Biomimetic polymers with phosphorylcholine groups as biomaterials for medical devices

Abstract

Biomimetic molecular designs can yield superior biomaterials. Polymers with a phosphorylcholine group, a polar group of phospholipid molecules, are particularly interesting. A methacrylate monomer, 2-methacryloyloxyethyl phosphorylcholine (MPC), was developed using efficient synthetic reactions and purification techniques. This process has been applied in industrial production to supply MPC globally. Polymers with various structures can be readily synthesized using MPC and their properties have been studied. The MPC polymer surface has a highly hydrated structure in biological conditions, leading to the prevention of adsorption of proteins and lipid molecules, adhesion of cells, and inhibition of bacterial adhesion and biofilm formation. Additionally, it provides an extremely lubricious surface. MPC polymers are used in various applications and can be stably immobilized on material surfaces such as metals and ceramics and polymers such as elastomers. They are also stable under sterilization and in vivo conditions. This makes them ideal for application in the surface treatment of various medical devices, including artificial organs, implanted in humans.

1. Introduction

Thrombus formation is a serious issue for some cardiovascular medical devices such as implantable artificial hearts and artificial valves, because they come into contact with blood.^1)-3) To counter thrombus formation in patients with such devices, the administration of drugs that inhibit blood coagulation is necessary, but this imposes a burden on patients. In addition, anticoagulants are necessary for extracorporeal blood circulation treatments such as hemodialysis.^4),5)

Therefore, since the 1960s, developing antithrombogenic materials has been a major focus in biomedical, medical, and pharmaceutical sciences, leading to several key concepts being developed by the 1970s. Figure 1 illustrates the historical development of material designs for obtaining antithrombogenic polymers.⁶⁾

Fig. 1

(Color online) Interfacial concepts for biomedical materials. PTFE: polytetrafluoroethylene, PDMS: polydimethylsiloxyane, PVA: poly(vinyl alcohol), PHEMA: poly(2-hydroxyethyl methacrylate). The figure was modified based on that in Ref. 6.

The first method for obtaining an antithrombogenic surface relied on its physicochemical properties, specifically its hydrophilic and hydrophobic nature. Surface free energy affects the adsorption of proteins and other biological components.^7)-9) Researchers have explored grafting water-soluble polymers onto surfaces to increase hydrophilicity and interface mobility.^10)-13)

To precisely design the polymer surface in contact with blood, microdomains can be created using block-type copolymers with polymer segments with varying elasticities and surface free energies, or by controlling surface electric charges with polyion complexes of polyelectrolytes. A notable success in this area is that of segmented polyurethanes, used in developing diaphragm-type blood pumps and the first total artificial heart implanted in humans.^14),15) Notably, Japanese researchers significantly advanced the development and testing of antithrombogenic polymers between 1982 and 1984 through support by the Ministry of Education for the “Design of Multiphase Component Materials”.¹⁶⁾ The project involved several researchers and led to various antithrombogenic polymers, such as hydrophobic-hydrophilic segment-block polymers,¹⁷⁾ polyion complexes,^18),19) and amorphous crystalline surfaces.²⁰⁾ These polymers, measured in vitro, were found to inhibit platelet adhesion and activation on heterogeneous, neutral, or completely hydrophilic surfaces. Polyion complexes and amorphous-crystalline surfaces were later used in temporary medical devices such as cell separation filters and cannulas, respectively. Poly(2-hydroxyethyl methacrylate (HEMA)-block-styrene (St)) was used in small-diameter artificial vessels and exhibited good patency in animal models for more than three years; however, it could not be developed as a clinical implant.²¹⁾

The use of bioactive biomolecules to inhibit blood coagulation on surfaces has been explored. For instance, heparin-immobilized surfaces reliably delay blood coagulation, aiding the development of temporary medical devices such as catheters and oxygenators.²²⁾ Urokinase, an enzyme that dissolves an immobilized thrombus, is used on catheters for a short period.²³⁾

Unlike these polymer designs, liposomes, aggregates of natural phospholipid molecules, are used in the pharmaceutical field for dissolving materials for drug administration. Liposomes are garnering interest as potential blood coagulation inhibitors due to their lack of significant biological reactions, even when directly administered into the bloodstream.²⁴⁾ Hence, fundamental biochemical studies have explored phospholipid aggregates in diverse forms as cell membrane models, offering a pathway for developing novel biomimetic materials.

2. Cell membrane mimicking polymers

2.1. Monomer design concept.

Cells are vital within biological systems, functioning independently and in response to chemical and physical signaling among cell groups, often regulated by the cell membrane. Figure 2 illustrates the cell membrane, a molecular hybrid primarily consisting of phospholipids, proteins, and polysaccharides.²⁵⁾ The cell membrane provides mechanical support for cell morphology and regulates chemical concentrations in the cytoplasm. Additionally, cell-cell junctions facilitate communication between neighboring cells.^26)-28) Due to its diverse functions, the cell membrane is an appealing model for fabricating nanostructured biomimetic materials for bio-, nano-, and information technologies.

Fig. 2

(Color online) Illustration of the cell membrane structure and representative phospholipids constructed using a cell membrane with a double layer structure. PC: phosphorylcholine, PE: phosphorylethanolamine, PS: phosphorylserine.

Phospholipids in the cell membrane serve to effectively separate the intracellular cytoplasm from the extracellular environment while providing a scaffold for presenting glycoprotein receptors and membrane proteins on the cell surface.^29),30) Comprising hydrophobic alkyl chains with a hydrophilic polar head group, phospholipids spontaneously assemble into continuous bilayer membranes in aqueous media. A key phospholipid polar group on the outer surface of the cell membrane is the phosphorylcholine (PC) group, which is electrically neutral and zwitterionic (Fig. 2). In contrast, the interior of the cell membrane contains numerous electrically negative phospholipid polar groups, such as phosphatidylserine (PS) group and phosphorylethanolamine (PE) groups.³¹⁾ In biomimetic chemistry, phospholipid molecules have been employed to prepare cell membrane-like structures such as liposomes and Langmuir-Blodgett membranes. However, a significant drawback of this molecular assembly is its insufficient chemical and physical stability. Enhancing the stability of phospholipid assemblies is crucial for constructing interfaces between living and artificial systems. One strategy to address this challenge is designing a novel polymer incorporating phospholipid polar groups, which offers improved chemical and physical stability due to the bonding between the phospholipid-like moieties.^32),33) The bioinspired approach to polymer development involves selecting specific phospholipid polar groups and incorporating them into polymer chains. The PC group, a polar group for obtaining bioinert polymers, is often preferred. Additionally, conventional polymerization processes can be used to synthesize various molecular structures and architectures.

2.2. Synthetic route of methacrylate with the phosphorylcholine group.

Methacrylate bearing a PC group on its side chain, 2-methacryloyloxyethyl phosphorylcholine (MPC), is widely regarded as the most suitable monomer.³⁴⁾

Figure 3 illustrates the total synthetic route to producing MPC. Improved and elegant syntheses need to be developed to widen the scope of MPC polymer science. Even trace amounts of water markedly influenced the purity of the resulting compound because of the hygroscopic properties of MPC.³⁵⁾ The initial step in MPC preparation involves the reaction between cyclic phosphoryl chloride, 2-chloro-2-oxo-1,3,2-dioxaphospholane (COP), and HEMA to form 2-(2-oxo-1,3,2-dioxaphospholan-2-yloxy)ethyl methacrylate (OPEMA). During the reaction, triethylamine (TEA) is required to trap hydrogen chloride as triethylammonium chloride. The subsequent step involves ring-opening of the cyclic phosphate group in OPEMA using trimethylamine (TMA), leading to the formation of PC groups. Sufficient cooling of the product resulted in the crystallization of MPC from the reaction mixture. The final product is obtained as a white powder with a melting point of approximately 140℃. The chemical structure of MPC obtained was confirmed by ¹H-NMR, ¹³C-NMR, and IR spectroscopies.³⁵⁾ The results in the literature are indicated as follows: ¹H-NMR (CDCl₃): δ = 1.90 (-CH₃, 3H), 3.27-3.36 (-N(CH₃)₃, 9H), 3.70-3.80 (-CH₂N, 2H), 4.00-4.10 (POCH_2－, 2H), 4.21-4.31 (-OCH₂CH₂OP-, 4H), 5.60 (=C-H, 1H), and 6.10 (=C-H, 1H). ¹³C-NMR (CDCl₃): δ = 18.59 (-CH₃), 54.25 (-N(CH₃)₃), 59.44 (-CH₂N), 63.34 (-POCH₂-), 64.41 (-CH₂OP), 66.20 (-OCH₂-), 126.09 (=CH₂), 136.22 (=C=), and 167.26 (-CH₃). IR (cm^－1): 1710 (C=O), 1640 (C=C), 1300, 1230, 1177, 1085 (-POCH₂).

Fig. 3

Synthesis pathway for MPC in the laboratory (a) and at industrial scale (b). HEMA: 2-hydroxyethyl methacrylate, TEA: triethylamine, THF: tetrahydrofuran; DIPA: di-i-propylamine; TMA: trimethylamine; MPC: 2-methacryloyloxyethyl phosphorylcholine.

Successful MPC synthesis and purification technologies were transferred to NOF Co. Ltd., a Japanese chemical company, with support from the Japan Science and Technology Agency (JST). In 1999, a facility for large-scale industrial MPC production was established.³⁴⁾ Minor adjustments to the synthetic procedure were made to comply with the regulations regarding the utilization of certain compounds. The most effective method involved substituting TEA with diisopropylamine to remove hydrogen chloride during condensation. This markedly increased the purity of OPEMA and the synthetic yield of MPC to ＞ 80%. The industrial facility currently operates smoothly, producing several tons of MPC annually. Establishing a global distribution network for MPC as a reagent may expand its potential applications and stimulate research for developing new functional polymers. Using the synthetic conditions for MPC, PC group-bearing monomers with various polymerizable groups, including acrylic, styryl, and fumarate groups, have been synthesized. MPC derivatives with various hydrocarbon chains or oxyethylene chains between the polymerizable and PC groups have also been synthesized.³⁶⁾

Inspired by MPC polymer research, methacrylate monomers containing different phospholipid polar groups have been synthesized, and the properties of their polymers are under investigation.^37),38) For example, polymers with PS groups show activation of the immune system and exert immunostimulatory effects.³⁹⁾ Additionally, research is underway to assess the biocompatibility of polymers containing sulfobetaine or carboxybetaine groups in the side chains.^40)-43) These are betaine polymers with positive and negative charges in molecules, such as the PC groups. Some of these polymers have been applied as surface treatment materials for medical devices intended for external use, for example, blood cell separation filters.⁴⁴⁾

2.3. MPC polymer synthesis.

MPC undergoes addition polymerization initiated by radical generators such as general methacrylates (Fig. 4).^{34),35),45),46)} Its copolymerization ability allows for the introduction of several functional groups into its polymers. Additionally, living radical polymerization reactions, atom transfer radical polymerization, and reversible addition-fragmentation chain-transfer polymerization can be employed, enabling control over molecular weight, molecular weight distribution, and terminal functional groups.^47)-49) Consequently, polymers with precise overall molecular structures, including block and graft polymers, have been obtained.^50),51) Recent reports indicate the production of polymer brushes with high-density grafted polymer chains by binding a living radical polymerization initiator to a substrate surface and utilizing it for MPC polymerization.^52)-55)

Fig. 4

Synthesis pathways for representative MPC polymers and derivatives. MPC: 2-methacryloyloxyethyl phosphorylcholine.

The solubility of MPC polymers in solvents is determined by their molecular structure, composition of MPC and other monomer units, and molecular weight. For instance, a homopolymer of MPC, poly(MPC) (PMPC), can readily dissolve in water and alcohol but is insoluble in acetone, acetonitrile, tetrahydrofuran, and specific compositions of aqueous ethanol solutions (60-92 vol% ethanol). This co-nonsolvency may be attributed to variations in the solvation states of the polymer chain and solvent molecules.^56)-58)

Introducing other monomer units into MPC alters its solubility. By adjusting the monomer feed ratio, the MPC unit composition in the polymer can be easily regulated. When MPC was copolymerized with various alkyl methacrylates such as n-butyl methacrylate (BMA; forming poly(MPC-co-n-butyl methacrylate [PMB]), t-butyl methacrylate, n-hexyl methacrylate (forming poly(MPC-co-n-hexyl methacrylate [PMH]), n-dodecyl methacrylate (forming poly(MPC-co-n-dodecyl methacrylate [PMD]), or n-stearyl methacrylate (forming poly(MPC-co-n-stearyl methacrylate [PMS]), the polymerization proceeded smoothly, resulting in a statistically random sequence.⁴⁵⁾ The solubilities of the MPC polymers in the solvent depended on the MPC unit composition, hydrophobicity of the alkyl methacrylate unit, and the molecular weight of the polymer. For PMB, increasing its molecular weight to ≥ 500 kDa enhances stability in biological conditions, preventing easy detachment and resistance to enzymatic or hydrolytic degradation. It remains stable in sterilized environments, such as γ-ray, ethylene oxide gas, and autoclave sterilization, making it suitable for medical device materials. This polymer is registered in the Master Files for Devices of the US Food and Drug Administration as Lipidure^®-CM5206 (NOF).³⁴⁾

Another crucial MPC reaction involves the Michael addition of thiol or amine compounds.^59),60) The reaction with the methacrylate group occurs smoothly in mild conditions, allowing the introduction of phosphorylcholine groups into compounds with various functional groups, including silane coupling, diols, and primary amino groups.

2.4. Hydration.

MPC polymers demonstrate unique interfacial properties when they are hydrated in aqueous medium. The representative interfacial properties of PMPC and PMB are summarized in Table 1.^61)-65) The water contact angles are very small, that is, they have super-hydrophilic nature. When they are under water, contact angles of air bubble and oil droplets are very large. That is, they cannot attach to the MPC polymer surface tightly, meaning that the surface provides anti-fouling properties. The surface ζ-potential of the polymers is almost zero. These findings clearly indicate the phosphate anion and trimethyl ammonium cation in the PC group form an internal salt.

Table 1

Representative interfacial properties of MPC polymers

	PMPC(Polymer brush on Si wafer)	PMB(Coating on the substrate)
Mole fraction of MPC unit in the polymer	1.00	0.30
Water contact angle under air (deg)	＜ 561)	＜ 206)
Air bubble contact angle in water (deg)	17061)	152.2 ± 1.964)
Hexadecane contact angle in water (deg)	17561)	―
Surface ζ-potential (mV)	0.2 ± 262)	0.019)
Amount of adsorbed protein (ng/cm2)	＜ 563)	＜ 44 ± 2065)

PMPC: poly(MPC) chains were fixed on the St wafer by surface-initiated graft polymerization by ATRP procedure. PMB: high-molecular weight poly(MPC-co-BMA) was used for coating on the substrate by a solvent evaporation procedure using ethanol solution the polymer.

PMPC exhibits excellent aqueous solubility, even in high ionic concentrations.⁶⁶⁾ Surprisingly, it dissolves in a saturated NaCl aqueous solution, remaining largely unaffected by different ionic species.⁶⁷⁾ Its unique solvent composition-dependent dissolution behavior in an aqueous-ethanol system suggests a distinct hydration structure compared with traditional water-soluble polymers. Conversely, the MPC polymer of PMB with BMA units of the hydrophobic groups does not dissolve in water but forms hydrogels, with water content increasing with temperature.³⁵⁾ The water content of hydrogels made of hydrophilic/hydrophobic polymers is generally known to decrease with increasing temperature in the range up to approximately 70℃. This has been explained with respect to the hydrogen-bonding properties of the water molecules hydrating the hydrogel. In other words, the temperature dependence of hydrogen bonding directly affects hydrogen bonding hydration and the resulting hydrophobic interactions. As the temperature rises from room temperature, water molecules bound to hydrogen in the polymer chains are desorbed by the bonding forces, influencing hydrophobic interactions between the polymer chains. In MPC polymers, the PC groups cover the polymer main chain almost completely because of their bulky structures, considerably affecting hydration of the polymer, a strong reflection of the nature of PC groups.^66),68) Water hydration involves hydrogen bonding-based hydration, ionic hydration, where water molecules strongly align with ionic functional groups, and hydrophobic hydration, where water molecules preferentially interact with the hydrophobic groups through hydrogen bonding.

2.5. Water state in the hydrated polymer.

Understanding phenomena at the interface of biomaterials with aqueous media is crucial for their application in living environments. Water, with its unique intermolecular interactions, exhibits special properties despite its low molecular weight, such as high boiling point attributed to hydrogen bonding between molecules in liquids. Thermal energy is required to break these bonds. Disruption of these hydrogen bonds leads to system instability, believed to be the origin of hydrophobic interactions. Némethy and Scheraga considered the energy levels of water molecules around hydrocarbon groups and proposed that adding van der Waals forces between the hydrophobic groups in water molecule clusters where four hydrogen bonds interact helps in the system becoming energetically stabilized.^69),70) However, the energetically strong hydrogen bonds replace the van der Waals forces when there are three or fewer hydrogen bonds, thereby becoming energetically unstable. Therefore, they demonstrated a mechanism of hydrophobic interactions that helps in system stabilization by agglomerating the hydrophobic groups to reduce the contact area with the water molecules.

Hydrophilic polymers undergo dissolution or swelling when exposed to aqueous media, suggesting hydration of the polymer chains through interactions with water molecules. This alters the cluster structure of hydrogen-bonded water molecules. Kitano et al. employed Raman spectroscopy to quantify the number of disturbed hydrogen bonds between water molecules caused by dissolving various hydrophilic polymers in water.^64),71),72) Figure 5 illustrates the number of destroyed hydrogen bonds in water upon dissolution of various water-soluble polymers, defined as the N value. PMPC produces an N value of －0.7, and the copolymer of MPC and BMA (PMB) produces an N value of －0.4. These values are markedly lower than those of representative non-electrolyte water-soluble polymers such as poly(ethylene glycol) (PEG) and poly(N-vinylpyrrolidon) (PVPy), which have N values of approximately 1.0. Polymers with hydrogen-bonded amide groups and polymers with undissociated carboxylic acids exhibit N values of 2-4, and above 5 for the polyelectrolyte, sodium polyacrylate (PAA-Na), and poly(L-lysin) hydrobromide (PLL-HBr), respectively.⁷¹⁾

Fig. 5

(Color online) Comparison of the hydration state of MPC polymers with other water-soluble polymers. MPC: 2-methacryloyloxyethyl phosphorylcholine; PCBMA: poly(carboxybetaine methacrylate); PSBMA: poly(sulfobetaine methacrylate); PEG: poly(ethylene glycol); PVPy: poly(N-vinyl pyrrolidone); PDMAAm: poly(N,N-dimethyl acrylamide); PAAm: polyacrylamide; PAA: poly(acrylic acid); PDMAPMAm: poly(dimethylaminopropyl methacrylamide); PAA-Na: sodium polyacrylate; PVS-Na: sodium poly(vinyl sulfonate); PLL-HBr: poly(L-lysin) hydrobromide.

Water-soluble polymer dissolution typically involves water molecules hydrating the polymer chains. Thus, hydration affects the surrounding water molecule structure. However, hydration is thought to have little or no effect in PMPC. As previously mentioned, the chemical structure of the PC group suggests potential hydrophobic hydration over strong hydrogen bonding or ionic hydration. Due to the MPC unit being electrically neutral, the phosphate anion and trimethylammonium cation are in close proximity, with the three methyl groups attached to the nitrogen atom located outward. This facilitates suitable regions for hydrophobic hydration. Adding van der Waals forces to the water cluster may occur, and they may stabilize further when water molecules cluster around hydrophobic functional groups. The state of water molecules at the phosphatidylcholine-assembled interface is similar to that of the bulk water phase. It has been reported that as measured using two-dimensional heterodyne-detected vibrational sum-frequency generation spectroscopy, the mobility of water molecules was unaffected by the interface.⁷³⁾ Additionally, molecular dynamics calculations have shown that ammonium ions from phosphatidylethanolamine break the hydrogen bonds between the water molecules in the second hydration layer. However, the trimethylammonium ions of phosphatidylcholines weakly bind to water molecules in the primary and secondary hydration layers, thereby increasing the hydrogen bonds between the water molecules.⁷⁴⁾ In other words, hydrophobic hydration is dominant around the PC groups; therefore, the water structure is similar to that in the bulk. Furthermore, this phenomenon occurs at the interfaces where PC groups are introduced onto the substrate surface.

2.6. Nonfouling properties against the biological components.

Biological systems demonstrate various defense mechanisms that maintain life. These mechanisms cooperatively act to protect the living body from pathogens and help prevent excessive bleeding due to injury. In contrast, when medical devices are used for medical treatment and clinical diagnosis, similar biological defense and foreign body recognition reactions occur.^75)-77) In other words, medical devices used in contact with living organisms require good compatibility with tissues. Therefore, this requires a surface to be created that reduces tissue irritation and maintains the functionality of the medical devices for a long period. Interfaces between living tissues are often fluid-mediated environments. In other words, in a complex fluid environment where several types of ions, lipids, and proteins are found in an aqueous medium and where cellular systems and biological tissue are present, the surface must demonstrate an affinity toward these ions, making material design extremely challenging for material science.

Figure 6 shows the biological responses observed on the material surface. When a material interacts with a living system, water molecules initially contact its surface, followed by ions and proteins. Protein adsorption triggers conformational changes of proteins, influencing subsequent cell adhesion, and then thrombus formation, tissue response, and infection. Thus, preventing protein adsorption onto the surface of the material is crucial when designing biomaterials. Furthermore, the formation of protein adsorption layers induces functional deterioration of the medical devices. Blood coagulation involves a complex cascade (hierarchical) reaction, comprising several blood components.⁷⁸⁾ They consist of three additional systems: the coagulation factor, complement, and platelet systems, which act in concert with each other.

Fig. 6

(Color online) Representative biological reactions observed at the interface between artificial substrates and biological systems.

Medical devices, including artificial skin, subcutaneous drug administration ports, and surgical sutures need soft tissue compatibility. The surfaces of percutaneous devices, such as catheters that bridge the internal and external body environments, may benefit from surfaces possessing tissue adhesive properties to deter infections.⁷⁹⁾ In contrast, constructing a surface that does not attach to the cells and tissues is necessary to prevent adhesion and infection caused by postoperative inflammatory reactions.

Additionally, soft tissues in living systems exhibit self-defense mechanisms that manifest as tissue reactions. Inflammatory reactions occur relatively early. During the acute phase, fibrin, red blood cells, and white blood cells are produced, followed by macrophages. Over time, the number of lymphocytes, plasma cells, and macrophages increases, with fibroblast activity, central to tissue repair, becoming more active, thereby leading to fibrosis progression.⁸⁰⁾ The cell reaction and the thickness of the active reaction and fibrogenic layers determine the tissue reaction strength to foreign bodies and materials. Different cells reside in the active reaction layer and are linked to tissue thickening and tumor formation.

Controlling bacterial infection is another important issue when using medical devices.⁸¹⁾ Bacteria multiply and form biofilms using their secretions when they attach to materials. This is impermeable to molecules, thereby preventing drug treatment. Surfaces that inhibit bacterial adhesion and growth must be created to control bacterial infections. Typically, techniques such as the sustained release of bactericidal silver ions are deemed effective; however, their effects on the surrounding tissue must be completely considered. Creating surfaces that hinder bacterial adhesion can help in preventing infection through the immune defense system. The initial state of the protein adsorption layer governs the adhesiveness of cells and tissues, adhesion of bacteria, and formation of biofilm. Therefore, surface designs that inhibit protein adsorption are strongly required.⁸²⁾

Figure 7 shows the effect of surface modification of substrates with MPC polymers on systematic bioreactions.³⁴⁾ It indicates that the bioreactions ranging from protein adsorption (Fig. 7(a)), cell adhesion (Fig. 7(b)), immune response with phagocytes by macrophages (Fig. 7(c)), thrombus formation (Fig. 7(d)), tissue reaction (Fig. 7(e)), and infection by bacteria adhesion and biofilm formation (Fig. 7(f)) can be suppressed effectively by the MPC polymers.

Fig. 7

(Color online) Effects of MPC polymer modification on various biological reactions. (a) Protein adsorption reduction, (b) cell adhesion inhibition, (c) phagocytosis prevention, (d) antithrombogenicity, (e) tissue compatibility, (f) bacteria adhesion/biofilm formation inhibition. Scale bar indicates 10 μm. MPC: 2-methacryloyloxyethyl phosphorylcholine; PMMA: poly(methyl methacrylate); PMB: poly(MPC-co-BMA); PSt: polystyrene; PMEH: poly(MPC-co-2-ethylhexyl methacrylate); SPU: segmented polyurethane; SUS: stainless steel.

2.6.1. Protein adsorption resistance.

Figure 8 illustrates a schematic description of protein adsorption on the polymer surface.^36),83) Water molecules are removed from the contact sites between the amino acid residues and the polymer surface. Some water molecules may remain trapped in spaces between the protein and surface if the spaces are larger than the size of a water molecule. The dotted circles represent hydrophobic amino acid residues. Solvation energy depends on the distribution of hydrophobic amino acid residues on the protein surface and polymer surface hydrophobicity.

Fig. 8

(Color online) Mechanism of protein adsorption and detachment. The figure was modified based on that in Ref. 36.

Attractive solvation interactions (hydrophobic interactions) occur when a water molecule is bound to a hydrophobic surface, the strength of which depends on the hydrophobicity of the surface or the surface group. Thus, computer calculation models should consider hydrophobic and hydration interactions, which act in attraction and repulsion, respectively.

By considering the partition coefficients of the amino acid residues and the repeating units of the polymers, the hydrophobic interaction energy for protein adsorption on polymer surfaces can be calculated. The attraction between hydrophobic molecules in water is known to be unusually strong, and this strong interaction cannot be accounted for by van der Waals forces. Although hydrophilic amino acid residues tend to be present on protein surfaces, attractive hydrophobic interactions are crucial during protein adsorption. This is partially due to some hydrophobic amino acid residues being present on the protein surface but mainly to water molecules being eliminated from the hydrophobic polymer surfaces. Upon contact with the surface, the protein may undergo spontaneous conformational changes due to alterations in its hydration state. A more stable protein form consists of hydrophobic amino acid residues in contact with the polymer surface.

Fundamental research has shown that MPC polymer surfaces prevent the adsorption and structural changes of different proteins present in body fluids.^84)-86) In all biological reactions, protein adsorption is the initial process, and achieving biocompatibility is essential. Thus far, the suppression effect of protein adsorption of MPC polymers is considered to be due to their unique hydration structure, in which a large amount of free water-like water exists on the surface.^61),81)

Using interface science techniques, Whitesides et al. considered material surfaces that do not induce protein adsorption and listed the following characteristics: (1) hydrophilic, (2) containing hydrogen bond acceptors, (3) not containing hydrogen bond donors, and (4) having an overall neutral charge.⁸⁷⁾ Water molecules form hydrogen bonds with one another and form clusters (molecular aggregates). A substance must have a hydrated structure when it dissolves in water, with water molecules arranged around it. In several cases, functional groups that form hydrogen bonds with water molecules and hydrates are necessary. Additionally, water molecules are oriented around their surroundings because of strong electrostatic interactions with polarized water molecules and hydrates when a substance has ionic functional groups. Furthermore, low-polarity hydrocarbon groups being directly hydrated with water molecules is challenging, and only extremely weak van der Waals forces act on them. Therefore, it is thermodynamically advantageous for water molecules to form hydrogen bonds with each other, leading to a free water-like cluster structure.

Oxygen and nitrogen atoms, which have relatively high electronegativities, are hydrogen bond acceptors. Additionally, hydrogen atoms that are covalently bonded to these atoms are hydrogen bond donors. In other words, nonionic hydroxyl, amide, and ionic amino and carboxyl groups are hydrophilic but are not effective in suppressing protein adsorption.⁸⁸⁾ Considering the MPC polymer structure, no hydrogen atoms are directly bonded to the oxygen or nitrogen atoms present in the molecule. Additionally, the charge is neutralized by the formation of salts within the molecule. Thus, the MPC polymer satisfies these four requirements for suppressing protein adsorption. As previously mentioned, PC groups can interact with water molecules extremely weakly and do not destroy the water structure.⁷²⁾ The effect of protein inhibition on the MPC surface was interpreted from a molecular perspective.

Considering protein adsorption and desorption mechanisms, even when a protein comes in contact with an MPC polymer surface, the interaction force at the interface is weak and does not cause protein structural change; therefore, the protein easily detaches from the contact interface.⁸⁶⁾ Measurements of the secondary protein structure in contact with polymer surfaces were performed using circular dichroism spectroscopy. The α-helix contents of bovine serum albumin and bovine plasma fibrinogen do not substantially change from that in bulk solution after contact with the PMB surface.⁸⁵⁾ In contrast, a remarkable conformational change in the protein was induced at the interface with PHEMA hydrogel.

The protein amount adsorbed on the PMB from human plasma determined via radioimmunoassay and immunogold-colloid labeling techniques revealed that the adsorbed protein amounts―major components such as albumin, γ-globulin, and fibrinogen as well as minor components such as the coagulation factor VIII, XII, and complement C5―are quite small and decreases with an increase in the MPC unit composition.⁸⁴⁾ These results have facilitated an increase in studies reported globally on the resistance of protein adsorption to MPC polymers and other zwitterionic polymers.

Recent advances in polymerization techniques have demonstrated the superiority of blood-compatible surfaces prepared using MPC. PMPC is prepared using atom transfer radical polymerization initiated at the substrate surface.^52)-54) A dense polymer brush surface of PMPC can be obtained. The amount of plasma proteins adsorbed on the surface was ＜ 5 ng/cm² when the PMPC chain density on the surface was 0.39 chains/nm².⁶³⁾ This value is extremely low and sufficient to suppress platelet adhesion. These findings indicate that materials with a PMPC brush structure on their surface are completely blood compatible. Well-defined polymer brush surfaces are suitable models for directly measuring the interaction force of a protein with a polymer surface. An atomic force microscopy (AFM)-based methodology can be applied to evaluate the changes in the interaction force.⁸⁹⁾ A specific protein is immobilized onto the AFM probe, and the modified probe contacts and detaches from the polymer brush surface. Force-distance curves are subsequently recorded during this process, and the interaction force between the protein and polymer surface is evaluated.^90),91) Figure 9 illustrates a representative force-distance curve of an AFM probe with immobilized fibrinogen for different polymer brush surfaces. Fibrinogen is a protein that is crucial in thrombus formation and subsequent cell adhesion, although its blood concentration is lower than that of albumin and immunoglobulins. Fibrinogen exhibits a large molecular weight of 330 kDa and three domains with varying functions within the molecule. It is a fibrin precursor, which is the main component of the thrombus. Additionally, fibrinogen contains cell-adhesion ligands and is crucial in the overall biological reactions. Therefore, analyzing the molecular interactions between fibrinogen and material surfaces provides meaningful insights into biomaterial creation. The amount of fibrinogen adsorbed from plasma components has been analyzed using immunoassays.⁸⁴⁾ The amount of fibrinogen adsorbed was extremely low as assessed using radioactively labeled antibodies after contact of human plasma with MPC polymers. However, on glass and poly(BMA) (PBMA), numerous proteins are adsorbed, and a time-dependent increase in the amount of adsorbed fibrinogen is observed, known as the Vroman effect.^92),93) This corresponds to an exchange adsorption reaction in which relatively high protein concentrations are adsorbed on the surface. No intermolecular forces are observed on either the polymer brush surface when the protein comes in contact with the surface. However, the force that the protein receives from the surface during desorption is observed once the protein contacts the surface. This force depends on the types of functional groups present on the surface. Owing to the presence of different amino acid residues, such as in fibrinogen, the interaction force is high when the layer of the polymer brush is charged or when hydrophobic functional groups are present. In some cases, when the protein adheres to the material, electrostatic interaction forces may be considered attractive forces; however, if the ionic strength is sufficiently high, no attraction is observed owing to electrostatic shielding effects. Additionally, protein adsorption onto hydrophobic surfaces is due to hydrophobic interactions that originate from changes in the water state occurring during adsorption. Figure 10 illustrates the relationship between the interaction force during desorption and the actual adsorption amount.⁹⁰⁾ A strong linear relationship is observed between these two parameters. From these results, the adsorption force developed owing to the interaction with the surface during the initial protein adsorption is reflected in the resulting protein adsorption amount. In other words, surfaces can be constructed to prevent protein adsorption if intermolecular interactions can be inhibited, and the PC group is an example that effectively fulfills this role.

Fig. 9

(Color online) Force-distance curve of a fibrinogen immobilized on an AFM probe and different polymer grafted substrates. AFM: atomic force microscopy; PTMAEMA: poly(2-(N,N,N-trimethylammonium)ethyl methacrylate chloride); PSPMA: potassium poly(3-sulfopropyl methacrylate); PMPC: poly(2-methacryloyloxyethyl phosphorylcholine); PBMA: poly(n-butyl methacrylate). The figure was modified based on that in Ref. 90.

Fig. 10

(Color online) Interaction force of fibrinogen immobilized on an AFM probe and different polymer substrates (a) and the relationship between the interaction force and amount of adsorbed fibrinogen on the polymer substrates (b). PTMAEMA: poly(2-(N,N,N-trimethylammonium)ethyl methacrylate chloride); PSPMA: potassium poly(3-sulfopropyl methacrylate); PMPC: poly(2-methacryloyloxyethyl phosphorylcholine); PBMA: poly(n-butyl methacrylate).

2.6.2. Inhibition of cell adhesion and reduction in tissue response.

The initial process of cell adhesion to a substrate involves using an adsorbed protein layer, wherein cell-adherent proteins capture cells through a specific ligand-receptor combination.^94)-96) The cells, once they adhere to the substrate, produce extracellular matrix proteins and form more organized tissues. This natural healing process is necessary for medical implants, such as long-term insertion catheters, medical sensor devices, and tooth implants, to prevent infection from outside the body. However, cell adhesion can result in severe bioreactions and thrombus formation in blood-contacting devices.⁸²⁾ Platelets adhere to, activate, and aggregate with fibrin resulting in the formation of a blood clot. Recent biological technologies have used microfluidic devices and microparticles to evaluate cell functionalities and drug-cell responses.^97)-99) Therefore, surfaces that can control cell adhesion via protein adsorption are urgently required.

The MPC polymer surface can reduce or sometimes inhibit cell adhesion.^100)-103) No capture of the cell through ligand-receptor interactions occurs because cell-binding proteins are hardly adsorbed on the MPC polymer surface. With half of the area of a tissue culture plate coated with the MPC polymer (PMB), when fibroblast cell culture was performed (Fig. 7(b)),³⁴⁾ cells did not adhere to or proliferate on the surface of the MPC polymer. On the untreated surface, the cells adhered, grew, and formed a cell monolayer. In other words, creating surfaces that allow cells to adhere to and proliferate as well as creating surfaces that completely prevent this is easily possible. Thus, creating cell adhesion sites with higher precision is possible by combining micropatterning and inkjet printing technologies. Cells have been shown to orient and adsorb when the surface is treated in a 200-μm wide striped pattern. MPC polymers are useful in controlling cell adhesion. Furthermore, when cell-specific ligand molecules incorporate onto a MPC polymer surface, it can be used as a cell separation device or three-dimensional cultured tissue plate.^104)-106)

When several cell types, including cancer cells, are cultured on substrates treated with MPC polymers, cell adhesion is prevented, regardless of the cell type. This indicates that the MPC polymer is highly effective in suppressing protein adsorption, which is the initial process of cell adhesion. In other words, the amount of cell-adhesion protein adsorbed does not reach the threshold for capturing cells or provide a sufficient scaffold for adhesion.

2.6.3. Bacterial adhesion and resistance to biofilm formation.

Bacteria proliferate and produce specific molecules when they adhere to the surface of a material, creating a favorable environment for their survival.^107),108) This complex of bacteria and the produced molecules forms a biofilm (see Fig. 6). Bacterial infections can be a major obstacle when medical devices are applied to living organisms, such as during catheterization or replacing artificial valves or joints, for treatment in medical procedures. Biofilm formation is challenging to prevent because bacteria defend themselves by reducing the therapeutic effects of common antibiotics and isolating themselves from the immune system. Therefore, when medical devices are used, bacterial contamination must be adequately controlled.

Pseudomonas aeruginosa adhering to various substrate surfaces has been investigated.¹⁰⁹⁾ Bacterial adhesion occurs through direct contact with substrate surfaces, and properties such as surface hydrophilicity, hydrophobicity, and charge state are deemed to have a substantial impact. Moreover, the hydrophobicity of bacteria is also considered a parameter involved in adhesion. The effects of coating stainless steel (SUS) surfaces with MPC polymer (PMB) on the adhesion and biofilm formation of P. aeruginosa, Staphylococcus aureus, and Escherichia coli was investigated (Fig. 7(f)).¹¹⁰⁾ The results demonstrated that surface modification with PMB effectively inhibited bacterial adhesion and biofilm formation on SUS. Furthermore, the bactericidal effect of antibiotics was substantial. Thus, several researchers have investigated the combination of MPC polymer coatings and sustained release of antimicrobial reagents. MPC polymers have been used for the surface treatment of metal implants in orthopedic and dental applications.^111)-115)

2.7. Lubrication by hydrated surface.

Living biological tissue is a water-dominated environment with a small water layer between the tissue and medical device. Therefore, lubricating the medical device surfaces must be considered to reduce the impact on biological tissue and prevent performance degradation of medical devices due to tissue reactions over a long period.¹¹⁶⁾ Several researchers have reported that hydrophilic polymers bonded to the surface of a solid substrate retain a water layer.^117)-121) Water is attracted to hydrophilic polymers on the surface, forming a hydration layer that plays a crucial role in lubrication. However, the flexibility of the polymer’s hydration layer does not support the load on its properties. Therefore, the surface layer of water may accept most of the load. Friction forces act on the sliding surfaces because the polymer chains adhere to opposing surfaces over time. This implies that water slowly exudes from the surface of the polymer hydration layer under a load, with or without sliding. As water is lost, the surface layer thickness decreases, and the water content of the surface layer decreases. Consequently, the degree of adhesion of this surface to the opposite bearing surface increases with frictional force. Therefore, friction depends on the water content of the surface layer. Thus, stable “hydration lubrication” may result in low friction, leading to low wear of the material.^122)-124) In living systems, joints are mainly composed of bone substrates and hydrated biological molecules. This hydrated layer of polysaccharides and lipids can stably hold water and slide together through the mechanism of hydration lubrication.

Measuring the kinetic friction coefficient of a polymer substrate coated with a physically hydrophilic polymer over time initially shows good lubricity but subsequently increases to that of an untreated substrate. This suggests that the hydrophilic polymer-coated layer peeled off from the substrate surface because of sliding. Therefore, physical bonding via molecular interactions of surface-modified polymers is ineffective for large multidirectional loads. In contrast, polymer substrates grafted with hydrophilic polymers demonstrated extremely low and stable kinetic friction coefficients and volumetric wear compared with untreated substrates and substrates coated with hydrophilic polymers. To provide long-term stability, the surface hydrophilic graft polymer layer should be bonded to the substrate by chemical covalent bonds.

A polymer brush surface of polymer electrolyte was prepared, and the lubricity of the interface was evaluated.¹¹⁷⁾ Surface-initiated living polymerization provides brush-like polymer structures at the surface and allows chain density and length control of the grafted polymer. The polyelectrolyte brush surface exhibited excellent lubricating properties in water. This is a phenomenon in which the electrostatic repulsion between the polymer chains has a crucial effect. However, inorganic salts present in the in vivo environment can have an impact. In other words, for application to the sliding surfaces of artificial joints, the key is to form a polymer brush layer unaffected by the coexistence of different ions and biomolecules.

A uniform polyelectrolyte layer was formed on the polymer surface via graft polymerization using a photoreaction method.¹²⁵⁾ The thicknesses of the layers were 100-150 nm. The kinetic friction coefficients of the resulting surfaces in water ranged from 0.01 to 0.05, which were reduced by 40% to 85% compared with the kinetic friction coefficient of the untreated polymer substrate. The influence of the chemical structure of the grafted polymer chains on lubricity has been reported.¹²⁶⁾ The kinetic friction coefficients of poly(oligo(ethylene glycol) monomethacrylate) (POEGMA) and PMPC graft surfaces did not markedly differ for different lubricants. For all lubricants, the coefficient of kinetic friction of the PMPC-grafted substrates was considerably lower than that of the POEGMA-grafted substrates. Similarly, the kinetic friction coefficients of poly(N,N-dimethylaminoethyl methacrylate) (PDMAEMA) and poly(2-methacryloyloxyethyl phosphoric acid) (PMPA)-grafted substrates dramatically increased in simulated body fluid (SBF) and bovine serum, respectively. Furthermore, untreated substrates revealed increased friction with bovine serum lubricant compared with water. These findings show that the charge type on the grafted polymer layer (nonionic, cationic, anionic, or zwitterionic) affects the hydration and friction rates of the lubricated substrate surface. For instance, PDMAEMA-grafted substrates demonstrated a higher kinetic friction coefficient in bovine serum-containing proteins, such as albumin and γ-globulin, than in water or SBF. This is attributed to the fact that PDMAEMA becomes positively charged in the biological environment and induces adsorption when it comes in contact with proteins like albumin. PDMAEMA has positively charged -NH^＋(CH₃)₂ groups at neutral pH. These groups subsequently bind to negatively charged molecules. For instance, albumin molecules are negatively charged at physiological pH (pH 7.4). This implies that the presence of protein molecules at the bearing interface increases the resistance to sliding motion. In contrast, the negatively charged PMPA-grafted substrate surface demonstrated considerably lower lubricity in SBF during ball-on-plate friction tests. The chemical structure of negatively charged PMPA is characterized by numerous trapping sites for positively charged inorganic ions. Therefore, the poor surface lubricity of the PMPA-grafted substrates is because of the shrinkage or cross-linking of the negatively charged polyelectrolyte chains. This reduces the mobility of chains in solutions containing positively charged inorganic ions.

By mimicking the structure of this biogenic joint, using MPC polymers, particularly PMPC, with excellent hydration properties for the surface treatment of medical devices requiring sliding surfaces is advantageous.^123),127) The surface of the PMPC graft base material has an extremely low dynamic friction coefficient of 0.02-0.04 in water. Its performance remains stable over time, even in physiological conditions, including serum and plasma.

3. Biomedical applications

Table 2 lists representative medical devices that use MPC polymers and which have already been applied in clinical practice.^6),40)

Table 2

Medical devices using MPC polymers

Medical device	Product name	Modification process	Clinical introduction
Artificial heart (LVAD)	EVAHEART	Coating	2011
Stent
Cardiovascular	BiodivYsio	Reacting	2000
Drug eluting	Endeavor Endeavor Sprint	Reacting	2010
Neurovascular	Pipeline with Shield Technology	Reacting	2015
Artificial lung	Synthesis	Coating	2005
Artificial hip joint	Aquala	Grafting	2011
Gide wire/Catheter	Hunter	Reacting	1997
	Orphis CV Kit	Coating	2018
Contact lens
Soft hydrogel	Proclear (Omafilcon A)	Copolymerization	2005
Silicone hydrogel	Total 30 (Lehfilcon A)	Grafting	2022

MPC: 2-methacryloyloxyethyl phosphorylcholine; LVAD: left ventricular assist device.

3.1. Implantable artificial organs.

3.1.1. Cardiovascular stent.

The treatment of acute blood vessel closure involves using a coronary stent. Since 1995, MPC polymers have been coated onto their surfaces to reduce the thrombogenicity of metal stent frames. The polymer was based on the PMD addition of a 2-hydroxypropyl methacrylate and 3-trimthoxysilylpropyl methacrylate units as reactive moieties.^128),129) The reaction between the hydroxyl and triethoxysilyl groups in the polymer promotes the polymer layer stabilizing on the metal stent frame, even when the device is in its expanded state in the blood vessel and exposed to blood flow. This stent, termed BiodivYsio AS^® (Biocompatible. Ltd), is effective in preventing initial thrombus formation and was approved by the FDA in 2000. Stents were implanted in patients, and after the success of this particular device, a drug-eluting stent using an MPC polymer (Endeavor^® Sprint, Medtronic) has been developed.¹³⁰⁾ This approach is even more efficient in reducing in-stent restenosis and target lesion revascularization. These MPC polymers are highly effective as coating materials for drug-eluting stents because MPC polymers composed of alkyl methacrylates exhibit antithrombogenicity and good solute permeability. A bioactive reagent, Zotarolimus, which can suppress excessive cellular ingrowth, has been incorporated into the device, and Endeavor^® Sprint stents have been clinically implanted since 2010.

3.1.2. Artificial heart.

The development of artificial hearts dates back to the 1950s, initially featuring pulsatile designs. Recently, these have shifted to centrifugal pump types that use motor-driven propellers.¹³¹⁾ In this process, titanium is used as the base material because it are relatively resistant to blood clotting when polished to a mirror finish. Patients with artificial heart implants typically take antiplatelet drugs to suppress blood clotting. However, microthrombi can still form, depending on the state of blood flow and reactions with the surface, potentially blocking blood vessels in the brain and other areas. To mitigate this risk, an MPC polymer (PMB) was applied to the pump’s inner and outer surfaces and the artificial blood vessels connecting the pump to the heart.^132)-134) Repeated in vivo animal experiments demonstrated that the device could function for over 820 days without anticoagulation therapy. Clinical trials for Japan’s first implantable artificial heart (EVAHEART^®) began in 2005, and in 2011 it became available for clinical use.¹³⁵⁾ Over 240 cases have been reported, with patients thriving for up to 15 years. Since 2018, clinical trials of EVAHEART^® have been underway in the United States. The main purpose is to support blood circulation in patients with dilated cardiomyopathy; however, EVAHEART^® enables patients to transition from hospital care to normal social activities.

3.1.3. Neurovascular stent.

The development of new stents was prompted because of the discovery that surface modification of metal stents with MPC polymers can suppress thrombus formation at the initial stage of introduction and does not interfere with subsequent endothelial tissue reaction.¹³⁶⁾ Previously, by fixing the affected area with metal clips during craniotomy, cerebral aneurysms were treated.¹³⁷⁾ However, endovascular surgery has become widespread because it is less invasive. A method of treating an aneurysm was used in which a catheter was inserted into the blood vessel, and a platinum coil was placed in the affected area for blood flow suppression in the aneurysm.¹³⁸⁾ A demand existed for a higher occlusion rate for aneurysms, and new treatment techniques were required. In 2010, the use of vascular stents in cerebral artery treatment was considered, and several cerebrovascular devices made of metal were developed.¹³⁹⁾ These were compared with coronary artery stents. They exhibit thinner struts that make up the mesh and a higher mesh density. Initially, elastic modulus and flexibility were considered crucial, and the devices were fabricated by weaving metal wires. In 2015, attempts were made to add blood compatibility and lubricity to metal surfaces, one of which involved covering them with PC groups based on the MPC polymer chemistry.^140)-142) In vitro and ex vivo studies revealed that PC group-immobilization treatment was extremely effective. Furthermore, it was introduced into clinical practice as a Pipeline^TM Flex Embolization Device with Shield Technology^TM (Medtronic). Covalent bonding between the metal and PC groups was achieved using a silane coupling agent, and the polymer layer generated at the surface was extremely thin (3 nm in thickness). The immobilized PC group layer inhibits blood coagulation reactions on the surface, resulting in a 95% reduction in platelet adhesion and a 55% reduction in thrombin generation, without affecting the overall flexibility and modulus of the device. Additionally, it can provide a good clinical setting because of its excellent lubrication properties. The device has been properly set up in affected areas in clinical practice because of its excellent lubricity.

3.1.4. Artificial hip joint.

Total hip arthroplasty with an artificial hip is an effective treatment for severe arthritis, with its prevalence increasing due to aging populations.¹⁴³⁾ Surgical techniques have improved, including designing implants and applying robotic technology. However, periprosthetic osteolysis remains a serious issue that markedly limits the quality of life of patient. Up to 20% of patients implanted with conventional polyethylene hip joints develop prosthesis loosening within 10 years, and approximately half are disabled due to pain and loss of function.¹⁴⁴⁾ Consequently, the number of revision surgeries is increasing and is predicted to double by 2035. Unless a limiting mechanism is introduced to prevent periprosthetic osteolysis due to material wear, social and economic impacts linked to the treatment will continue to increase.

The human body maintains joint lubrication through synovia, comprising polysaccharide derivatives, proteins, and lipids.¹⁴⁵⁾ Restoring the joint function in which the same level of lubrication is artificially achieved is crucial. One method involves mimicking the surface functionality of synovia using artificial materials. The preparation of a PMPC graft layer on the substrate was considered to prevent wear by reducing the friction generated at the interface.^125),127) Super-low-friction surfaces were obtained by employing the established procedure for surface-initiated graft polymerization of MPC. A simple, new methodology involving MPC photoinduced polymerization on an ultra-high-molecular-weight polyethylene (UHMWPE) surface was developed.^127),146) The water-insoluble photoinitiator benzophenone was coated onto the surface of the UHMWPE substrate from an acetone solution and subsequently exposed to UV light in an MPC aqueous solution. MPC polymerization was initiated from the polyethylene substrate surface, and the molecular weight of the resulting PMPC correlated with the photoirradiation period.

Figure 11 illustrates some of the characteristics of PMPC-grafted UHMWPE substrates.^124),127) The thickness of the PMPC graft layer was approximately 100-200 nm. Analysis of the friction behavior of the sliding surface of the joint showed that the dynamic friction coefficient of the PMPC-treated UHMWPE substrate was approximately 0.002, which was approximately one-seventh of that of the untreated UHMWPE substrate. This is an extremely small value compared with the friction coefficient of a living cartilage surface. A mechanical walking simulation test loaded with 280 kgf was conducted on the modified and unmodified UHMWPE substrates as a liner using a hip joint simulator in an environment similar to that of a living organism. PMPC presence was confirmed on the sample surface even after 20 million tests (equivalent to 20 years of actual walking) (Fig. 11(c)). Moreover, the wear-inhibiting effect persisted even after ＞ 70 million continuous walking loads.¹⁴⁷⁾

Fig. 11

(Color online) (a) Transmission electron microscopic image of PMPC-grafted ultra-high-molecular-weight polyethylene (UHMWPE), (b) illustration of a PMPC-grafted UHMWPE liner of on artificial hip joint, (c) amount of wear of the crosslinked UHMWPE (CLPE) liner and PMPC-grafted CLPE liner during mechanical friction tests, (d) X-ray image of a human hip joint before artificial hip joint replacement surgery (Pre) and after 5 years after implantation, (e) the Japanese Orthopaedic Association (JOA) score after implanting an artificial hip joint installed using a PMPC-grafted liner. PMPC: poly(2-methacryloyloxyethyl phosphorylcholine). The figure was modified based on that in Refs. 126, 147, and 148.

Moro et al. applied this technique for the surface modification of a crosslinked UHMWPE (CLPE) cup of an artificial hip joint.¹⁴⁷⁾ The surface exhibited high lubricity, termed fluidic friction, and the wear of the PMPC-grafted CLPE cup was considerably reduced compared with that of the bare CLPE and that grafted with other hydrophilic polymers. Additionally, the PMPC-modified polyethylene particles did not stimulate osteoclast cells that cause bone resorption, whereas normal polyethylene particles generated from the friction wear of the artificial joint-induced osteolysis.¹²⁵⁾ In Japan, a novel hip joint system with a PMPC-grafted CLPE cup, Aquala^® (Kyocera) has been approved, with the joint expected to be suitable as a long-term hip replacement, in comparison with conventional alternatives (Fig. 11(d)). The Japanese Orthopaedic Association (JOA) score, which reflects the quality of life of patients, improved in the first year and was maintained for 10 years after the replacement operation (Fig. 11(e)).¹⁴⁸⁾ Over 100,000 hip joints with PMPC-grafted CLPE cups were implanted in Japan between November 2011 and September 2024. This surface modification technology with PMPC is also applicable for artificial knee joint development.¹⁴⁹⁾

3.2. Contact lenses.

Since the 1990s the characteristics of MPC polymers have been used to develop soft hydrogel contact lens materials.¹⁵⁰⁾ The material consists of MPC, HEMA, and a small amount of crosslinker monomer, and the resulting hydrogel has a water content of approximately 60%. This material, named Omafilcon A, is used by several contact lens companies globally.

A requirement exists for new contact lens materials with high functionality and long-term wearability. Modification of the silicone hydrogel surface of the contact lenses with MPC polymers, focusing on the surface structure of the cornea, has been reported.^151)-154) Hydrophilic biomolecules assemble on the corneal surface to form a hydrogel-like structure. Tear fluid can diffuse and move relatively freely, and oxygen and nutrients are supplied. A specific characteristic is hydrophilic biomolecule binding such as the binding of mucin to the surface, which constitutes a highly mobile interface.

This method is characterized by the uniform reaction of MPC polymers with reactive functional groups near the surface, which enables a clear definition of the grafted surface structure. The mechanical strength and optical properties of silicone hydrogels do not change after the reaction. Figure 12(a) illustrates near-surface observations of the MPC polymer (PMPC)-modified silicone hydrogel. When observed using an environmentally controlled scanning electron microscope in a wet environment, different contrast layers are observed on the silicone hydrogel material surface. Near-surface observations using AFM revealed that the PMPC layer on the surface was approximately 200-nm thick and homogeneously covered the surface (Fig. 12(b)), indicating the formation of a hydrated PMPC layer on the surface of the silicone hydrogel material. Transmission electron microscopy in the vicinity of the surface confirmed PMPC formation grafted onto the surface, which was similar to the corneal surface structure. The untreated silicone hydrogel material exhibited a high modulus of elasticity in the near-surface area, whereas surface modification with PMPC lowered the modulus to approximately one-ninth because of the highly hydrated surface state. This value is comparable to the cornea’s elastic modulus. The structure of the PMPC graft layer on the surface did not change after 30 days of wear and scrubbing (Fig. 12(c)).

Fig. 12

(Color online) (a) Environmental scanning electron microscope image, (b) AFM image, and (c) transmittance microscopic image of the interface between a PMPC layer and silicone hydrogel contact lens substrate (Lehfilcon A), (d) lipid adsorption and absorption to a silicone hydrogel contact lens, (e) relationship between the equilibrium water content of contact lenses and oxygen permeability. AFM: atomic force microscopy; PMPC: poly(2-methacryloyloxyethyl phosphorylcholine). The figure was modified based on that in Ref. 155.

The surface friction coefficient of the PMPC-modified silicone hydrogel is less than one-third that of untreated material, closely matching the friction coefficient of the living cornea. This significantly enhances comfort in wearing contact lenses, reduces impact on the cornea and eyelids, and protects living tissue. This silicone hydrogel material with an PMPC graft-modified layer on its surface was named Lehfilcon A in 2022.¹⁵⁵⁾

Figure 12(d) illustrates lipids adsorption when silicone hydrogel contact lenses were immersed in lipid (cholesterol ester and triglyceride lipids) solutions for 30 days.¹⁵⁶⁾ Lipids are adsorbed on the surface and diffuse into the interior of the material in conventional silicone hydrogel materials because the silicone phase present in the material is exposed on the surface layer, which is the lipids’ pathway for adsorption and diffusion. However, no lipid adsorption was observed on the surface or inside the lens when the surface was modified using a PMPC graft layer. Additionally, the contact angle of oil droplets on hydrated PMPC-grafted surfaces was shown to be ＞ 170°.⁶¹⁾ This corresponds to the fact that the oil droplets barely adhered to the surface and were easily removed by gentle washing. Furthermore, proteins and bacteria barely attached to the surface.

Figure 12(e) shows the relationship between the water content and oxygen permeability of the contact-lens materials. Several contact lens materials have been developed for this purpose. Oxygen permeability increased depending on the water content for contact lens materials with a common water-containing gel. Large oxygen permeation coefficients were observed in silicone hydrogel materials aimed at improving oxygen solubility. However, their low water contents rendered them uncomfortable to wear. The oxygen permeability tended to decrease as the water content increased when hydrophilic functional groups were introduced to solve this problem. Silicone hydrogel lens materials have been developed to solve this problem, with a water content of approximately 45% and an oxygen permeation coefficient of over 100 × 10^－11((cm²/sec)(mL O₂/mL × mmHg)). Lehfilcon A, in which MPC polymer is grafted near the surface, has a water content of 55% (almost 100% at the surface) and an oxygen permeation coefficient of 127 × 10^－11((cm²/sec)(mL O₂/mL × mmHg)), a good balance between these two parameters.¹⁵⁵⁾ Contact lenses made from Lehfilcon A have been sold globally since 2022 as long wearable lenses (Total 30^TM, Alcon).

3.3. Other biomedical devices.

In recent years, new testing methods and treatments that support medical technology have been researched using cells as targets.^157),158) Cell diagnosis, cell therapy, and the production of useful bioactive molecules using cells as production reactors are major goals in molecular biology, cell biology, and cell engineering.^159),160) In particular, stem cell and iPS cell availability will accelerate progress in this field. However, traditional cell culture methods are used when handling cells, and advances are necessary for technological development. In artificial environments, cultured cells are produced and their growth and differentiation processes differ from those of living cells. The future challenge is to make this approach closely mimic cellular response in a biological environment. Adjusting cell adhesion to cell culture substrates can be associated with the control of cell responses. Cells communicate with each other via the extracellular matrix. The artificial reproduction of this situation is necessary for creating biological tissue analogs via three-dimensional culture. With these considerations in mind, MPC polymer applications that can control cell adhesion properties are being studied at present. First, when a cell suspension is introduced into a normal cell culture plate, cells adhere to the bottom surface of the substrate and wall surface.¹⁶¹⁾ Creating cell aggregates that can be controlled to a certain size is necessary to form target cell tissues using stem cells. In this case, conventional cell culture plates exhibit low efficiency. However, some cells will form aggregates within each well if a well plate with a U-shaped bottom is used. At this point, eliminating cells that adhere to the well wall is necessary. Furthermore, coating the surface with an MPC polymer (PMB) can effectively prevent cell adhesion to the wall. This increases the efficiency of stem cell differentiation in the production of artificial tissues.¹⁶²⁾

Similarly, microfluidic channels manufactured using microfabrication techniques are used to separate cells via flow control or to produce useful compounds by enabling their attachment to specific locations.^163)-166) In addition, microfluidic devices have been used to verify drug safety and efficacy. MPC polymer modification in microfluidic channels to control cell adhesion has been investigated. Using the non-cell-adhesion properties of MPC polymers, co-immobilizing ligand biomolecules that specifically react with cells can be used to induce cell adhesion. This technology can be used to define cell separation and also enables patterning and adhesion of multiple cells. Surface treatment technology is applied to cell culture plates and microchannels as well as membrane reactors and selectively permeable membranes by binding different cells to both sides of a membrane.¹⁶⁷⁾

Additionally, MPC polymers are effective in encapsulating and physically immobilizing cells. MPC polymer hydrogels, capable of reversible gel formation and dissociation using chemical crosslinking and intermolecular force interactions between the polymers, are currently being studied.¹⁶⁸⁾ These consists of a water-soluble MPC polymer with phenylboronic acid groups (PMBV), and other water-soluble polymers with numerous hydroxyl groups, and poly(vinyl alcohol) (PVA). Mixing these solutions results in the formation of a crosslinking reaction and hydrogel (PMBV/PVA hydrogel) is obtained.¹⁶⁹⁾ However, phenylboronic acid groups react with low-molecular-weight sugar compounds depending on the binding constant. Therefore, when fructose or sorbitol is added to the generated PMBV/PVA hydrogel, it dissociates and returns to a homogeneous solution. By gelling the cell suspension, the cells can be three-dimensionally fixed within the PMBV/PVA hydrogel.^169)-172) Small changes in the elastic modulus of the gel have been shown to control the growth process despite the cells being maintained in normal cell culture conditions. In other words, cells are fixed in a state where the elastic modulus is high, a medium is added to the cells to swell them further, and cell proliferation is initiated when the elastic modulus is lowered.^173),174) An artificial extracellular matrix that can control cell dormancy and activity is created. Because the cell proliferation cycle affects the induction of cell differentiation, cells can be induced more efficiently than in normal differentiation-inducing conditions by immobilizing stem cells, controlling their cell cycle, and timing the addition of cell differentiation-inducing factors. This enables the induction of differentiation. Finally, sorbitol is added to dissolve the PMBV/PVA hydrogel and allow recovery of the cells. This method is crucial for tissue regeneration when homogeneous cells are required.

Additionally, biosensing and bioimaging devices can be coated with MPC polymers to ensure cytocompatibility and tissue compatibility. MPC polymer-modified biosensing devices exhibit long-term functionality. For instance, a biosensor attached to an artificial pancreas, regulating blood glucose levels through insulin injection,¹⁷⁵⁾ utilized fluorescent quantum dots (QDs) coated with MPC polymers.¹⁷⁶⁾ These QDs, bound to a cell-permeating peptide, octaarginate (R8), penetrated the cell membrane and remained stable within cells, preserving cell function after encapsulation of the MPC polymer-modified QDs. As the cells differentiated, intracellular QDs were distributed within each cell. These results offer a promising technique for tracing implanted cells in tissue regeneration medicine. When cell sheets fabricated from cells with QDs inside were inserted under the skin of pigs and left for 10 days, distinguishing the original tissue from the embedded tissue was possible using fluorescence observation.

4. Conclusions and future perspectives

Thirty-seven years have passed since the successful development of methacrylate monomers with phospholipid polar groups, MPC, employing a biomimetic molecular design.³⁴⁾ Significant advances in synthetic routes and purification processes have enabled industrial-scale production, making MPC available as an industrial material. This has facilitated the commercialization of products for in vitro use, including cosmetics and contact lenses. Through systematic analysis of their reactions with living organisms, polymers containing MPC units have exhibited excellent biocompatibility, from protein to biological tissue levels. This has been recognized as a surface treatment that can effectively solve the problems related to the clinical use of medical devices. Consequently, MPC has found widespread applications in various medical devices globally, demonstrating successful in vivo implantation as part of artificial organs for over a decade, significantly contributing to patient quality of life. Its potential extends to tissue regeneration medicine and cellular gene therapy. MPC is now commercially accessible as a reagent, with numerous researchers utilizing it as a single functional monomer, uncovering new functions of MPC polymers.

Recently, MPC polymers have been suggested to be used in separating porous membranes for oil-water mixtures with high efficiency and separation rates,^177)-179) transparent films that prevent frost and freezing as well as contamination from the environment,^180),181) and in lithium-ion batteries to enhance ionic conductivity.^182)-185) Additionally, it can be used as a binder to increase the photoelectric conversion efficiency in perovskite solar cells.^186),187) The application range extends from medicine to environmental and energy problem-solving. By imagining the biomolecular structure and function and combining the necessary elements with those involved in polymer chemistry, these polymers present significant opportunities for functional polymer research across diverse fields.^188)-190)

Advancements in polymer science have revealed diverse attributes of zwitterionic structures through MPC polymer investigations, leading to exploration in polymer research involving sulfobetaine, carboxybetaine, choline phosphate groups, and amino acid residues.^{43)-45),187),191),192)} In addition, there is increasing focus on hybridizing polymer chains with positive or negative charges to study the properties of zwitterionic polymer-like surfaces.^{43),193),194)} Notably, research on MPC polymers is paving the way for a fresh realm of functional polymers.

Acknowledgments

The author offers his grateful appreciation to all collaborators, staff, and students who contributed to the progress of this research. He also thanks the companies for developing medical devices based on his research. Most importantly, he extends his sincere thanks to his family for their support during his research career.

Notes

Edited by Susumu KITAGAWA, M.J.A.

Correspondence should be addressed to: K. Ishihara, Division of Materials & Manufacturing Science, Graduate School of Engineering, Osaka University, 2-1 Yamada-oka, Suita, Osaka 565-0971, Japan (e-mail: [e-mail]).

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Profile

Dr. Kazuhiko Ishihara was born in Osaka in 1956. He graduated from Waseda University and received his Ph.D. degree in applied chemistry in 1984. He then worked as a research associate at the Sagami Chemical Research Center from 1984 to 1986, moved to Tokyo Medical and Dental University in 1987 and started research on molecular design of bio-inspired polymers for medical applications, and was an associate professor at Tokyo Medical and Dental University from 1991 to 1998. He became an associate professor in the School of Engineering at the University of Tokyo in 1998, a professor there in 2000 in the Department of Materials Engineering, and a co-professor in the Department of Bioengineering in 2006. After he retired from the University of Tokyo (2021), he became an emeritus professor at the University of Tokyo. He is currently a Specially Appointed Professor at Osaka University and a Visiting Professor at Kansai University. During this period, he has been a sub-project leader of the 21st Century COE, “Creation of Human-Friendly Materials Based on Chemistry” project and “The University of Tokyo Center for Nano-Bio Integration Research”. He is also a project leader of the Ministry of Education, Culture, Sports, Science and Technology Grant-in-Aid for Scientific Research on Innovative Areas “Nanomedicine Molecular Science” and the Japan Science and Technology Agency S-Innovation Project “Development of Ultra-Biofunctional Surface Construction Technology Based on the Creation of Materials Photochemistry”. To implement his research results in society, he has collaborated with companies under the Japan Science and Technology Agency Commissioned Development Scheme and has received two successful accreditations (1999 and 2011). He has received academic awards from the Society of Polymer Science, Japan, the Japanese Society for Biomaterials, and the Japanese Society for Artificial Organs, as well as the Inoue Harushige Award (2004), The American Academy of Orthopedic Surgeons Frank Stinchfield Award (2006), Society for Biomaterials Clemson Award (2009), Minister of Economy, Trade and Industry Award for Advanced Technology (2011), Prizes for Science and Technology from Ministry of Education, Culture, Sports, Science and Technology, Japan, Development Category (2017), The Prize of Ministry of Economy, Trade and Industry, National Commendation for Invention (2018), The Japan Medical Research and Development Grand Prize, Award of Ministry of Health, Labour and Welfare, Japan (2018), Technology Management and Innovation Award, Society of Science, Technology and Economy (2020), Special Award of Bioindustry, Japan Bioindustry Association (2021), NIMS Award (2022), among others, and was awarded the Medal with Purple Ribbon (2024).