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The layer-by-layer assembly of sequentially adsorbed, alternating polyelectrolytes has become increasingly important over the past two decades. The ease and versatility in assembling polyelectrolyte multilayers (PEMs) has resulted in numerous wide ranging applications of these materials. More recently, PEMs are being used in biological applications ranging from biomaterials, tissue engineering, regenerative medicine, and drug delivery. The ability to manipulate the chemical, physical, surface, and topographical properties of these multilayer architectures by simply changing the pH, ionic strength, thickness, and postassembly modifications render them highly suitable to probe the effects of external stimuli on cellular responsiveness. In the field of regenerative medicine, the ability to sequester growth factors and to tether peptides to PEMs has been exploited to direct the lineage of progenitor cells and to subsequently maintain a desired phenotype. Additional novel applications include the use of PEMs in the assembly of three-dimensional layered architectures and as coatings for individual cells to deliver tunable payloads of drugs or bioactive molecules. This review focuses on literature related to the modulation of chemical and physical properties of PEMs for tissue engineering applications and recent research efforts in maintaining and directing cellular phenotype in stem cell differentiation.
Layer-by-layer (LbL) assembly of polyelectrolytes (PEs) is a method for preparing highly tunable thin film polymer coatings. Such coatings or films are denoted henceforth as polyelectrolyte multilayers (PEMs).1–3 PEMs self-assemble due to electrostatic interactions between sequentially deposited, alternately charged PEs.2,4–6 The chemical composition and concentration of the individual PEs direct the self-assembly process of PEMs. Since electrostatic interactions enable the links between PEs, multilayer properties can be modulated through the adjustment of deposition conditions. Typically, parameters that are varied are the pH of the PE solution,4,7–12 ionic strength,7,12 the number of layers,13,14 the order in which the layers are deposited,15 and post-assembly modifications.10,11,16–19 A commonly used post-modification process is cross-linking, which further influences film morphology, thickness, and structure.16,17,20 Virtually any combination of cationic and anionic PEs can be used to assemble PEMs with highly tunable properties applicable in a wide range of applications. Such applications include microelectronics,11 nanofluidics,21,22 virucidal coatings,23 drug delivery,24,25 and tissue engineering.26,27 In this review, applications of PEMs in tissue engineering will be discussed.
PEMs have been assembled from synthetic as well as naturally occurring polymers.1–3,28–30 Synthetic PEMs have been assembled from cationic PEs such as poly (allylamine hydrochloride) (PAH) and poly (diallyldimethylammonium chloride) (PDAC) and anionic PEs such as poly (acrylic acid) (PAA) and sulfonated polystyrene (SPS).8,29,31,32 PEMs have also been fabricated from naturally occurring PEs such as, polypeptides, polysaccharides, DNA, and proteins.26,31,33–37 One of the earliest reports of PEMs derived from a combination of synthetic and natural PEs was by Lvov et al. where ultrathin films comprised of DNA and PAH were assembled for applications in ecological sensors.1 More recently, Dubas and Schlenoff explored combinations of poly(styrene) and PDAC to systematically determine the factors governing LbL assembly such as thickness, ionic concentration, and pH.12 Elbert et al. formed PEMs comprised of poly (L-lysine) (PLL) and alginate (ALG) to render the surfaces biologically inert and thereby block the adhesion of human fibroblast cells.33 Berg et al. assembled PAA and arginine-glycine-aspartic acid (RGD)-modified PAH multilayers to promote the cellular adhesion of murine wild-type NT6 fibroblasts and to determine the effect of RGD-PAH spatial configuration (i.e., RGD-PAH line widths ranging from 10 to 50μm) on cytoskeletal protein organization.38 Due to a wide range of PEs available to assemble multilayer films, as well the ability to tailor their properties by varying the deposition conditions, PEMs are being incorporated into a variety of biological applications. Their uses include coatings to either promote or prevent cell adhesion, and more importantly in directing or maintaining cellular phenotype.29 In addition to PEs, multilayer films have been assembled from DNA strands using complimentary base pairing,39 click chemistry utilizing covalent bonding,40 and hydrogen bonding to create degradable multilayers for drug release applications.41 The remainder of this review will focus on the recent literature regarding LbL assembly of PEs and their applications in the field of tissue engineering.
PEMs have garnered immense interest in the field of biomedical engineering to render implantable material surfaces more biomimetic.42 Largely due to their tunable nature, PEMs have had a significant impact in nearly every aspect of biomaterial design and tissue engineering.10,18,22,24,26,27,30,31,38,42–55 PEMs have been used to modulate protein adsorption,33 promote cell adhesion,16,29,34,38,55–57 or regulate the inflammatory response.58–60 In addition, PEM coatings have the potential to mimic the complex extracellular matrix found in vivo. To recapitulate the in vivo environment, studies have been conducted to tailor the surface topography, chemical composition, mechanical properties, and the degradation of PEMs.61–63 In subsequent sections, the effect of chemical and mechanical properties of PEMs and their impact on tissue engineering, cell adhesion, maintaining, and, more importantly, directing cellular phenotype will be discussed. Other novel applications of PEMs in tissue engineering such as drug delivery and drug reservoir construction for controlled release will also be reviewed.
The adhesion of cells in vitro and their subsequent performance is often dependent on the physical properties of the underlying substrate.52,62 The mechanical properties of a PEM can be varied by changing the thickness, the pH of the solutions used to assemble the PEM, and through postassembly modifications. In the following sections, we will review the approaches to tailor the mechanical and chemical properties to optimize cellular response.
Various factors such as the assembly conditions and the number of bilayers in the PEM contribute to the diversity in mechanical properties. In a recent study, the response of smooth muscle cells (SMCs) was directly linked to the Young's modulus of the multilayer on which the cells were cultured.16 Richert et al. demonstrated that the number of layers within the self-assembled film play a dominant role in SMC response.16 PEMs were assembled using PLL and hyaluronic acid (HA) as the cationic and anionic PEs, respectively. Increasing the PLL/HA bilayer number from 20 to 60, directly correlated with increasing thickness from 3 to 15μm, respectively. However, the Young's modulus was inversely related to bilayer number, with values of 90 and 40kPa observed for 20 and 60 bilayers, respectively. Thicker PEMs were less rigid due to a higher degree of hydration.5,64 When SMCs were cultured on the PLL/HA multilayer, projected cell areas were observed to be ~1100μm2 on substrates comprised of 20 bilayers, compared to 800μm2 for a substrate coated with 60 bilayers.16 In another study by the same group, the adhesion of human chondrosarcoma cells increased with decreased PLL/poly (L-glutamic acid) (PLGA) PEM thickness.56 Although the thickness of a PEM can affect cellular adhesion and spreading, it appears that swelling and mechanical properties such as Young's modulus may play a more significant role in cellular response.56
A self-assembled PEM can grow in a linear2 or exponential manner33 (Fig. 1). Linear growth typically occurs when the PEs are highly charged and do not diffuse freely throughout the PEM.2 Highly charged PEs exhibiting linear growth include cationic PAH and anionic SPS.2 On the other hand, exponential growth occurs in the presence of weak PEs (PLL, ALG, and PLGA), characterized by diffusion, and hydrogen bonding.13,33,34,65 In a few cases, linear growth in the presence of weak PEs has also been observed.66 Exponential PEM growth is largely due to the diffusion of at least one PE through the film,34,65 resulting in an exponentially increasing thickness, as additional bilayers are added to the film.66,67 Picart et al. attributed the exponential growth observed in PLL/HA PEMs to the diffusion of PLL chains into the interior of the PEM when in contact with the PLL solution.28 Upon immersion in a HA solution, the PLL chains diffused to the surface and interacted with the highly charged HA chains, resulting in large PE aggregates at the surface resulting in an exponential increase in PEM thickness as subsequent bilayers were deposited.28 Researchers have exploited the two different modes of growth within the same PEM to create distinct regions. For example, the combination of linear and exponential growth modes can be used to create compartments where the PEs diffuse freely within an exponentially growing region of bilayers and tightly bound in areas of linear growth.68 Taken together, this methodology exhibits enormous potential for applications in drug delivery, specifically as drug reservoir compartments.68 Multicompartment PEMs will be discussed later in this review.
The height of a PEM is also dependent on the pH of the individual PE solutions and their ionic strength.4,5,12,29,55 Polymers can be assembled that exhibit thick, loopy structures at lower charge density or thin, rod-like confirmations at a higher charge density.8 Mendelsohn et al. tuned the adhesion of murine NR6WT fibroblasts on PEMs comprised of PAH and PAA or PDAC and SPS by modulating the pH of the PE solution.29 PEMs comprised of PAH and PAA, assembled at pH 6.0, and exhibited excellent cell adhesion compared to PEMs assembled at pH 2.0. When PEMs were assembled at higher values of pH, the multilayers were more rigid in contrast to films assembled at pH 2.0. The observed decrease in cell adhesion was attributed to a higher extent of hydration.29 Thompson et al. assembled identical PEMs to directly relate stiffness, which was modulated by pH, to microvascular endothelial cell adhesion.55 Multilayers of PAH and PAA assembled at pH 6.5 exhibited a Young's modulus of 150MPa. On day 6 in culture, the cell densities were found to be ~100 cells/mm2.55 However, when the multilayer was assembled at pH 2.0, the Young's modulus was found to be 200kPa and the cell density on day 6 was ~20 cells/mm2. These trends suggested that less rigid multilayers may result in lower rates of proliferation.55 In separate studies, similar properties were obtained by simply changing the ionic concentration of the PE solutions.7,12 Typically PEMs are assembled under specific conditions, for example, pH or ionic concentration, to obtain a desired set of properties. However, post-assembly modifications may result in changes in the PEM thickness, stiffness, or porosity.10,29
Postassembly chemical crosslinking is a method to further modulate mechanical properties. For PEMs with the appropriate functional chemistry, such as amine and/or carboxyl groups, water-based covalent chemical crosslink chemistries have been investigated.16,32,69–71 Carbodiimide crosslinkers such as 1-ethyl-3-[3-dimethylaminopropyl] carbodiimide hydrochloride (commonly referred to as EDC) are crosslinkers used to couple carboxyl and amine groups leading to the formation of amide bonds.70,72–74 In a recent study, the Young's modulus was varied for PLL/HA PEMs to modulate myoblast adhesion, differentiation, and myotube formation using EDC concentrations ranging from 5 to 100mg/mL.57 Increasing the concentration of EDC was found to be directly proportional to the mechanical stiffness of the PEM. Myoblast adhesion increased with higher crosslinker concentration and well-defined focal adhesions were observed only for the stiffest PEMs along with well-organized f-actin fibers. The morphology of the myotubes varied significantly with thinner, more elongated tubes forming only on the stiffest PEMs (Young's modulus of 400kPa).57 It is also possible to crosslink PEMs during LbL assembly as demonstrated by Croll et al. using EDC.75 In this study, EDC and N-hydroxysuccinimide (NHS) were added to the HA solution before LbL assembly. PEMs cross-linked by EDC did not degrade in phosphate-buffered saline, whereas those assembled without the crosslinker rapidly degraded. These stable hydrated PEMs were used to prevent protein and cellular adhesion.75
Glutaraldehyde (GA) is another common chemical used to crosslink PEMs.76–78 Tong et al. crosslinked PEM microcapsules comprised of PAH and SPS to promote stability for potential applications in drug delivery. GA crosslinking resulted in a 2.3-fold increase in Young's modulus, from 290 to 680MPa.79 The critical pressure, upon which microcapsule collapse occurred, doubled upon GA crosslinking for 1h.79 More recently, GA crosslinking was utilized to improve the stability as well as to tune the Young's modulus of free-standing HA and chitosan (CHI) PEMs.18 Upon exposure to GA, the stability of the films in aqueous solutions increased to >90% weight retention over a period of 7 days. However, uncrosslinked films deteriorated in aqueous solutions within 5min.18 The Young's modulus increased from 90MPa for uncrosslinked PEMs to 310 and 480MPa for PEMs crosslinked with GA exposure for 1 and 2min respectively.18 BALB/c 3T3 fibroblasts colonized the entire surface of the crosslinked PEMs within 6 days and demonstrated well-defined actin cytoskeletal structure.18
In addition to chemical crosslinkers, other covalent methods can also be used. Recently, Moussallem et al. thermally crosslinked PAH and PAA PEMs to form covalent amide crosslinks.20 The hydrated Young's modulus increased from 6MPa to 8 GPa for uncrosslinked and crosslinked PEMs, respectively. A7r5 rat aortic SMCs cultured on uncrosslinked PEMs (6MPa) expressed phenotypic markers, such as vimentin and nonmuscle myosin heavy chain IIB,20 whereas SMCs expressed markers of a contractile phenotype such as calponin and smooth muscle α-actin when cultured on the 8 GPa crosslinked PEMs.20 For most cell types, increased substrate stiffness promotes cellular adhesion.16,29,38,53,55,56,62,69 The range of mechanical properties in PEMs due to assembly conditions can be used to probe cellular mechanics as well as promote cellular functions for tissue engineering (Table 1).
The surface composition of a PEM film can drastically impact cellular adhesion.61 This feature can be altered by varying the final PE deposited during the LbL assembly.34,55,56,80 Richert et al. observed that PEMs terminating in PLL demonstrated higher human chondrosarcoma adhesion forces, ~80nN, in comparison to PLGA-terminated PEMs, where an eightfold lower adhesion force was observed.56 More recently, Wittmer et al. assembled PLL and dextran sulfate PEMs.80 In situ mass adsorption measurements with a quartz crystal microbalance with dissipation showed a 40% increase in fibronectin adsorption on PLL-terminated PEMs compared to dextran sulfate-terminated multilayers. Further, significant spreading of human umbilical vein endothelial cells was observed only on the PLL-terminated PEM, due to the increased fibronectin adsorption and terminal PLL layer.80 However, certain polycations have displayed cytotoxic effects. For example, Brunot et al. determined the cytotoxicity of poly (ethyleneimine) (PEI) used as a base layer for PEMs and as a PE coating.81 PEI demonstrated potential cytotoxicity for both human periodontal fibroblasts as well as MG63 osteoblast-like cells.81 Additionally, Fischer et al. investigated the cytotoxicity of polycation solutions such as PEI, PLL, and PDAC.82 The MTT assays for L929 mouse fibroblasts indicated that PEI was more cytotoxic compared to PLL and PDAC.82 Together, these results demonstrate that choosing a noncytotoxic polyion is a critical design parameter when assembling PEMs for biological applications.
Surface composition does not always significantly impact cellular adhesion. Thompson et al. observed that terminal layer had no statistical significance when PEMs were sufficiently stiff to support cellular adhesion.55 When compliant PAH and PAA multilayers with an approximate Young's modulus of 200kPa were used as substrates for microvascular endothelial cells, the terminal layer (comprised of either PAA or PAH) had no statistically significant effect on cell density on day 7.55 Similarly, chondrosarcoma adhesion on crosslinked PLL- and HA-terminated PEMs did not vary, with both substrates almost entirely colonized by day 6.83 These studies suggest that there is a subtle interplay between cellular adhesion and chemical and physical properties of PEMs.
Topographical cues play an important role in vivo, since cells are exposed to environments ranging from the micro- to the nanoscale.61,63 The ability to control PEM surface topography is extremely important in tissue engineering. Basement membranes in vivo typically exhibit porosities ranging from the nano- to the microscale to enable the diffusion of gases and nutrients to the cells.84–86 When PAA and PAH multilayers were exposed to an acidic aqueous solution (pH 2.5), microporous PEMs were obtained (Fig. 2). The transition from a nonporous to porous state was attributed to polymer cleavage and reorganization.10 The characteristic pore length was 100–500nm after 60s exposure to an acidic solution (pH 2.4). Upon treatment at neutral pH for >1h, pore size diminished to ~50–200nm. By heating PEMs above 200°C, amidization occurred and created inter- and intrachain crosslinks, and the films were found to be stable under ambient conditions up to 18 months.10 In another study, Hiller et al. used a similar approach to introduce nanoscale pores in PAH/PAA films.11 By reducing the pH of the acidic solution to 1.8, pore sizes were reduced to 15–80nm.11 In addition to lowering the pH, salts such as NaCl and MgCl2 were added at low concentrations to promote structural rearrangements.11,87 Recently, nano- and microporous PEMs were used to mimic the basement membrane in the cornea (typical porosity ranging from 20 to 200nm).44,84,85 PAA and PAH multilayers exhibiting pore sizes ranging from 50 to 240nm were used to culture human corneal epithelial cells (HCECs).44 HCECs cultured on nanoporous PEMs exhibited two-fold higher migration speeds in comparison to microporous films, 30 and 15μm/h, respectively.44 In addition, HCECs exhibited well-defined actin cytoskeletal structures and vinculin focal adhesions only on the nanoporous substrates.44 These examples demonstrate that by simply varying the surface topography of a PEM, cellular response can be controlled.44,88
Maintenance of cellular phenotype is paramount in any tissue engineering application. The ability to measure in vitro cellular responses to conditions that recapitulate the environment in vivo can provide comprehensive insights into the behavior of implanted biomaterials. PEMs are often designed with finely tuned bulk properties like Young's modulus and degradability, in addition, such films can be further modified to maintain or induce specific phenotypic characteristics.89,90 In the following sections, the use of PEMs in the fields of regenerative medicine and stem cell differentiation will be highlighted.
The cellular adhesion and phenotypic properties of SaOS-2 (human osteoblast-like cells) and human periodontal ligament cells on multilayers derived from various PEs were investigated by Tryoen-Toth et al.54 In this study, anionic SPS and PLGA were utilized in combination with cationic PEI, PAH, and PLL to construct a wide range of PEMs. In these studies, the terminal PE layer drastically affected cellular adhesion, but more importantly, this feature had a significant effect on influencing cellular phenotype. Measurement of alkaline phosphatase mRNA levels indicated the terminal PE layer must be negatively charged to maintain osteoblast phenotype in both cell lines. Negatively charged PGA and SPS terminating films resulted in stable expression levels, unlike cationic PEI and PAH terminal films.54 Interestingly, it was also shown that these cells interacted only with the terminal PE layer in the substrate. For example, PGA deposited directly on top of PAH exhibited positive phenotypic effects by the osteoblast-like cells, whereas reversal of these two layers resulted in negative phenotypic expression.54 These results suggest the terminal PEM layer is a dominant factor in biocompatibility and must be carefully considered based on desired cellular interactions for a particular application. Many additional studies have investigated adhesion, proliferation, spreading, and phenotypic expression profiles of many different cell types on PEMs constructed from combinations of PEs (Table 2).
In addition to facilitating phenotypic maintenance, it is important for PEMs to stimulate cellular proliferation. Various growth factors can stimulate proliferation, and to this end their incorporation and controlled release from PEMs may be necessary for regenerative medicine. A common problem with growth factor administration is uncontrolled release leading to high initial concentrations followed by periods of rapid clearance, thus limiting possible therapeutic effects.91,92 The release of fibroblast growth factor 2 (FGF-2) has been controlled with the construction of poly (β-aminoester)s/heparin (HEP)/FGF-2/HEP tetralayers deposited on top of PEI/SPS multilayers.43 This approach was similar to another study using tetralayers for controlled hydrophobic drug and polysaccharide release.93 The incorporation of the poly (β-aminoester) polymer into the tetramer was to prevent the localized degradation of the PEM due to cellular interactions. Ester bonds in the poly (β-aminoester) were hydrolytically cleaved during cellular degradation.93 The incorporation of methylene units to increase hydrophobicity resulted in decreased degradation. The incorporation of HEP into tetralayers was necessary due to its known sequestration of FGF-2,94 thereby inhibiting the diffusion of free FGF-2 between tetralayers. Thus, the tetralayer method of encapsulating FGF-2 proved to be very useful since the protein from a single layer had to be completely released before release from the underlying layers. The bioactivity of the released protein was verified through its exposure to preosteoblasts and by monitoring their proliferation.43 Information on additional studies investigating the delivery of bioactive molecules using PEMs is provided in Table 2, further supporting the versatility of PEMs in stimulating phenotypic maintenance.
When short functional peptide sequences are tethered to a substrate, they can modulate a wide range of responses. For instance, the tri-peptide sequence, RGD, has been extensively utilized to promote cell adhesion and migration.95,96 This approach has been particularly successful for cells that preferentially adhere to fibronectin.19,97 A comprehensive review of the chemistry involved in RGD and other peptide conjugation to synthetic polymers has been previously been reported.98 To demonstrate the efficacy of PEMs functionalized with short peptide sequences, Tsai et al. constructed PAH/PAA PEMs comprised of five bilayers, with a terminal RGD-PAH layer.99 In this study, the incorporation of RGD stimulated osteoblast adhesion, proliferation, and phenotype.99 The assembly of the multilayer was conducted at pH of 2 and 6.5. At lower pH, the anionic PE was incorporated at increased nonstoichiometric ratios to compensate for the fully ionized cationic PE (~70% PAA, 30% PAH at pH 2.0).99,100 Consequently, when the terminal RGD-PAH layer was deposited, the concentration of RGD molecules at the surface increased fivefold at pH 2.0.99 Despite the increased RGD surface concentration, osteoblast adhesion and proliferation were unaffected since the concentration of RGD exceeded 0.6pmol/cm2.99 This value has been reported to be the threshold concentration, necessary to stimulate osteoblast adhesion and spreading.101 However, at pH 2.0 a threefold increase in the deposition of calcium by osteoblasts was observed. Although the exact mechanism of increased calcium deposition in this study remained unknown, the authors hypothesized that decreased PEM mechanical stiffness, 100MPa at pH=6.5–1MPa at pH=2.0,55 or RGD conformational changes due to pH changes may be responsible for the observed increased calcium deposition. This study serves as a good example on the ability to elicit a desired phenotypic response by simply varying the assembly conditions of a PEM.
Many studies have focused on maintaining cellular phenotype through direct cellular interaction with PEMs. The adhesion of hepatocytes and maintaining their phenotype is one such example.96 In general, primary hepatocytes cultured in vitro as monolayers tend to rapidly de-differentiate. They lose their ability to produce urea (indicative of liver specific carbohydrate, lipid, and amino acid metabolism) and albumin (representing hepatic protein production). PAA/PAH multilayers have been used to culture hepatocytes and these substrates appear to elicit increased urea and albumin production in comparison to hepatocytes cultured on a collagen surface.102 Similarly, Wittmer et al. have demonstrated that both PLL/ALG and PLL/PLGA PEMs facilitate adult rat hepatocyte adhesion and increased albumin production when compared to cells cultured on a collagen surface.103 Further, in both reports, the adsorption of collagen on the terminal PE layer enhanced hepatocyte adhesion and phenotypic functions.102,103 Hepatocytes are generally stable when cultured between two collagen gels (collagen sandwich).104,105 When these cells are cocultured with nonparenchymal hepatic cells, their function can be further enhanced.48,106–108 Studies by Kidambi et al. have shown that hepatocyte and fibroblast attachment can be spatially controlled on patterned PDAC/SPS PEMs with superior performance.47,48
More recently, three-dimensional (3D), layered liver mimics have been designed using PEMs.26,27 In these 3D tissue mimics, hepatocytes were seeded on a collagen gel and a HA-CHI PEM was assembled above live cells. Next, liver sinusoidal endothelial cells (LSECs) were seeded above the PEM-coated hepatocytes.27 In these cultures, the PEM served as a synthetic substitute for the Space of Disse, the interface between hepatocytes and LSECs in vivo, thus using the PEM to create a physiologically relevant 3D culture system. In these cultures, the physical properties of the PEM, the number of layers, and the relative concentrations of each cell type affected hepatocyte-specific functions.27 The shear modulus of the PEM within the 3D liver mimic was ~100kPa, a value similar to that previously shown to maintain hepatic phenotype.109 Furthermore, the activity of CYP1A1/2, a cytochrome P450 enzyme involved in the metabolism of cyclic aromatic compounds was increased 16-fold over conventional monolayer cultures.27 These results suggest that PEMs can serve as the interface between cell layers and used to assemble a wide range of stratified 3D tissue mimics.
A growing body of work is emerging in which PEMs are being tuned to direct phenotypic differentiation, particularly progenitor cells, toward specific lineages. Directing the lineage of stem cells is dependent upon complex interactions between soluble factors, the extracellular matrix, intra- and intercellular communications, and biophysical stimuli.110–113 Although soluble growth factors are critical to induce directed stem cell differentiation, Engler et al. recently showed that mesenchymal stem cells (MSCs) could be directed to either neuronal, muscle, or osteogenic lineages by simply modulating substrate elasticity from <1, to 8–17, to 25–40kPa, respectively.53 Since PEMs exhibit mechanical properties in this range, the results from this study can be used in directing phenotype.16,17,57 PEMs composed of alternating layers of PAH and SPS have been shown to stimulate the differentiation of endothelial progenitor cells.114–116 Endothelial progenitor cells are a specialized type of progenitor cells with a limited number of potential lineages. However, when these cells were cultured on PEMs in the presence of culture medium containing growth factors, such as vascular endothelial growth factor and basic FGF, the differentiation into endothelial phenotype was observed.114,117 PEMs derived from PEI and SPS with embedded carbon nanotubes enabled neuronal progenitor cell differentiation into the three main classes of neural cells.118 These studies show PEMs can aid in directing the differentiation of progenitor cells.
Liu et al. studied the adhesion and differentiation of MSCs when cultured on PEMs comprising of alternating layers of CHI with either gelatin (GEL), HA, or HEP.119 In this study, MSC adhesion increased for the following PEM combinations: GEL/CHI, HA/CHI, and HEP/CHI.119 It was further reported that MSC spreading was inversely proportional to cell adhesion. Although cell spreading could be modulated through the careful selection of PEs, the authors concluded a high degree of cell spreading, observed on GEL/CHI, was not indicative of MSC phenotype. Studies by Semenov et al. further validated these findings by culturing MSCs on PEMs exhibiting a range of mechanical properties showing chondrogenic and osteogenic lineage differentiation was dependent on the Young's modulus.6
One of the most critical aspects of stem cell research is selecting and isolating a pure population of cells. Fluorescence-activated cell sorting is a powerful technique used in the isolation of populations of pure cells. However, due to the inherent complexity of this technique, and in obtaining viable, antibody-free cells, this technique presents limitations.120–122 A very interesting application of PEMs was reported by Tsai et al. where fetal liver stem/progenitor cells could be selected from a gross tissue digest using PLGA-terminating PLL/PLGA multilayers. Further, these substrates stimulated cell proliferation over time.120 The development of such a cell selection method could yield a more precise and less expensive approach to selectively isolate pure cell populations.
Another novel use of PEMs in cell culture was to culture MSCs on thermo-responsive multilayers made from SPS and PAH, copolymerized with N-isopropylacrylamide.123 The use of thermo-responsive PEMs enabled MSC proliferation and passaging without enzymatic treatment. This procedure yielded an increased number of MSCs capable of osteogenic and adipogenic lineage differentiation after multiple passages.123 The combination of these two studies illustrates the application of PEMs for the selection and culture of stem cells in a noninvasive manner.
Although advances have been made toward the incorporation of bioactive molecules into PEMs, an unmet challenge continues to be the inclusion of small, uncharged, hydrophobic molecules.49,124,125 Such compounds account for ~40% of the drugs approved by the Food and Drug Administration.41 A recent development has been the construction of free-standing multilayer films comprised of block copolymer micelles and PAA.41 Kim et al. demonstrated that the hydrophobic antibacterial drug triclosan,41 remained active after micelle incorporation and could be successfully released at physiological conditions. It was also shown that the rate of film degradation and drug release could be tailored based through changes in the degree of crosslinking within the PEM. A similar approach to protect and deliver bioactive molecules is their incorporation within polymersomes.126–128 Polymersomes are bilayer membrane capsules formed through the self-assembly of ampiphilic block copolymers.129 Biodegradable functionalized polymersomes exhibit potential for site-specific therapeutic delivery through incorporation into PEMs. Further expanding on this idea of drug encapsulation is the construction of capsosomes. Capsosomes are constructed from alternating layers of liposomes and an alternately charged polymer using LbL assembly.130,131 The multiple liposome layers allow multiple reactions to occur simultaneously while being spatially separated within a single capsosome.130
Another novel application of PEMs in drug delivery is the construction of multiple drug reservoirs within a single degradable PEM. Drug reservoirs are constructed by exploiting the characteristic diffusivity associated only with exponentially growing PEMs, not normally observed under linear growth.68 The physical differences between exponentially and linearly grown PEM films have already been discussed in detail in the section Assembly conditions modulate the mechanical properties of PEMs. Drug reservoirs take advantage of exponentially grown PEMs in which bioactive molecules are incorporated. Due to the higher diffusivity associated with such regions, bioactive molecules can easily diffuse, thereby providing a means for localized delivery at high concentration.13 A schematic of the layered architecture required to create a PEM film containing multiple drug reservoirs is shown in Figure 3. The construction of such reservoirs requires the presence of impermeable barriers on either side of the exponentially grown PEMs. Impermeable barriers to prevent drug diffusion include linearly grown films,65 an annealed layer of charged wax particles,132 or a layer of biodegradable PLGA.68 Garza et al. created two drug reservoirs within a single film in which each reservoir was separated by a layer of PLGA. In this instance the PLGA layer was deposited above the terminal PEM layer using a novel spray deposition technique to provide the biocompatible biodegradable layer separating the reservoirs.68 The PLGA layers prevented direct cell contact with the drugs in each reservoirs up to 5 days, a length of time which correlated directly with time the cells took to degrade the PEM.68 It has also been shown that cells do not have to come in direct contact with the reservoirs, but can internalize drugs through localized degradation of the PLGA layer.133,134
A final application of PEMs to be discussed in tissue engineering involves living cells functionalized with multilayer patches.50 This work represents an elegant combination of several PEM applications and techniques outlined in this review. Recent advances in PEM applications have led to the construction of multilayer assemblies encapsulating entire living cells135–139; however, the entire surface chemistry of the cell is likely altered affecting cell communication and environmental interactions.50 In the work by Swiston et al. a three-tiered construct was developed where each PEM tier had a specific function: the releasable region, the payload region, and the cell-adhesive region.50 The three-tiered construct was built from a substrate on which a photolithographic patterning method was used to create the first tier, a layer capable of being detached from the substrate through a combination of pH and thermal conditions. The second tier, the payload region, of the cell patch was engineered as a drug reservoir to contain a range of bioactive molecules. In this case, the payload PEM consisted of alternating fluorescent-PAH and anionic super-paramagnetic nanoparticles. The final tier in the patch was the cell-adhesive region constructed from a HA/CHI PEM. After patch construction, individual cell-patch attachment was observed only for 10μm patches and cell aggregates were observed on 15μm patches.50 After release from the substrate, these cells could be magnetically manipulated due to the presence of the PEM patch. A significant consequence was that the patches did not interfere with the phenotype and physiology of the cell.
The tunability of PEMs provides unique opportunities to mimic the complex in vivo extracellular matrix environment. Chemical, mechanical, and topographical features can be varied to promote cellular adhesion, differentiation, migration, and gene expression of virtually any cell type. Currently, the majority of PEM applications in tissue engineering are focused on rendering two-dimensional surfaces compatible for biological applications. However, it is well known that providing a 3D environment to cells has the potential to elicit an optimal response.61 Modifications in the current techniques used to deposit multilayers on two-dimensional substrates will be critical to successfully incorporate PEMs into the design of 3D scaffolds.
An ultimate objective in tissue engineering is to functionalize implantable materials directing cellular adhesion, phenotype, and enhance site-specific regeneration of damaged or diseased tissues. One such instance is the construction of 3D tissue constructs involving multiple cell types. To this point, most PEM-related tissue engineering applications have focused on surface coatings. However, due to the tunable nanoarchitecture of PEMs they can be used in the construction of 3D liver mimics.27 The PEM must provide a mechanical support for LSEC culture, a physical barrier between hepatocytes and LSECs, but most importantly allow diffusion of signaling molecules to facilitate cell–cell communication and phenotypic maintenance of both cell types. Recent efforts to create free standing PEMs comprised of synthetic and biologically derived PEs18,140,141 exhibit tremendous potential to be utilized as scaffolds in tissue engineering. An application with much potential in biological therapeutic applications is the coating of individual cells with biocompatible PEMs. Use of PEMs to coat a cell will likely alter its biological activity, but a carefully chosen combination of PEs may result in optimal physiological function.
Finally, the number of tissue engineering application employing PEMs are rapidly expanding and providing useful insights into cellular responses to different combinations of mechanical and chemical signals. In the near future, PEM applications studied in vitro for drug delivery and directing cellular phenotype will undoubtedly proceed to in vivo applications and clinical trials. Current research and emerging ideas to functionalize surfaces for biological applications will benefit from the unique and highly tunable nature of PEMs as they can be deposited on any surface with a wide range of mechanical and chemical properties.
We gratefully acknowledge financial support from the National Science Foundation (DMR-090750, P.R.), the National Institutes of Health (NIDDK-1R21DK077802, P.R.), the Institute for Critical Technology and Applied Sciences (P.R.), and the MILES-NSF IGERT program at Virginia Polytechnic Institute and State University (A.L.L. and P.R.).
No competing financial interests exist.