve structures, pacemaker leads, blood oxygenators, tubing, microfluidic devices

[51, 52, 53, 54, 55, 56, 57, 58, 59, 60, 61, 62, 155]

Polytetrafluoroethylene (PTFE) (Teflon)

Artificial vascular grafts, catheters

[64, 65, 66, 67, 68, 69, 70, 71, 72, 73, 74, 75, 76, 77, 78, 79, 80, 81, 82, 83, 84, 85, 86, 158]

Polyethylene terephthalate (PET) (Dacron)

Artificial vascular grafts, sewing rings on artificial heart valves, fabrics in implants, surgical mesh, catheter cuffs

[88, 89, 90, 91, 92, 93, 94, 95, 96, 97, 98, 99, 100, 101, 102, 103, 104, 159]

Polyvinylchloride (PVC)

Catheters, blood bags, tubing

[106, 107, 108, 109, 110, 111, 112, 114, 161, 162]

Polysulfone (PSF), polyethersulfone (PES)

Hemodialysis membranes

[116, 117, 118, 119, 120, 121, 122, 123, 124, 125, 126, 127, 128], [129, 130, 131, 132, 133, 168]

Polyacrylonitrile (PAN)

Hemodialysis membranes

[135, 136, 137, 138, 139, 140, 141, 142, 143, 169]

Polymethylmethacrylate (PMMA)

Hemodialysis membranes and blood pumps


Polyamide (nylon)

Sutures, balloons for catheters

[145, 146, 147, 177]

Polypropylene (PP)

Membranes in blood oxygenators, disposable syringes


Polyglycolide (PGA)



Polylactide (PLA), poly(L-lactide) (PLLA)

Sutures, stents


It should be noted that despite using the term blood compatible polymers for this chapter, in general, synthetic polymers and other biomaterials are not considered completely blood compatible. However, significant research has been carried out to develop polymeric materials that have improved blood compatibility and create less of a reaction in blood. An understanding of polymer-blood interactions is important to achieve this goal and minimize adverse responses.

1.1 Polymer-Blood Interactions and Testing

Similar to other types of biomaterials, polymeric biomaterials induce numerous adverse responses in contact with blood such as thrombosis, infection, as well as immune and inflammatory responses. Such responses can ultimately lead to the failure of the device. Thrombosis and the complications associated with it remain the main challenge in the context of blood contacting polymers. Figure 1 shows the sequence of events during blood-material interactions [148]. As shown, protein adsorption happens as the first event upon polymer-blood contact. This results in activation of the coagulation cascade, platelets, and the complement system. Blood coagulation involves various proteins, cells, and other components in a series of complex reactions. Thrombin is a central enzyme in coagulation and thrombosis and is responsible for converting fibrinogen to fibrin. Fibrinogen also plays an important role in platelet adhesion and activation. Thrombus formation can cause issues at the site of the device or can embolize and cause issues elsewhere in the body.
Fig. 1

The sequence of events in blood-material interactions. (Taken from Ref. [148] with permission)

Testing of polymers in contact with blood should thus investigate reactions not only at the polymer surface but also in circulating blood. Protein adsorption and platelet adhesion are often studied as initial indicators of how well a material may interact with blood. However, in vitro measurements don’t always correlate with the in vivo performance of a material and rigorous characterization and analysis is needed. ISO 10993-4 provides guidance on the selection of tests for blood interactions including thrombosis, coagulation, platelets, leukocytes, and the complement system. It is important for studies to examine multiple factors, use appropriate controls, and adhere to standardized methods. The future of this area relies on continual development to better understand the mechanisms involved in blood-material interactions along with improvement towards testing standardization.

1.2 Polymer Modification for Blood Contacting Devices

For most of the currently used blood contacting devices, thrombosis is prevented therapeutically through anticoagulation and antiplatelet therapies [149]. Bleeding associated with the use of such therapies, however, necessitates the use of other approaches such as materials modification. Hence, there has been ongoing research to modify materials to reduce thrombotic complications. The general approaches used in the modification of polymers to reach this goal include: bulk and surface modifications. Bulk modifications of copolymers used in blood contacting devices such as polyurethanes have been addressed in several studies as described in Sect. 2.1. Surface modification, however, has been of greater interest due to the fact that polymer interactions with blood mainly happen on the surface. In the context of surface modification, the following strategies have generally been employed: passivation of the polymer surfaces that would prevent protein adsorption (bioinert surfaces), incorporation of bioactive molecules that would inhibit coagulation (bioactive surfaces), or modification with bioactive molecules in combination with the protein-resistant molecule.

1.3 Bioinert Surfaces

The goal of creating bioinert (also known as nonfouling, antifouling, or protein-resistant) surfaces is to minimize blood protein adsorption and passivate the surface by using “protein resistant” molecules. The most common molecule studied to create such surfaces is polyethylene oxide (PEO), also referred to as polyethylene glycol (PEG). PEO is a linear polymer with repeating units of –CH2-CH2-O-. The protein-resistant properties of PEO are associated with hydration and steric effects (Fig. 2). Numerous research studies have looked into improving blood compatibility of polymers by incorporating PEO. Typically, the surface of the polymer is modified to introduce functional groups. These functional groups can then react with PEO. Other protein-resistant molecules studied for polymer modification include PEO-containing copolymers, albumin, poly(2-methacryloyloxyethyl phosphorylcholine) (poly(MPC)), poly(2-hydroxyethyl methacrylate) (PHEMA), poly(sulfobetaine), poly(carboxybetaine), poly(2-methoxyethylacrylate) (poly(MEA)), polyacrylic acid (PAA), poly(methacrylates), hydroxyethylmethacrylate phosphatidylcholine, poly(vinyl alcohol) (PVA), polyvinylpyrrolidone (PVP), and glycidyl methacrylate (GMA). Almost all studies in this area have indicated that despite the success of such surfaces in reducing protein adsorption, they still need further modification to inhibit thrombosis.
Fig. 2

Schematic of PEO-modified surface (Excluded volume-steric repulsion of proteins by PEO)

1.4 Bioactive (Biofunctional) Surfaces

The aim of creating bioactive surfaces is to support specific interactions between the polymer surface and blood, mainly to inhibit coagulation. The most commonly used approach for inhibition of coagulation is heparinization. Heparin is a widely used anticoagulant that binds to the protein antithrombin in blood and inhibits thrombin, along with other coagulation factors, including factor Xa. It does this through its active pentasaccharide sequence. Many commercially available blood contacting devices are surface modified with heparin. Examples of such surfaces are the Carmeda® Bioactive surface (CBAS® Heparin) by W.L. Gore & Associates, Inc., and the Trillium® Biosurface by Medtronic (Fig. 3) [150]. Hirudin is another anticoagulant that also inhibits thrombin; however, unlike heparin that acts indirectly through antithrombin, hirudin directly binds to thrombin’s active site. Other molecules that have been used to create bioactive surfaces include antithrombin, an antithrombin-heparin covalent complex, a combination of heparin with other molecules such as chitosan, and small peptides. Certain bioactive molecules such as lysine have also been incorporated to take advantage of the fibrinolytic system and to lyse the clot that is formed on the polymer surfaces. Another approach in developing bioactive surfaces has focused on mimicking the nitric oxide releasing properties of endothelial cells to prevent platelet adhesion and activation. Endothelialization of polymeric surfaces has also been explored as a way of preventing thrombosis and in order to mimic the lining of blood vessels. Finally, in the past decade, researchers have focused on modification of polymers to contain both protein resistant and bioactive molecules (Fig. 4). More details on these strategies are presented throughout this chapter.
Fig. 3

Examples of commercially available heparinized surfaces: Schematic of (a) Trillium® Biosurface and (b) CBAS® Heparin [150]

Fig. 4

Schematic of biofunctional-protein resistant surface

In the next sections, the major polymer types that are used in blood contacting devices are described along with the efforts that have been carried out to improve each of these polymers for such applications.

2 Polymers Used in Current Blood Contacting Devices

2.1 Polyurethane (PU)

Polyurethanes (PUs) are block copolymers made of alternating blocks of soft elastomeric and hard crystalline segments. Figure 5 shows the typical polyurethane synthesis procedure. The hard segment is usually made of a diisocyanate and contains a diol or diamine as the chain extender (Table 2). The soft segment is typically a high-molecular-weight long chain macrodiol such as a polyester, a polyether, a polycarbonate, or a polybutadiene (Table 2). The phase segregation between the hard and soft segment provides PUs with their exclusive physical and mechanical properties. Furthermore, the wide range of choices available for soft and hard segment make PUs ideal polymers to be used in biomedical devices [151]. Medical grade PU is available from many different manufacturers. Table 3 shows a list of these medical grade PUs under various trade names. Pellathane® and Biomer® are among the very first commercial biomedical grade PUs synthesized. These PUs are made of polytetramethylene oxide (PTMO), methylene-bis-phenyldiisocyanate (MDI), and butanediol (BD) or ethylenediamine (ED) as the chain extender, respectively. Despite their current successful performance as catheters, these PUs lack properties required for long-term blood contacting devices. To improve blood compatibility, new polydimethylsiloxane (PDMS)-based PUs, polycarbonate-based PUs, and PDMS-polycarbonate-based PUs were commercially synthesized and are currently available commercially (Table 3).
Fig. 5

PU synthesis

Table 2

Common macrodiols, diisocyanates, and chain extenders used in polyurethane biomaterials

Common macrodiols used as soft segment

Common diisocyanates

Common chain extenders

Polytetramethylene oxide (PTMO)

Polyethylene oxide (PEO)

Polypropylene oxide (PPO)

Hydroxyl-terminated polybutadiene (HTPB)

Polyhexamethylene carbonate glycol


Methylene bis(p-phenyl isocyanate) (MDI)

Methylene-bis(p-cyclohexyl isocyanate) (H12MDI)

Toluene diisocyanate (TDI)

1,5-Naphtalene diisocyanate (NDI)

Ethylene glycol

1,4-Butane diol

Ethylene diamine

Hexane diol

1,4-Butane diamine

Table 3

Some examples of commercially available polyurethanes used in blood contacting devices

Trade Name


Biomer™ (Ethicon)

Polyether urethane urea

Pellethane® TPU (Lubrizol) (Dow)

Aromatic polyether urethane

Bionate® PCU (DSM)

Polycarbonate polyurethane

BioSpan® SPU (DSM)

Segmented polyurethane

CarboSil® TSPCU (DSM)




Elasthane™ TPU (DSM)

Thermoplastic polyether-urethane

PurSil® TSPU (DSM)

Thermoplastic silicone-polyether-urethane

Tecoflex HR™ (Thermedics)

Aliphatic polyether urethane

Cardiothane-51™ (Kontron)

Polyether urethane modified with PDMS

Corathane® (Corvita)

Polycarbonate polyurethane

Rimplast™ (Petrarch Systems)

Silicone-polyurethane mix

Some of the most common applications of PU in blood contacting devices are in the development of catheters, left ventricular assist devices, and pacemaker lead insulation [152]. They have become one of the preferred polymer types as a catheter for long-term use in contact with blood. This is mainly due to lower instances of infection, relatively better hemocompatiblity and tunable mechanical properties when compared with other polymers. Furthermore, compared with many other polymers, application of PU in pacemaker lead insulation has been promising [153]. A number of PUs have also been used to make left ventricular assist devices (LVADs). When used as the chamber wall material in LVADs, PU showed desired mechanical properties compared with several other polymers [1]. Without any modification, however, PU would fail as a long-term LVAD due to associated thrombosis and infection.

There is ongoing research on further improving the blood compatibility of PUs for long-term blood contacting devices. In this section, major efforts in bulk and surface modification of PU for improved blood compatibility are described.

Many initial studies on segmented polyurethanes (SPUs) were focused on using a selection of different hard and soft segments to improve blood compatibility. In one study, PEO-containing PUs showed more thrombogenicity than the pure PTMO-based PUs. It was also found that there was a threshold concentration of PEO in the PUs (when PEO/PTMO mixture was used as the soft segment) affecting the blood contacting properties of this polymer [2]. The polydimethylsiloxane (PDMS)-based polyurethanes showed the lowest platelet adhesion when compared with PUs synthesized with either of the PEO, PTMO, hydroxyl-terminated polybutadiene or polyisobutylene as the soft segments [3]. Another study showed no significant improvement in blood compatibility when PUs were synthesized by substituting the PTMO soft segment with 5 and 15 wt% of PDMS [4]. Increasing molecular weight of various soft segments in polyurethaneureas (PUUs) impacted the blood compatibility of these polymers due to creating a more microphase-separated structure. This indicated that the hydrophilic and hydrophobic balance of PUU film surfaces could affect their blood compatibility [5].

Incorporating ionic groups into segmented PUs to mimic the structure of heparin, a known anticoagulant, has widely been investigated. Substituting urethane hydrogen with sulfonate groups in PU resulted in lower platelet adhesion and activation but higher fibrinogen adsorption [6]. Use of sulfonated chain extenders in PU synthesis has also been investigated. These studies, however, have not been conclusive in evaluating blood compatibility. In PTMO-based PU, sulfonation resulted in decreased platelet activation but higher fibrinogen adsorption whereas in PEO-based PU, sulfonation caused both enhanced fibrinogen adsorption and platelet activation [7, 8]. Another study showed increased fibrinogen adsorption, platelet adhesion, and albumin adsorption but prolonged thrombin times for PUs synthesized with sulfonated chain extenders [9, 10]. Reduced platelet adhesion was observed on PUs with carboxylated chain extenders compared with PUs with sulfonated chain extenders [11]. Protein adsorption and platelet adhesion was studied on PUs with carboxylate groups grafted into their soft segment [12]. This work showed that the protein adsorption and platelet adhesion were dependent on the graft chain length with the thiopropionic graft being the most hemocompatible followed by thiosuccinic and thioglycolic grafts. Polybutadiene-based PUs with carboxylate groups incorporated into their soft or hard segment were synthesized and were investigated for protein adsorption and platelet adhesion [13, 14]. This study showed that for these types of PUs, and in the absence of known protein-resistant chains such as PEO, ion-containing PUs did not show low protein adsorption and platelet adhesion compared with the non-ion-containing PU.

Blood-polymer interactions mainly take place on the surface. Hence, many studies have been focused on improving the interactions on surfaces while preserving bulk properties of the PU biomaterials. As mentioned previously (in Sect. 1), one of the main strategies in this regard, is to create “bioinert” or “nonfouling” surfaces by the attachment of hydrophilic or other protein-resistant molecules to the surface of PUs.

In one study, PU surfaces modified with PEG and PEG-SO3 were compared with unmodified PU in terms of fibrinogen and albumin adsorption from plasma [15, 16]. PU grafted with PEG only (no sulfonate groups) was found to have the least fibrinogen and albumin adsorption compared to the other surfaces. In another study, PUU was modified with hydroxyl- and amino-terminated PEGs of different chain lengths [17, 18]. A decrease in protein adsorption was observed with an increase in PEG chain length. Amino-terminated PEG surfaces showed lower protein adsorption compared with hydroxyl-terminated PEG. Although PEG modification reduced protein adsorption, the reduction was independent of protein size or charge. When PU substrate containing monobenzyloxy PEG was modified with PEG and monomethoxy PEG as fillers, higher protein adsorption was observed [19]. This study concluded that backfilling with PEG caused the benzyloxy-containing PEG chains to stretch, which resulted in higher protein adsorption.

Use of other protein-resistant molecules such as poly(oligo(ethylene glycol) methacrylate) (poly(OEGMA)) has also been explored. Attachment of poly(OEGMA) on PU through atom transfer radical polymerization (ATRP) significantly reduced fibrinogen adsorption [20]. Similarly, zwitterionic compounds such as sulfobetaine, carboxybetaine, and phosphorylcholine have been shown to reduce protein adsorption and to improve antithrombogenic properties of PU [21]. PU surfaces modified with poly(2-methacryloyloxyethyl phosphorylcholine) (poly(MPC)) using ATRP showed significantly lower fibrinogen and lysozyme adsorption compared with nonmodified surfaces [22]. Another study showed reduced protein adsorption and platelet adhesion on poly(carbonate urethane)s containing fluorinated alkyl phosphatidylcholine side groups [23]. Yuan et al. reported decreased platelet adhesion on PU modified with sulfobetaine [24, 25, 26]. In an additional study, they used a three-step functionalization reaction to introduce carboxybetaine to PU surfaces and showed low platelet adhesion on the modified surfaces [27].

Approaches to modify the surface of PUs by attaching functional or “bioactive” molecules have also been frequently studied as a means to improve blood compatibility. This strategy is often combined with the approach to provide a bioinert layer, for example, with PEO grafting on the PU surface. These bioinert molecules can then provide functional groups to attach bioactive molecules of interest, as will be described in the studies detailed below.

Immobilization of anticoagulants on the surface of PU is frequently used to prevent thrombus formation. As discussed previously, the anticoagulant heparin has been used as a surface modifier on various polymeric materials due to its ability to catalyze thrombin inhibition. Heparinization of PU has been extensively studied and there have been various different approaches taken to create antithrombogenic heparin materials on a range of PUs [28, 29]. Immobilization of different types of heparin on PU has also been explored. Low-molecular-weight- (LMWH) and unfractionated heparin (UFH) were compared for their thrombogenicity in terms of protein and platelet interactions on PU [30]. Both LMWH and UFH modified PU surfaces demonstrated less protein adsorption and platelet adhesion than the PU controls, but variations were observed with molecular weight, particularly at shorter incubation times. An important factor in attaching bioactive molecules to polymer materials is maintaining functionality of the molecule of interest, and this can be especially important for heparin due to its need to bind antithrombin through its active pentasaccharide sequence. Liu et al. used spacer chains of different lengths and terminal functional groups to immobilize heparin to a PUU [31]. Thrombus formation on the heparinized PU was less than on controls and the adhesion and activation of platelets was suppressed.

Molecules such as PEO can also be used as spacer molecules for attachment of bioactive moieties in addition to providing a passivating layer. Various studies have investigated the combination of both PEO and heparin modification on PU [32, 33, 34]. Park et al. used PEO of different chain lengths to attach heparin onto a segmented PUU (Biomer) [35]. Protein adsorption and platelet adhesion were reduced and heparin bioactivity was enhanced with the spacers. Furthermore, a prolongation of occlusion times with immobilized heparin demonstrated the ability to create a thrombo-resistant material.

As an alternative to heparin, a covalent antithrombin-heparin (ATH) complex was developed and various studies have immobilized it on PU substrates to provide active anticoagulant function. Attachment has been through direct covalent grafting [36] and also through a base coat applied to PU catheters and grafts with PEO [37, 38, 39]. Additional studies have investigated the chemical and physical properties of ATH-modified PU [40] and have also explored the grafting of PEO with a range of MWs to elucidate the mechanisms of protein and platelet interactions [41, 42]. Overall, ATH-modified PU has demonstrated the ability to provide both noncatalytic and catalytic thrombin inhibition and shows potential as an anticoagulant surface in a range of blood contacting applications.

Surface modification of polyurethane with the direct thrombin inhibitor hirudin has also been explored. Recombinant hirudin was covalently immobilized on a poly(carbonate) urethane surface and a vascular graft using an albumin basecoat [43, 44]. In both studies, the PU modified with hirudin showed greater thrombin inhibition and binding than controls and was able to maintain potent antithrombin activity.

Promoting fibrinolysis on PU surfaces has been investigated as a method to lyse clots and decrease thrombosis [45, 46, 47]. Chen et al. created dual functioning PU surfaces by combining PEG grafting, to provide resistance to nonspecific protein adsorption, and immobilizing lysine, to promote plasminogen binding. Fibrinogen adsorption was significantly reduced on the modified PU and plasminogen bound to the surfaces with high selectivity [48]. Although the surfaces also bound tissue plasminogen activator (t-PA) from plasma and were able to lyse clots, a greater concentration of t-PA may be required. Wu et al. continued this work and loaded PU with t-PA as an exogenous plasminogen activator source. The rapid release of t-PA into plasma was confirmed and formed blood clots were dissolved [49]. These methods demonstrate significant potential for reducing thrombus formation by clot lysis on PU materials in contact with blood.

Another strategy that involves modification of PU to provide an active function is nitric oxide release and generation for inhibiting platelet adhesion and aggregation. There have been many different chemistries studied to incorporate these materials in PUs for a wide range of applications, and they have particular advantages in contact with blood for catheters and extracorporeal circuits. This area has been discussed in detail in a recent book chapter [154].

2.2 Silicones

Silicones consist of a chemical backbone with alternating silicon-oxygen bonds and organic groups attached to each silicon. The most common silicone elastomer is polydimethylsiloxane (PDMS), and it has significant use in many areas including the biomedical field (Fig. 6). Common applications where silicones are used in contact with blood include catheters, blood pumps, heart valve structures, pacemaker leads, devices for extracorporeal circulation (ECC), including blood oxygenators and tubing, as well as microfluidic devices. PDMS has also been used as a primary reference material in hemocompatibility, biocompatibility, inflammatory response, and in vivo studies [50].
Fig. 6

Silicone structure

Some of the main advantages of silicones are their thermal and chemical stability, low toxicity, oxygen permeability, ease of processing, and low cost. However, as biomaterials, in particular for blood contacting applications, their high degree of hydrophobicity is a disadvantage, since it promotes interactions at the blood-material interface including protein adsorption, platelet adhesion, red blood cell lysis, and other potentially adverse responses. To reduce the risk of thrombosis, systemic anticoagulation is typically needed with the use of silicone devices.

In order to improve their compatibility with blood, silicones can be modified using various strategies to alter their surface properties. These modification methods can be classified as surface (chemical and physical), bulk (blending, copolymerization, IPNs, functionalization), and other methods as reviewed by Abbasi et al. [51]. Since silicones do not naturally possess suitable functional groups for subsequent attachment of molecules of interest, in many cases they must first be activated. Some techniques that have been used to functionalize and modify silicone surfaces will be described in further detail below.

PDMS rubber has been modified by an irradiation and grafting method to create materials with a range of wettabilities for comparing relative blood compatibilities [52]. PDMS surfaces were treated with a CO2-pulsed laser, creating superhydrophobic surfaces, and also grafted with hydroxyethylmethacrylate phosphatidylcholine, resulting in superhydrophilic surfaces. The two extreme hydrophobic and hydrophilic surfaces demonstrated a reduction in platelet adhesion and activation compared to the control PDMS.

Zwitterionic polymers, including sulfobetaine and carboxybetaine, can provide nonfouling characteristics to PDMS and have been investigated for their ability to reduce nonspecific protein adsorption. PDMS was functionalized with carboxybetaine and blended with PDMS elastomer to create films. These materials demonstrated reduced protein adsorption (bovine serum albumin), bacterial adhesion and show promise as antifouling surfaces in various applications, including blood contacting devices [53]. Carboxybetaine groups were also covalently grafted on the surface of Si-H functionalized PDMS through a hydrosilylation reaction. These modified materials displayed a reduction in bovine serum albumin adsorption, a prolonged dynamic clotting time, and were able to decrease bacterial adhesion, demonstrating their antifouling properties [54].

As with many other polymers, modification with polyethylene oxide (PEO) has been studied extensively to provide greater hydrophilicity on PDMS. Chen et al. have completed numerous studies where silicones have been modified with different PEO polymers. PDMS-based elastomers were prepared by incorporating monofunctional PEO during curing with classic room temperature vulcanization chemistry [55]. PDMS with PEO was able to decrease fibrinogen adsorption from both buffer and plasma. Further protein adsorption studies were completed on PDMS incorporated with both mono- and bifunctional PEO using a similar rubber formation process [56]. Again, reductions in protein adsorption (fibrinogen, albumin, and lysozyme) were observed with the addition of PEO to PDMS. Compared to the controls and the bifunctional PEO surfaces, the monofunctional modification showed greater protein selectivity.

PEO of various molecular weights and functionality have also been grafted to PDMS after introducing a high density of Si-H groups on the surface by acid-catalyzed equilibration in the presence of methylhydrosiloxane (MeHSiO)n and linking PEO by a platinum-catalyzed hydrosilization reaction [57]. These surfaces demonstrated protein-resistant properties due to their ability to reduce fibrinogen adsorption from buffer and plasma, and albumin from buffer. Reductions of greater than 90% fibrinogen adsorption were achieved on the modified PDMS versus the controls. This modification method was extended to a generic process where after functionalization of PDMS with Si-H groups, hydrosilylation with PEO-NSC then provides activated esters that various biomolecules containing amines can readily react, including proteins, peptides, and glycosaminoglycans such as heparin [58]. For blood contacting applications, heparin was successively grafted to the PEO functionalized PDMS and covalently bound antithrombin providing effective thromboresistant properties.

A unique antithrombin-heparin (ATH) covalent complex was immobilized on PDMS using PEO as a linker, providing an anticoagulant surface in contact with blood [59]. Using thromboelastography ATH-modified PDMS was shown to delay the initiation of coagulation and clotting time. Polydopamine was also used as a bonding agent in additional studies investigating ATH-modified PDMS in the design of microfluidic blood oxygenators [60].

Surfaces with both resistance to nonspecific protein adsorption and fibrinolytic properties have been obtained with modification of PDMS. A PEG spacer was used to attach ε-lysine for selective adsorption of plasminogen from plasma. The modified surfaces significantly reduced the adsorption of fibrinogen, were able to bind plasminogen, and through conversion to plasmin effectively dissolved fibrin clots [61].

Platelet adhesion and activation can be naturally inhibited by nitric oxide released from endothelial cells. As with other polymers, this idea has been applied to silicones to create nitric oxide releasing surfaces. Zhang et al. used a three-step process to synthesize nitric oxide releasing silicone rubbers with covalently linked diazeniumdiolate groups. The materials were coated on the inner walls of silicone rubber tubing for extracorporeal circulation (ECC) in a rabbit model. The nitric oxide releasing materials showed a decrease in platelet consumption and platelet activation and overall had improved thromboresistance [62]. In a similar rabbit model, a solvent swelling method based on S-nitroso-N-acetylpenicillamine (SNAP) was used to load silicone rubber tubing with nitric oxide [155]. The SNAP-loaded silicone rubber ECC circuits achieved efficient release of nitric oxide, preserving the blood platelet count at 64% of baseline and providing a 67% reduction in thrombus formation compared to controls. The results suggest that these modifications have the potential to improve the hemocompatibility of silicones used in blood contacting devices.

2.3 Polytetrafluoroethylene (PTFE; ePTFE)/Teflon

Polytetrafluorethylene (PTFE), also known as Teflon, is a thermally and chemically stable, hydrophobic polymer (Fig. 7).
Fig. 7

PTFE structure

The main application of this polymer in blood contacting devices has been for catheters and vascular grafts. PTFE was developed by DuPont in 1938 and was marketed under the trade name Teflon. Later on, W.L Gore & Associates Inc. introduced expanded PTFE (ePTFE). ePTFE is more porous than regular PTFE and is used as prosthetic vascular conduits.

In vascular surgery, use of autologous saphenous vein as the graft is desirable. When this choice is not available for reasons such as insufficient size or length, prosthetic grafts are used. Along with polyethylene terephthalate (PET)/Dacron, ePTFE is the most common polymer used in this application. ePTFE has shown better performance compared with Dacron when used in certain locations in the body [156, 157]. Examples of commercially available ePTFE-based grafts include the following: Gore-Tex (W.L. Gore & Associates, Inc.); VascuGraft (B. Braun Medical Inc.); Exxcel Soft Vascular Graft (Boston Scientific); Impra CenterFlex (Bard Peripheral Vascular); Exxcel Soft (Atrium Medical Corporation), which are all unmodified; Gore Propaten (W.L. Gore & Associates, Inc.), which is heparin modified with Carmeda Bioactive Surface (CBAS); Impra Carboflo, Venaflo II, Distaflo, Dynaflo (Bard Peripheral Vascular), which are all carbon coated, Advanta VXT, Advanta SST, Advanta Supersoft (Atrium Medical Corporation), which are all reinforced; Lifespan (Angiotech/Edwards Life Sciences), Rapidax (Vascutek Ltd.), Advanta SST (Atrium Medical Corporation), which are all trilaminar; and Maxiflo, SealPTFE, Taperflo, CannulaGraft (Vascutek Ltd.), which are all unsealed and gelatin sealed [63]. Most of these are either approved or in clinical trials for peripheral bypass and hemodialysis applications.

Despite the successful performance of ePTFE when used as a large diameter vascular graft, it is prone to failure as a small diameter graft (DI < 6 mm). This failure happens due to thrombus formation as well as mismatch between the graft and the native vessel. The mismatch causes the development of neointimal hyperplasia. Efforts to improve the performance of ePTFE grafts range from antithrombotic surface modifications and coatings to endothelial cell seeding and nitric oxide modification.

Surface coating of ePTFE with polymers, proteins, and other molecules known to improve blood compatibility has been studied by a number of researchers. Yang et al. modified ePTFE grafts with poly(1,8-octanediol citrate), a biodegradable polymer, and studied the biocompatibility of these grafts in vitro [64]. They observed delayed plasma clotting as well as enhanced attachment of porcine endothelium-like cells on modified grafts compared with nonmodified ones. In a study by Sato et al. [65], ePTFE films were coated with poly(2-methoxyethyl acrylate) (PMEA) and seeded with human umbilical vein endothelial cells or aorta smooth muscle cells. Both of these cells adhered and proliferated well on these films while platelets did not adhere, showing promise for applications involving endothelialization and the development of blood vessels. In another study, Karrer et al. modified ePTFE with polypropylene sulfide-PEG and showed decreased thrombogenicity on modified grafts in a porcine shunt model [66]. Photopolymerization of an acrylate phospholipid on ePTFE grafts showed low platelet adhesion and fibrinogen adsorption in a baboon shunt model [67]. Jin et al. functionalized ePTFE capillaries with poly(vinyl alcohol) (PVA)–glycidyl methacrylate (GMA) copolymers. This was followed by immobilization of human serum albumin (HSA) onto the functionalized surfaces. Their data showed a sharp suppression of platelet adhesion on PVA coated and PVA–GMA–HSA coated PTFE capillaries [68]. The protein, P15, a cell-binding peptide, has also been attached to ePTFE grafts. These grafts exhibited a higher amount of endothelialization in a sheep model [69]. In another study, Rotman et al. coated ePTFE grafts with anti-CD34 antibodies, a protein that captures circulating endothelial progenitor cells and showed improved endothelialization on these grafts after 72 h in vivo [70]. Among other molecules used for ePTFE graft modification is carbon. Kapfer et al. compared carbon impregnated ePTFE vascular grafts with standard unmodified grafts in patients but observed no statistical difference in patency [71].

Several studies have used bioactive reagents with antithrombotic properties to modify ePTFE grafts. The most widely studied and commercially available reagent in this regard is heparin. One of the most known commercially available ePTFE grafts modified with heparin is the Gore-Tex Propaten graft by W.L. Gore. The modification of this graft is done by the covalent end-point attachment of heparin (Carmeda BioActive Surface, CBAS). A few studies have looked at these heparin-modified ePTFE grafts in a canine model. Significant reductions in platelet deposition and thrombus formation were observed on heparin-coated grafts [72, 73]. In a clinical study carried out on 86 patients, Carmeda Bioactive modified ePTFE grafts showed an improved outcome compared with unmodified ePTFE grafts [74]. A more recent clinical study with 569 patients concluded that the overall risk of primary graft failure was significantly reduced (by 37%) on these heparin-modified ePTFE grafts [75].

Modifying ePTFE with heparin through other methods has also been studied. ePTFE tubes were modified with aminated poly(1,8-octanediol-co-citrate) (POC) via POC carboxyl functional groups [76]. Heparin was then attached to the amino terminal of the POC. POC-heparin-coated ePTFE grafts showed significant reduction in platelet adhesion and performed better in whole blood clotting kinetic studies. Lu et al. coated ePTFE grafts with an anti-CD133 antibody-functionalized heparin/collagen multilayer [77]. Prolonged coagulation time and less platelet activation and aggregation was observed on modified ePTFE grafts compared with the bare ePTFE grafts. Furthermore, these modified surfaces showed a higher rate of endothelialization in a porcine carotid artery transplantation model. Zhu et al. studied ePTFE grafts modified with a chitosan/heparin complex [78]. These grafts exhibited reduced platelet adhesion in vitro and no clot formation 2 weeks postimplantation in a canine vein model. Greisler et al. used a combination of fibroblast growth factor-1/fibrin glue with heparin to improve ePTFE graft performance [79]. They observed greater endothelialization on these grafts compared with control grafts in a canine model.

Other antithrombotic molecules explored for ePTFE modification are hirudin and tissue plasminogen activator (t-PA). Heise et al. used hirudin in combination with iloprost (a vasodilator and an inhibitor of platelet aggregation) along with PEG to modify ePTFE [80]. They demonstrated improved blood flow in modified grafts and less restenosis compared with unmodified ePTFE. Greco et al. modified ePTFE grafts with a tissue plasminogen activator-iloprost combination and observed improved patency in a rat model [81].

Another approach to improve the blood compatibility of ePTFE grafts has been surface modification by seeding endothelial cells. Endothelial cells (ECs) are a natural lining of the blood vessels and secrete substances that prevent thrombosis and neointimal hyperplasia. In most studies utilizing this approach, ECs were isolated from the artery or vein of a patient, expanded in cell culture, and seeded onto the graft. In a clinical study by Deutsch et al., the EC-seeded ePTFE grafts showed significantly better patency than a control group after 9 years [82]. Magometschnigg el al. investigated the EC-seeded grafts for crural reconstruction. These grafts exhibited 39% improvement in patency after 30 days compared with the control nonmodified ePTFE grafts [83]. Laube et al. used a similar technique by seeding venous endothelial cells onto ePTFE grafts for coronary arteries. Their results demonstrated that 91% of the grafts used (out of 21 grafts) were patent after 28 months [84]. Griese et al. coated ePTFE with fibronectin and modified it further with endothelial progenitor cells (EPCs). Rapid endothelialization and reduced neointima deposition on these grafts were observed in a rabbit model [85].

Endothelial cells in the body are constantly producing nitric oxide (NO), which is essential for vascular regulation. Furthermore, NO prevents platelet aggregation, a necessary requirement for blood compatibility. Thus, modification of vascular grafts with NO to improve graft patency has been investigated [86, 158]. In a study by Pulfer et al., pores of ePTFE grafts were modified with polymeric diazeniumdiolate polyethyleneamine/NO microspheres. These grafts exhibited NO release in vitro for more than 150 h [158]. There is potential for this technique to produce grafts with improved resistance to thrombosis.

2.4 Polyethylene Terephthalate (PET)/Dacron

Polyethylene terephthalate (PET) is one of the most common polymers of the polyester family (Fig. 8). It is frequently referred to by its trade name Dacron, which is PET manufactured into fibers that can be in knitted or woven form. When PET is in film form, it is often called Mylar, also a DuPont trade name. The aromatic rings in the backbone of PET and its stiff polymer chains provide it with numerous advantages including a high melting point, excellent tensile strength, toughness, fatigue resistance, and chemical stability. When in fiber form, PET also has good crease and abrasion resistance [87].
Fig. 8

PET structure

PET has been used for many years in blood contacting medical device applications, most commonly for vascular graft prostheses. It has also been used for sewing rings on artificial heart valves, fabrics for fixation of implants, surgical mesh, and as a catheter cuff material. As an artificial blood vessel, PET has been successful for large and medium diameter sites with high flow, however, similar to PTFE, in small diameter graft applications, such as coronary artery bypass, there are major issues. Acute graft failure can occur from surface-induced thrombogenic events and delayed graft failure results from the development of intimal hyperplasia along with thrombosis.

In order to overcome these problems, a range of strategies to modify PET have been explored including physical adsorption through coatings and chemical coupling to the surface. Since PET does not possess any chemically active functionality, appropriate functional groups must first be introduced on the surface to achieve covalent grafting. Molecules that have been immobilized on PET include those that provide a passivating effect, such as PEO, and those with a particular active function, such as anticoagulants.

Studies have investigated a range of methods of immobilizing PEO on PET. Desai and Hubbell used a solution technique to incorporate PEO molecules of various molecular weights (MWs) into several biomedical polymers including PET. PEO-modified PET of MW 18,500 showed a substantial decrease in albumin adsorption compared to the PET control. Platelet adhesion was reduced to a large extent with this MW PEO and to a less extent with PEO of lower and higher MWs [88]. These authors also covalently grafted PEO to the surface of PET and investigated biological responses [89]. High MWs of PEO, specifically 18,500, again demonstrated decreases in albumin adsorption, along with fibrinogen adsorption, and platelet adhesion compared to the control. Gombotz et al. also covalently grafted various MW PEO molecules to better understand the mechanisms of low protein adsorption. In this study, it was found that a lower surface density of high MW PEO was more effective in providing decreases in protein adsorption than a higher density of low MW PEO [90].

Other methods of immobilizing PEO containing molecules on the surface of PET have been explored including modification with block copolymers and polymer brushes. A two-step procedure was used to graft two types of PEO triblock copolymers to PET, first PEO-polybutadiene-PEO (PEO-PB-PEO), and then the commercially available Pluronic PEO-poly(propylene oxide)-PEO (PEO-PPO-PEO) [91]. The surface was primed with PEO-PB-PEO to create a double bond rich layer and PEO-PPO-PEO was then coated on the surface followed by γ-irradiation. PEO-grafted surfaces showed a decrease in fibrinogen of more than 90% compared to control PET along with a reduction in platelet adhesion. PEO-containing polymer brushes have also been grafted onto PET. Li et al. utilized surface initiated atom transfer radical polymerization (SI-ATRP) to tether various chain lengths of poly(ethylene glycol) methacrylate (PEGMA) to the surface of PET. The relationship between polymer chain length and resistance to protein adsorption was assessed and a reduction in fibrinogen adsorption was achieved on PET surfaces grafted with PEGMA [92].

In early years, collagen and gelatin were used to coat PET grafts. However, these grafts have continued to cause issues due to their high thrombogenicity. Albumin coatings on PET grafts have also been explored for many years. The use of albumin as a surface modifying agent has been investigated due to its ability to provide a relatively inert passivating layer. An in vitro study of an albumin coating on Dacron arterial prostheses was carried out to assess thrombogenicity in terms of platelet activation, fibrin formation, and leucocyte interactions [93]. The albumin-coated surfaces prevented platelet accumulation on the surface, diminished coagulation activation and fibrinopeptide A formation, and had minimal leucocyte deposition on the surface. There have been several albumin-coated PET grafts used clinically including the Bard albumin-coated DeBakey Vascular® II and the albumin saturated, autoclaved DeBakey Soft Woven graft [94]. Marois et al. performed in vivo studies to compare the Bard albumin-coated vascular graft (ACG) to the uncoated Vascular® II. With the albumin-coated graft, there was no increased presence of an immune reaction and the healing response was equivalent to the uncoated prosthesis after albumin was resorbed. There was a decrease in platelet and fibrin uptake on the luminal surface of albumin-coated grafts; however, the differences were small [95].

The widely used anticoagulant, heparin has been attached to PET to provide it with bioactive functionality. To immobilize heparin, Kim et al. exposed PET to oxygen plasma glow discharge and the resulting peroxides on the surface were used to graft acrylic acid. PEO was then coupled to the surface to act as a spacer and for subsequent attachment of heparin and insulin [96]. On the surfaces that contained heparin, the plasma recalcification time and activated partial thromboplastin time were prolonged and a decrease in platelet adhesion was observed. Various heparin-modified grafts have been commercially available, including the InterVascular Hemaguard grafts and InterGard Knitted Heparin-Bonded Vascular Prosthesis by InterVascular [63]. Clinical studies with the Intervascular Hemaguard heparin-bonded collagen-coated grafts have been reported [97]. For infrainguinal bypass, these grafts produced comparable primary patency rates to other synthetic grafts. When used for femoropoliteal bypass grafts, collagen-coated heparin-bonded Dacron from InterVascular were found to have significantly better patency rates than PTFE at 3 years, but no difference at 5 years [98].

Another anticoagulant that has been immobilized on PET is the highly specific direct thrombin inhibitor, hirudin. Phaneuf et al. modified the surface of PET fibers using alkaline hydrolysis to create carboxyl groups and attach a bovine serum albumin (BSA) base coat. Recombinant hirudin (rHir) was then covalently linked to the modified surface and thrombin inhibition was assessed [43]. The PET with immobilized rHir had significantly greater thrombin adhesion and high thrombin inhibition compared to controls. Studies were continued and these r-Hir-BSA-PET vascular grafts were evaluated to determine the stability of surface bound rHir and the interaction of thrombin-rHir in an in vitro flow model [99]. The grafts coated with rHir demonstrated their stability and reduced local thrombin concentration under physiological flow. A further in vivo assessment of PET with covalently linked rHir was carried out with patches implanted in a canine thoracic aorta under high flow and shear rates [100]. After exposure to nonheparinized arterial blood flow for 2 h, the Dacron-BSA-r-Hir explanted patches demonstrated retained antithrombin activity, and compared to controls, had a much thinner pseudoinima of proteins and platelets along with significantly less thrombus formation on the graft surface. These studies have shown potential for covalent linkage of biologically active proteins like hirudin to the surface of PET vascular grafts.

Another study coated a polydopamine mussel adhesive inspired layer on PET films to then covalently graft zwitterionic cysteine [101]. Following surface characterization of the materials, protein adsorption, platelet adhesion, and hemolytic testing were carried out. The cysteine-modified PET demonstrated a significant decrease in BSA adsorption, reduced platelet adhesion and aggregation, along with a low hemolysis rate.

Various studies have utilized nitric oxide’s ability to inhibit platelet adhesion through particular surface modifications on PET. Duan and Lewis covalently grafted L-cysteine to PET to promote NO transfer to the polymer and release from it. In this study, platelet adhesion on the cysteine-modified PET was reduced by more than 50% compared to the control [102]. In a subsequent study, various thiol-containing groups, including those with L-cysteine moieties and more L-cysteine sites, were attached to PET and compared for their ability to promote NO transfer and release [103]. All polymers demonstrated a significant reduction in platelet adhesion with L-cysteine having the greatest decrease.

L-arginine immobilization has also been investigated for its potential ability to release NO when in contact with blood leading to a reduction in thrombus formation. Studies have covalently attached L-arginine to the surface of PET using glutaraldehyde as a crosslinker [104, 159]. In initial work, successful surface modification of PET with L-arginine was confirmed and the modified films demonstrated a slight reduction in protein adsorption and significant decreases in platelet adhesion and thrombus formation [159]. Further investigations with L-arginine immobilized on PET looked at various properties of anticoagulant activity and hemolysis [104]. A blood clotting test showed good antithrombogenic activity for modified PET and in clotting time assays the plasma recalcification time (PRT) and activated partial thromboplastin time (APTT) were both prolonged compared to unmodified PET. The hemolysis ratio for L-arginine modified PET was found to be low indicating that when the surfaces contact blood, red blood cells remain intact. Overall, these modifications on PET show promise for blood contacting medical device applications.

2.5 Polyvinylchloride (PVC)

Polyvinyl chloride (PVC), also known as polychlorethene, is a thermoplastic polymer made by addition polymerization of vinyl chloride monomer (Fig. 9). Additives, known as plasticizers, are typically added in biomedical applications to make this hard polymer soft and flexible. The ease of sterilization as well as the flexibility of the plasticized PVC has made this polymer a good candidate for blood contacting devices, in applications such as catheters and blood bags. The plasticizers, however, can leach out of the polymer in vivo, which leads to adverse effects such as inflammation.
Fig. 9

PVC structure

Examples of commercially available PVC blood bags include the following: Transfufol3226 (Cerus Corporation); PL146 and PL 1813 (Fenwal Inc.); Compoflex® (Fresenius Kabi); CPD bag (Grifols); A (MacoPharma); Standard PVC (Pall Corporation); XT-150 and bags with different volumes (TerumoBCT) which are all made of PVC with di(2-ethylhexyl)phthalate (DEHP) as the plasticizer; PL2209 (Fenwal Inc.); Compoflex® (Fresenius Kabi); CPP (Haemonetics); B (MacoPharma), which is made of PVC with butyryl-tri-n-hexyl-citrate (BTHC) as the plasticizer; B (MacoPharma), which is made of cyclohexane 1,2-dicarboxylate (DINCH); PL1240 (Fenwell Inc.); and Compoflex® (Fresenius Kabi), Transfer bag (Grifols), A (MacoPharma), ELX (Pall Corporation), which are all made of PVC with tris-octyl tri-mellitate (TOTM) as the plasticizer [160]. PVC bags can be used to store blood components such as red blood cells, platelets, and plasma. The plasticizer used in such applications has typically been di(2-ethylhexyl)phthalate (DEHP). Reduced red cell haemolysis has been observed when blood bags were made of DEHP-plasticized PVC compared to storage in glass. Exposure to DEHP has also shown to cause a range of adverse effects such as toxicity in animal models. In 2002, the FDA issued a notification regarding the exposure to DEHP-plasticized PVC. In recent studies, DEHP is labeled as a category 2 toxic compound, with recommendations on finding an alternative additive to replace this plasticizer in certain applications [105].

In order to minimize or eliminate the problems associated with DEHP-plasticized PVC, two main approaches have been explored. The first approach focuses on modification of DEHP-plasticized PVC to improve blood compatibility of this polymer and to reduce DEHP migration and release. Surface modification of DEHP-plasticized PVC is one of the strategies employed by researchers to reduce DEHP migration. One study showed that sulfonated DEHP-plasticized PVC has reduced DEHP migration compared to a nonsulfonated version of this polymer [106]. In contact with blood, the sulfonated polymer also exhibited decreased inflammatory response in vitro and in rodent models. Another study looked into improving blood compatibility of DEHP-plasticized PVC via surface heparinization [107]. This modification technique caused decreases in fibrinogen and Factor XII adsorption, and improvements in thrombin-antithrombin (TAT) complex and complement component C3a generation. Furthermore, lower content of DEHP at the surface of the modified polymer was reported. Gamma radiation was explored as another technique to modify DEHP-plasticized PVC surfaces [108]. Higher doses of gamma radiation were shown to minimize the amounts of leached DEHP. Bulk modification of DEHP-plasticized PVC has also been studied. Yu et al. attempted to stabilize DEHP by incorporating 2,3,6-per-O-benzoyl-β-cyclodextrin (Bz-β-CD) [109]. This modification technique had no negative impact on physical properties of DEHP-plasticized PVC but decreased the leaching of DEHP.

The second approach to eliminate problems associated with leaching of DEHP suggests the use of alternative plasticizers that can be safe in blood contacting applications. Di -(2-ethylhexyl)-adipate (DEHA), acetyl-tri-n-butyryl-citrate (ATBC), di-iso-nonyl phthalate (DINP), glycerides, castor-oil-mono-hydrogenated acetates (COMGHA), cyclohexane 1,2-dicarboxylate (DINCH), di(2-ethylhexyl) terephthalate (DEHT), tri-2-ethylhexyl trimellitate (TETM), tris-octyl tri-mellitate (TOTM), and butyryl-tri-n-hexyl-citrate (BTHC) are some of the main alternative plasticizers under investigation for this purpose. Information related to these alternative plasticizers is available in several review papers [110, 111, 112, 161]. Based on these reviews, studies on alternative plasticizers have tried to evaluate the alternatives based on the following criteria: how the plasticizer affects the stored blood component (most of the plasticizers have been shown to impact the storage of red blood cells); how toxic the plasticizer is; and how much of it is going to leach over time. There are, however, limited data on toxicity of DEHP and the alternative plasticizers. These reviews suggest that there is still a need for systematic toxicological studies of the alternative plasticizers in PVC. They recommend continued investigation on the development of DEHP-free PVC as well as alternative polymers that can be used in lieu of PVC in blood bag applications, especially red blood cell containers.

Another application of PVC in medical devices is in catheters and tubing, mostly for short-term purposes [113, 114, 162]. Studies have shown that in contact with blood, similar to other polymers, complications such as thrombosis on the surface of noncoated PVC can occur. Incorporation of heparin or hydrophilic polymers have been tested to improve blood compatibility and have proven to be effective in reducing thrombus formation [114, 162].

2.6 Membrane Materials: Polysulfone (PSF), Polyethersulfone (PES), Polyacrylonitrile (PAN), and Polymethylmethacrylate (PMMA)

In this section, polymers that are used as membrane materials in blood contacting applications will be discussed. The focus will be on polysulfone (PSF), polyethersulfone (PES), polymethylmethacrylate (PMMA), polyacrylonitrile (PAN), and ethylene vinyl alcohol copolymer (EVA), all materials that are widely used for hemodialysis membranes.

Hemodialysis involves the use of semipermeable membranes to filter waste products from the blood due to loss of kidney function. The membranes are usually configured as hollow fibers, which have a high surface area that contacts and interacts with blood and its components. The blood compatibility of polymers used as hemodialysis membranes is therefore extremely important.

One of the initial materials used as a membrane in hemodialysis was cellulose. Cellulose membranes, however, showed complications in contact with blood, mainly by activating the complement system as well as reducing patients’ immune function leading to infection [163, 164, 165, 166]. Due to these complications, use of synthetic polymers such as PSF, PES, PAN, PMMA and EVA as the membrane in blood contacting applications were tested and provided evidence of minimizing or eliminating such complications [167].

Polysulfones (PSFs) are thermoplastics that contain aromatic groups joined by an SO2 group (Fig. 10). The multiple aromatic rings can be connected directly, as in the case of polysulfone (PSF), or can be joined by an oxygen, resulting in polyethersulfone (PES) (Fig. 10). Due to this structure, PSFs possess high mechanical strength, thermal stability, good chemical resistance and are able to be easily processed into flat sheets and hollow fibers. They have an asymmetric microstructure and are inherently hydrophobic. For medical applications, their ability to withstand a wide range of sterilization techniques including steam, ethylene oxide, and gamma radiation is a great advantage. However, their hydrophobicity makes them especially prone to inducing adverse reactions such as complement activation, blood coagulation, platelet adhesion and activation, and other cellular responses when in contact with blood. Since the blood compatibility of both PSF and PES membranes is insufficient, systemic administration of anticoagulants is required during their use in hemodialysis.
Fig. 10

PSF and PES structures

Polysulfone was introduced as a commercial hemodialysis membrane by Fresenius [115]. Fresenius continues to manufacture various lines of hollow fiber membranes for hemodialysis including high performance membranes, along with both Toray and Asahi. Numerous strategies to modify PSF membranes have been investigated in order to improve their blood compatibility. Polyvinylpyrrolidone (PVP) is added during the manufacturing of PSF in order to decrease the hydrophobicity and to change pore size distribution. PSF membranes have been modified with PVP by various methods including physical blending, chemical grafting, and other surface modifications [116, 117]. PSF membranes blended with PVP were also evaluated on the nanoscale using atomic force microscopy (AFM) [118]. The study determined that the PSF/PVP surface was covered with PVP and protein adsorption greatly decreased with increasing content of PVP.

Many other methods have also focused on decreasing the hydrophobicity of PSF using hydrophilic molecules such as PEO, phophorylcholine, and other bioinert molecules. PEG was incorporated into PSF membranes through the synthesis of amphiphilic graft copolymers with PSF backbones and PEG side chains (PSF-g-PEG) [119]. These membranes were hydrophilic and exhibited high resistance to protein adsorption and cell attachment, showing potential for use in hemodialysis. In another study, PSF was coated with Pluronic™, triblock copolymers of polyethylene oxide and polypropylene oxide (PEO-PPO-PEO), with varying PEO block length [120]. Pluronic™-coated PSF membranes showed a decrease in the adsorption of plasma proteins (albumin, globulin, and fibrinogen) along with a reduction in platelet adhesion. PSF has also been blended with both PVP and PEG and these membranes were compared to a surface treatment method using trimesoyl chloride and m-phenylene diamine, and a polyacrylonitrile (PAN) unmodified membrane [121]. Although all membranes displayed similar protein adsorption, the blend and surface-modified membranes exhibited less platelet adhesion and thrombus formation indicating improved blood compatibility.

Ishihara et al. have modified PSF membranes with the phospholipid polymer 2-methacryloyloxethyl phosphorylcholine (MPC) using a blending method. Surface characterization showed that the MPC additive was concentrated at the surface and provided a reduction in fibrinogen adsorption [122]. Further studies found that with an increase in MPC composition, the density of the adsorbed total protein from plasma decreased. The MPC modified membranes were able to reduce platelet adhesion and aggregation, demonstrating their potential for improving blood compatibility [123]. Asymmetric porous membranes were also prepared with the MPC modifications and the decreases in protein adsorption and platelet adhesion were still evident [124].

Strategies to attach functional molecules to PSF membranes have also been investigated. PSF membranes have been grafted with a layer of poly(acrylic acid) to covalently bind conjugated linoleic acid (CLA), a naturally occurring substance with antioxidant properties [125]. These CLA-immobilized membranes reduced protein adsorption and hemolysis, prolonged coagulation time, and effectively improved blood compatibility of PSF. Heparin, the well-known anticoagulant, was covalently immobilized on PSF after activation of the surface with amino groups and EDC/NHS chemistry to bind heparin [126]. The modification demonstrated increased hydrophilicity of the membranes along with improved blood compatibility, in terms of longer coagulation times and reduced platelet adhesion [126]. PSF has also been modified with both chitosan and heparin. In a study by Yang and Lin, membrane surfaces were first treated with ozone to introduce peroxides and then grafted with either acrylic acid or chitosan, followed by the covalent immobilization of heparin with glutaraldehyde [127]. The hydrophilicity of PSF membranes increased with the addition of chitosan and heparin and both protein adsorption and platelet adhesion decreased, along with a prolongation of coagulation time. A recent study utilized the popular mussel-inspired method of polydopamine coatings to graft heparin and bovine serum albumin (BSA) covalently on PSF membranes through the reactive polydopamine layer [128]. After modification, the membranes showed improved hydrophilicity, reduced protein adsorption and platelet adhesion, and extension of blood coagulation times. Overall, these methods show promise for improving the blood compatibility of PSF membranes used in hemodialysis applications.

More recently, polyethersulfone has been investigated and used as a membrane material in hemodialysis due to its advantageous properties including its strength and stability. Although surface modification can be difficult due to its stable chemical structure, a range of methods have been utilized to provide better blood compatibility including bulk modification, blending, coating, and grafting methods. Like PSF and other polymers, PVP and PEG have been added to PES to increase its hydrophilicity [129, 130]. Surface modifying macromolecules (SMMs) have also been blended with PES to create surfaces with fluorinated functional groups [130]. These materials show great potential for improving the blood compatibility of PES as well as other polymers in contact with blood [132]. Many other modification strategies for PES have been studied and several review papers discuss these in further detail [133, 168].

Polyacrylonitrle (PAN) (Fig. 11) is a semicrystaline polymer with the general formula (C3H3N)n. It offers some advantages when used as a membrane material in hemodialysis systems, including good mechanical and thermal stability. Furthermore, its nitrile groups (−CN) can be easily modified to offer functional groups for surface modification. In hemodialysis, the PAN hollow fiber membrane allows small- to middle-size proteins to be removed from blood. Use of these membranes in blood-related systems is usually accompanied by the injection of anticoagulants. This is due to the hydrophobic properties of PAN membranes, which can cause biofouling, thrombosis, and immunoreactions in blood contacting applications. One of the early PAN membranes made was AN-69, a copolymer of acrylonitrile and sodium methallylsulfonate (made by Rhone-Poulenc, France). These membranes demonstrated improved outcomes in terms of their immune response when compared with cellulose membranes [134]. Despite wider use of PSF membranes, PAN membranes still demonstrate some advantages over PSF membranes. Easier surface modification can be achieved on PAN membranes to improve blood compatibility. Numerous research studies have been carried out in this area, mainly to increase hydrophilicity of the PAN membrane. The modifications range from incorporating hydrophilic polymers to immobilizing biomolecules.
Fig. 11

PAN structure

Blending PAN with 20 wt% polyvinylidine fluoride (PVDF) increased the hydrophilicity of the membrane and was shown to reduce protein adsorption, platelet adhesion, and thrombus formation [135]. In another study, membranes made of a copolymer of acrylonitrile with N-vinyl-pyrrolidone showed protein-resistant properties compared with the blended PAN with poly (N-vinyl-2-pyrrolidone) alone [136]. Both platelet adhesion and albumin adsorption decreased with an increase in N-vinyl-pyrrolidone content in the copolymer [137]. Copolymerization with maleic anhydride or 2-hydroxyethyl methaxrylate to obtain poly(acrylonitrile-co-maleic anhydride) (PANCMA) or poly(acrylonitrile-co-HEMA) (PANCHEMA), respectively has also been studied. These copolymers were specifically useful in grafting further hydrophilic polymers or biomolecules onto the surface of the membrane. Attaching polyethylene glycol (PEG) on this copolymer improved protein-resistant properties and reduced platelet adhesion and macrophage attachment [138]. Immobilization of heparin, insulin, or chitosan onto the PANCMA membrane decreased platelet and macrophage adhesion [139, 140].

Ubricht et al. modified the surface of PAN membranes using plasma treatment. They further grafted poly hydroxyl-ethylmethacrylate (PHEMA) on the plasma-treated surfaces. Both plasma-treated and poly(HEMA)-grafted surfaces showed reduced protein adsorption [141]. Ulbricht et al. also grafted polymerized hydrophilic monomers such as acrylic acid, 2-hydroxymethyl methacrylate and poly(ethylene glycol) methacrylate on PAN membranes using UV radiation graft polymerization. They observed very little bovine serum albumin adsorption on these surfaces. Another approach to modify the PAN membranes took advantage of hydrolyzing nitrile groups of PAN into carboxyl groups. These carboxyl groups can then be further used to attach biomolecules. Yang and coworkers used this strategy to attach a chitosan/heparin complex onto PAN membranes. This modification provided a reduction in protein adsorption, platelet adhesion, and thrombus [142]. The same technique was also used to graft conjugated linoic acid onto the PAN surface. These modified membranes showed increased coagulation time in vitro compared with the unmodified PAN membrane [143]. In another study, Yang and coworkers immobilized platelet-adhesion promoting collagen (COL) and platelet-inhibiting human serum albumin (HSA) onto PAN membranes [169]. HSA-modified membranes demonstrated lower platelet adhesion, longer blood coagulation time, and higher thrombin inactivity levels, as well as lower complement activation compared with the unmodified and COL-modified PAN.

Polymethylmethacrylate (PMMA) is a synthetic polymer with the structure shown in Fig. 12. Its use as a membrane in hemodialysis has been of interest due its good solute permeability and relatively good biocompatibility [170, 171]. Several studies have shown the advantages of using PMMA over other synthetic membranes in reducing a complication known as uremic pruritus. This complication happens when there is excessive urea in the blood due to renal failure. Some studies have shown that PMMA membranes could remove some of the high-molecular-weight uremic toxic solutes, which cannot be typically removed by membranes made of other synthetic polymers [172, 173]. This property was associated with the ability of the PMMA to adsorb such solutes and its improved permeability [174]. The main manufacturer of commercial PMMA membranes is Toray Medical Co. Ltd. This membrane is made by mixing isotactic (iso) and syndiotactic (syn) PMMAs. Their newly developed PMMA membrane called the BG-U series has an increased homogenized number of pores and has shown to be very effective in removing a range of solutes with low platelet adhesion, low complement activation, and high degree of cytokine removal.
Fig. 12

PMMA structure

There are other less studied synthetic polymers used as hemodialysis membranes. An example of one is ethylene vinyl alcohol copolymer (EVA). EVA is a hydrophilic polymer with protein-resistant properties. Two commercial versions of this polymer, EVAL and EVOH (by Kawasumi Laboratories, Japan) have been produced. In several research studies, EVA membranes have exhibited reduced activation of platelets and decreased formation of platelet-neutrophil aggregates compared with cellulose diacetate, PSF, and PMMA membranes in in vivo and in vitro models [175, 176].

2.7 Other Polymers

In the previous sections, polymers that have been most extensively studied for blood contacting applications were discussed. In this section, a brief overview is given of a few additional polymers that have been used in specific blood contacting applications. For these polymers there have been less research studies completed addressing the enhancement of their blood compatibility, but they still warrant discussion.

2.7.1 Polyamide

Polyamide is the general term used for any polymer containing an amide linkage. In this section, we use this term for synthetic polyamides, mainly nylons, used in blood contacting applications such as sutures. The high tensile strength of nylon makes it a suitable candidate to be used in sutures (specifically as nonresorbable sutures), and in tendon and ligament repair [144]. Another application of synthetic polyamides is in balloons for catheters, where they are typically used as a copolymer such as a polyamide-polyurethane copolymer. Research studies on the biocompatibility of polyamide sutures are generally focused on complications associated with infection. Modification strategies to render antimicrobial properties have, hence, been studied by several researchers. These modification techniques have been addressed in a review paper by Shmack et al. [145]. Several studies have also been carried out on the modification of polyamide for blood compatibility. For example, a copolymer of amidoamine and methylmethacrylate with heparin binding capacity was studied for blood compatibility [146, 147]. The heparin-bounded copolymer showed low thrombus formation (<0.22 mg in 60 min) and a low percentage of hemolysis (<5%). In another study, Nagase et al. synthesized polyamides containing phospholipid (PC) moieties. After incubation with platelet rich plasma for 1 h, the modified polyamides showed reduced platelet adhesion compared with nonmodified polyamides (ten times reduction) [177].

2.7.2 Polypropylene

Polypropylene (PP) is widely used in syringes as well as in membranes for blood oxygenators. Most disposable syringes today are made of polypropylene. The first microporous polypropylene membrane oxygenators were developed in the 1970s. These membranes can act as an artificial lung in clinical extracorporeal circuits. An example of a commercially available oxygenator with a polypropylene membrane is the Affinity NT Oxygenator by Medtronic. In most applications, these membranes are coated with anticoagulants such as heparin for improved blood compatibility. Medtronic, for example, offers two heparin coated Affinity NT membranes: the Carmeda® BioActive Surface (CBAS) and Trillium® Biosurface.

2.7.3 Polyglycolide (PGA), Polylactide (PLA), and Poly(lactide-co-glycolide) (PLG)

PGA, PLA, and PLG are all biodegradable polymers used in applications such as tissue engineering scaffolds, drug delivery systems, and biodegradable sutures. Two examples of biodegradable sutures currently used in clinical applications are Vicryl® (Ethicon Inc.) (composed of a copolymer made from 90% glycolide and 10% L-lactide) and Dexon (composed of PGA). The degradation of these sutures happens by hydrolysis and they are typically used in ophthalmic and orthopedic surgeries.

PLA and PLG have also been used for the development of bioresorbabale stents and for coating drug eluting stents. These types of stents were made in response to the limitations associated with metallic stents (restenosis, thrombosis, and impairment of vessel geometry) [178]. Lincoff et al. showed that stents coated with low-molecular-weight PLA (80 kD) elicit an intense inflammatory response as opposed to stents coated with high-molecular-weight PLA (321 KD) in a porcine coronary artery model [179]. Using a porcine animal model and high-molecular-weight PLA stents, Tamai et al. observed no thrombosis (6 months postimplantation) and no inflammation (16 weeks postimplantation) [180]. Similar promising results were observed for stents made of PLG [178].

3 Conclusion

Polymers are frequently used in devices that come in contact with blood in both short- and long-term applications. Their advantageous mechanical properties have provided improved function in many cases, but due to their foreign nature in the body, these synthetic materials continually illicit adverse responses. As a result, thrombosis remains a major issue and a limiting factor for the extended use of polymeric blood contacting devices. Some devices function sufficiently, particularly in short-term applications, for example, polyurethane and silicone peripheral catheters. However, in other applications, such as small diameter vascular grafts, current PTFE and PET materials have been unsuccessful and thrombus formation is detrimental to their use. In hemodialysis, membrane materials include PSF, PMMA, and PAN, but due to the large membrane surface area in contact with blood, adverse interfacial reactions create limitations. As discussed in this chapter, there are numerous other polymers used in blood contacting applications, but improvements to these are also still needed.

To overcome thrombosis and attempt to improve the blood compatibility of polymers, both bulk and surface modification methods have been researched. The common surface modification strategies that have been explored among polymers include the creation of bioinert surfaces, bioactive surfaces, and a combination of the two. Although bioinert modification with PEO and other molecules has provided improved results, these approaches have been inadequate when used in devices clinically. Functionalizing surfaces with bioactive molecules has shown potential but further work is required. It is expected that a combination of these strategies to best mimic the multifunctionality of native endothelium may lead to improved blood compatibility. Furthermore, continued investigations are needed to fully understand the complicated processes involved in blood-material interactions along with more standardized approaches to testing polymers for their level of blood compatibility.


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© Springer Nature Switzerland AG 2018

Authors and Affiliations

  1. 1.Wentworth Institute of TechnologyBostonUSA
  2. 2.McMaster UniversityHamiltonCanada

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