Interpenetrating polymer network hydrogels as bioactive scaffolds for tissue engineering

Abstract

Tissue engineering, after decades of exciting progress and monumental breakthroughs, has yet to make a significant impact on patient health. It has become apparent that a dearth of biomaterial scaffolds that possess the material properties of human tissue while remaining bioactive and cytocompatible has been partly responsible for this lack of clinical translation. Herein, we propose the development of interpenetrating polymer network hydrogels as materials that can provide cells with an adhesive extracellular matrix-like 3D microenvironment while possessing the mechanical integrity to withstand physiological forces. These hydrogels can be synthesized from biologically-derived or synthetic polymers, the former polymer offering preservation of adhesion, degradability, and microstructure and the latter polymer offering tunability and superior mechanical properties. We review critical advances in the enhancement of mechanical strength, substrate-scale stiffness, electrical conductivity, and degradation in IPN hydrogels intended as bioactive scaffolds in the past five years. We also highlight the exciting incorporation of IPN hydrogels into state-of-the-art tissue engineering technologies, such as organ-on-a-chip and bioprinting platforms. These materials will be critical in the engineering of functional tissue for transplant, disease modeling, and drug screening.

Introduction

The holy grail of tissue engineering is the construction of functional tissues and organs for transplant, disease modeling, and drug screening applications. Despite promising preclinical studies in animal models and preliminary clinical trials, tissue-engineered constructs have yet to make a measurable impact on patient outcomes (Geris and Papantoniou, 2019; Hoffman et al., 2019). One outstanding challenge that has prevented the clinical application of these cell-seeded scaffolds is the dearth of biomaterials that recapitulate the complexity of human tissue (Chaudhuri, 2017; Khademhosseini and Langer, 2016). Human tissue consists of dozens of hierarchical biological polymers that simultaneously contribute to mechanical stabilization and bioactivity in a fluid-dominated complex environment. In contrast, many pioneering studies in tissue engineering were initially conducted on two-dimensional (2D) surfaces (Baker and Chen, 2012). Perhaps not surprisingly, it has now been reliably demonstrated that cells display aberrant behavior on 2D substrates, including, but not limited to, flattened cell shape, an abnormal distribution of integrin-adhesion ligand complexes, and super-physiologic stiffness (Caliari and Burdick, 2016). Therefore, tissue scaffolds should (1) closely mimic the cues of the native extracellular matrix (ECM) and (2) encapsulate cells in three dimensions (3D). 3D hydrophilic cross-linked polymer networks, i.e., hydrogels, fulfill both of these critical criteria by promoting cell viability, ensuring nutrient transport, and recapitulating the bioactivity and microarchitecture of the native ECM (Caliari and Burdick, 2016; Crosby and Zoldan, 2019; Huang et al., 2017).

Bioactive hydrogels consist of a single biologically-derived polymer (isolated from a living organism), a single synthetic polymer (synthesized ex vivo), or a composite of two or more polymers. Biologically-derived polymers such as collagen, fibrin, and hyaluronic acid are excellent mimics of the ECM because they contain cell-binding and degradable peptide motifs and can self-assemble into micro-scale fibers that are vital for cell migration and environmental sensing. Biologically-derived polymers such as collagen rely on physical cross-links (e.g., molecular entanglement and non-covalent forces) that endow the hydrogels with physiological viscoelasticity while simultaneously weakening them, as physical interactions tend to be considerably weaker than chemical cross-links (Spicer, 2020). In contrast, synthetic polymers such as polyethylene glycol (PEG), poly(lactic acid) (PLA), and polycaprolactone (PCL) are engineered to display ideal mechanical properties and to incorporate cell-adhesive and degradable peptides. However, it has proven challenging to mimic the fibrillar architecture of the ECM in synthetic hydrogels, and many cross-linked synthetic polymers do not recapitulate the viscoelastic properties of human tissue (Lou et al., 2018; Matera et al., 2019). Furthermore, most biologically-derived and synthetic single-polymer hydrogels for tissue engineering applications tend to be soft, brittle, and weak (Gong, 2010). Therefore, the creation of composite hydrogels can overcome these limitations by incorporating one polymer to stimulate bioactivity and another to enhance the mechanical properties or degradability of the construct.

Composite hydrogels consisting of two polymers can be classified as (a) polymer blends (b) graft copolymers (c) block copolymers (d) AB-graft copolymers (e) semi-interpenetrating polymer networks (sIPNs) or (f) full IPNs (Figure 1A) (Sperling, 1994). In this review, we will focus on IPN hydrogels because they remarkably improve the strength and ductility of single-polymer hydrogels and offer tunability in biological settings. Briefly, IPNs are commonly defined as a class of polymer composites composed of two cross-linked networks that are topologically entangled and yet cannot be separated without disrupting existing chemical bonds (Jenkins et al., 1996; Khan et al., 2020; Suthar et al., 1996). Additionally, IPNs do not dissolve when immersed in a solvent, are resistant to creep and flow, and exhibit differential material properties than their constituent monomers or polymers (Gupta and Srivastava, 1994). IPNs can be synthesized from monomers or linear polymers simultaneously or sequentially (Figure 1B). Sequential IPNs result from the swelling of a monomer/linear polymer and cross-linker into an already polymerized single-polymer network or by selectively cross-linking one network before the other (e.g., ultraviolet radiation) (Dragan, 2014; Myung et al., 2008). When IPNs consist of hydrophilic polymers, they swell and form IPN hydrogels.

Figure 1:

Introduction to IPN chemistry and applications (A) Possible physical and chemical interactions between two linear homopolymers in solution. Briefly, polymer blends are a physical mixture of the two polymers, graft copolymers are defined by the binding of one or more side chains of one polymer to the backbone of the other polymer, block copolymers are defined by the end-to-end binding of two different polymers, AB graft copolymers are cross-linked networks of two or more species of homopolymers, sIPNs contain a linear polymer physically entrapped by a cross-linked network of a different polymer, and IPNs are defined by the intertwining of two polymer networks. Reprinted (adapted) with permission from (Sperling, 1994). Copyright (1994) American Chemical Society. (B) IPNs can be (1) cross-linked simultaneously, (2) synthesized sequentially by selective cross-linking, or (3) synthesized by adding a linear polymer of monomers to an existing single-polymer hydrogel. (C) The number of manuscripts linking IPNs to biomedical applications from 1988–2019, demonstrating an approximately exponential increase (data from pubmed.gov, PubMed database keyword search: interpenetrating polymer network). (D) Applications of IPN hydrogels to tissue engineering constitute approximately 23% of the manuscripts cataloged in (C).

The first naming and thorough characterization of an IPN is generally credited to J.R. Millar, who synthesized an IPN consisting of styrene and divinylbenzene in 1960 (Millar, 1960). However, IPNs were not evaluated for biomedical applications until the late 1980s. Even then, IPNs were viewed mainly as potential acellular connective tissue replacements (e.g., cartilage). Not until 1996 was the biocompatibility of an IPN (poly(2-hydroxyethyl methacrylate)-gelatin) hydrogel evaluated in vitro (Peluso et al., 1996), and cells were encapsulated in a 3D IPN microenvironment 12 years later (Liu and Chan-Park, 2009). However, the 2010s have witnessed an exponential rise in the number of manuscripts detailing the use of IPNs in biomedical applications (Figure 1C); of these publications, most (~38%) identify IPNs as drug delivery vehicles, structural biomaterials (~26%), or tissue engineering scaffolds (~23%) (Figure 1D). In two critical reviews published in 2013 and 2014, Matricardi et al. and Dragan, respectively, identified the potential of IPN hydrogels in tissue engineering (Dragan, 2014; Matricardi et al., 2013). However, tissue engineering applications were not covered in-depth. In this review, we focus on the translation of IPN hydrogels to tissue engineering applications within the past five years. Specifically, we detail how IPNs can improve hydrogel properties that are critical for their function as tissue scaffold syntheses, such as mechanical strength, bioactivity, degradation, and swelling. Additionally, we highlight the potential of IPN hydrogels to be deployed in 3D bioprinting and organ-on-a-chip applications. Lastly, we discuss future challenges in the IPN hydrogel-based tissue engineering field and evaluate which grand challenges outlined by previous reviews were addressed and which remain unsolved problems.

IPN hydrogels tissue scaffolds derived from biological polymers

Tissue engineers have long sought to mimic the hydrophilicity and hierarchical architecture of native tissue. Together, these two properties permit large-scale molecular rearrangements of the constituent bioactive polymers, which leads to the emergence of unique material properties (Fan and Gong, 2020). As such, a commonly-employed strategy is to isolate polymers from human, animal, or insect tissue, modify their structure, and re-construct ECM-mimicking hydrogels. A host of biologically-derived polymers have been used to construct tissue-engineered scaffolds, such as collagen, fibrinogen, hyaluronic acid (HA), silk, chitosan, starch, and alginate. The advantages and drawbacks of hydrogels synthesized from these polypeptides or polysaccharides have been well-summarized in a recently published review (Spicer, 2020). Biologically-derived polypeptides typically contain cell-binding and degradable motifs that add necessary bioactivity to the final hydrogel. Polysaccharides such as HA can be modified through a series of well-established chemical reactions to independently tune the resulting stiffness, degradability, viscoelasticity, or bioactivity of the hydrogel (Highley et al., 2016). Furthermore, biologically-derived hydrogels can be synthesized with minimal chemical modification, which ensures that the final product remains biocompatible. However, single-polymer biologically-derived hydrogels typically gel via physically cross-linking mechanisms and therefore tend to be mechanically weak and exhibit significant batch-to-batch variability (Yan and Pochan, 2010). The synthesis of an IPN hydrogel from two biologically-derived polymers, hence “biological-biological IPNs,” can enhance mechanical strength and modulate cell and tissue response by varying the stiffness and load-bearing capacity of the IPN hydrogel.

Enhancing the bulk mechanical strength of tissue-engineered scaffolds

While single-polymer biological hydrogels are biocompatible and facilitate cell adhesion, they possess low tensile or compressive strength (Gong, 2010). In contrast, biological-biological IPN hydrogels can possess storage/compressive moduli orders of magnitude greater than their constituent components. For example, a chitosan-gelatin methacrylate (GelMA) IPN hydrogel exhibited compressive moduli 33 times and 3 times greater than corresponding single polymer hydrogels, respectively (Suo et al., 2018). To form this IPN, the authors photo-crosslinked GelMA and then immersed the resulting sIPN in 0.1 M NaOH to repulse -NH₃⁺, which encouraged the otherwise positively charged chitosan to cross-link via hydrophobic and hydrogen bonds. The authors attributed this dramatic increase in compressive strength to the increased entanglement of chitosan with GelMA, which created a more porous mesh structure with thicker walls. GelMA was also photopolymerized and physically entangled with silk fibroin to manufacture load-bearing soft tissue (such as cartilage) (Xiao et al., 2019). The maximum compressive modulus of the silk fibroin-GelMA IPN hydrogel was measured to be 600 kPa, which lies within the physiological range of biological cartilage (Liu et al., 2015). However, the methods listed above either involve highly cytotoxic reagents (NaOH) or the modification of a biological protein, which is costly and may degrade its function when implanted in vivo.

In contrast, the enzymatic cross-linking of biological-biological IPN hydrogels can enhance the strength of tissue engineering scaffolds while avoiding the cytotoxicity that may be induced with physical or chemical cross-linking (Chirila et al., 2017). Park et al. synthesized a gelatin-silk fibroin hydrogel to cross-link gelatin with microbial transglutaminase (mTG), thereby avoiding the chemical modification of gelatin required by the synthesis of GelMA (Park et al., 2019). They were able to synthesize a construct with a compressive modulus of up to 1.5 MPa, which is of sufficient strength to be used in bone and cartilage tissue engineering applications. Fan et al. also used mTG to cross-link gelatin while also adding horseradish peroxidase (HRP) to cross-link HA in a gelatin-HA IPN hydrogel (Fan et al., 2015). This IPN hydrogel promoted cell spreading and proliferation and increased the compressive modulus by a factor of 10 when compared to pure HA hydrogels. Similarly, a fibrin-silk fibroin IPN hydrogel synthesized through a dual enzymatic cross-linking method with thrombin and HRP promoted cell proliferation and exhibited a storage modulus 10 and 12 times greater than that of single-polymer fibrin and silk fibroin hydrogels, respectively (Goczkowski et al., 2019).

The stiffness of IPN hydrogels may also be modulated to be more physiologically relevant and approximate the stiffness of native tissue. Since IPNs are independently cross-linked and are not chemically bound, it is possible to fine-tune the bulk stiffness of the hydrogel by separately varying the concentration of each component. For example, the stiffness of a sericin-alginate hydrogel was lowered from ~30 kPa to ~2 kPa by adjusting the ratio of sericin to alginate in the hydrogel (Zhang et al., 2015). The natural bioactive properties of sericin also promoted the adhesion, proliferation, and migration of mouse myoblasts (C2C12), which indicated that this polymer could be incorporated in future tissue engineering scaffolds to enhance cell viability. The addition of nanoparticles can also lend mechanical stability to IPN hydrogels that exhibit physiological porosity and viscoelastic properties. Cellulose nanocrystals were added to a sodium alginate-gelatin IPN to mimic the mechanical properties of cartilage (Figure 2A) (Naseri et al., 2016). The resulting sIPN hydrogel exhibited viscoelastic behavior under compressive tests and comparable stress and strain values to biological cartilage.

Figure 2:

Biological-biological IPN hydrogels in tissue engineering (A) The doping of cellulose nanocrystals into sodium alginate-gelatin 3D IPN hydrogels increases the mechanical integrity of the biological-biological constructs and makes these scaffolds ideal scaffolds for cartilage tissue regeneration. Reprinted (adapted) with permission from (Naseri et al., 2016). Copyright (2016) American Chemical Society. (B) In 3D alginate-collagen IPN hydrogels, fibroblasts display an extended morphology in softer hydrogels; in contrast, an increase in stiffness encourages fibroblasts to maintain a more rounded phenotype while increasing the production of inflammation-associated genes and proteins. Reprinted (adapted) with permission from (Branco da Cunha et al., 2014). Scale bars are 100 μm. (C) Increasing the stiffness of alginate-collagen IPN hydrogels results in an increased metastatic response, as multicellular aggregates were more elliptical and significantly larger after 17 days in 3D culture. Scale bars are 70 μm.

The synthesis of IPN scaffolds with enhanced mechanical properties has been applied to intervertebral disc regeneration. The intervertebral disc (IVD) is a highly hydrated tissue composed predominantly of collagen and proteoglycan and experiences loads ranging from 0.1–1 MPa (Neidlinger-Wilke et al., 2014). A triple IPN, composed of dextran, chitosan, and teleostean (DCT), was shown to remain within the disc space without structural damage and promote mesenchymal stem cell (MSC) viability and proliferation (Gullbrand et al., 2017). Following in vivo studies revealed that the IPN remained within the disc space for two weeks of load-bearing activity. Similarly, an injectable collagen-low molecular weight hyaluronic acid (collagen-LMW HA) sIPN was created to mimic the rheological characteristics of the nucleus pulposus of the IVD (Tsaryk et al., 2015). The nucleus pulposus, the inner part of the IVD, is an avascular, highly hydrated structure composed of mostly ECM and water with just 1–3% of the structure’s total volume occupied by cells (Erwin and Hood, 2014). The swelling and compression of the IVD are mainly responsible for force dispersion in various motions of the spine (Cramer, 2014). The viscosity of this collagen-LMW HA sIPN decreased substantially with increasing shear rate, thereby implicating shear-thinning behavior and highlighting this novel material as an ideal candidate for IVD injection (Portnov et al., 2017). More importantly, this IPN hydrogel supported MSC growth and chondrogenic differentiation both in vitro and in vivo. In summary, IPNs have significantly enhanced the mechanical strength of biologically-derived hydrogels. Specifically, the introduction and refinement of enzymatic cross-linking methods have allowed for the development of cell-friendly cross-linking reactions. Furthermore, these mechanical enhancements in IPN hydrogels have been rigorously tested in animal models for IVD regeneration. This area is promising for future tissue engineering advances because this tissue does not require a relatively large number of cells to be physiologically functional.

Tuning stiffness to modulate cellular response

In the previous section, we reviewed different approaches to improve the bulk mechanical properties of biological-biological IPNs to match the mechanical properties of native tissue. Alternatively, IPN stiffness has been modulated to monitor the response of encapsulated cells to physical forces (i.e., mechanotransduction) and the microarchitecture of their local microenvironment. Of particular interest, substrate stiffness has been shown to influence a variety of cell behaviors in 2D and 3D (Engler et al., 2006). Here, we will focus on the modulation of IPN stiffness to regulate cell morphology, proliferation, and differentiation.

From fibroblast elongation to platelet activation, the connection between cell morphology and cell function has been well-established. Therefore, by carefully engineering the stiffness of the local microenvironment, tissue engineers can instruct tissue-scale behavior by regulating cell morphology via mechanotransduction. The stiffness of alginate can be tuned by varying the amount of calcium ion in the cross-linking solution, and thus this biological polymer is an attractive candidate for exploring the impact of substrate stiffness on cell behavior. For example, calcium ion concentration was varied to study the relationship between cell morphology, stiffness, and inflammation in alginate-collagen IPN hydrogels (Figure 2B) (Branco da Cunha et al., 2014). In these 3D microenvironments, fibroblasts exhibited spindle-like morphologies and elongated after a few hours within softer IPNs, while fibroblasts exhibited a more spherical morphology in stiffer IPN hydrogels. Increased IPN hydrogel stiffness simultaneously promoted a three-fold increase in the production and secretion of IL-10, a well-implicated inflammation-associated protein, thereby highlighting a possible link between matrix stiffness and wound healing potential. As we have previously outlined (Crosby and Zoldan, 2019), stiffness needs to be decoupled from bulk polymer density to extract the impact of substrate stiffness on cell behavior. More specifically, decoupling the effects of stiffness from density on vascular cell spreading and microvascular network topology is critical for the development of angiogenic biomaterials. To this end, Vorwald et al. systematically tuned the stiffness of fibrin-alginate IPN hydrogels by varying the concentration of thrombin and calcium chloride, which are needed to cross-link fibrin and alginate, respectively (Vorwald et al., 2020). Predictably, high-stiffness IPN hydrogels significantly reduced cell spreading. Low-stiffness IPN hydrogels containing more fibrin, in contrast, promoted network formation from endothelial cells that remained stable for one week. Therefore, it may be necessary to manufacture vascularized scaffolds that are stiff enough to be handled by surgeons but, at the cellular level, are soft enough to encourage vascular cell migration and tubulogenesis. Kavari et al. varied the alginate concentration in alginate-collagen IPN hydrogels to create a range of physiological stiffness (1.85 to 5.29 kPa) microenvironments to probe the impact of 3D substrate stiffness on the metastatic potential of multi-cellular aggregates (MACs) (Khavari et al., 2016). Stiffer IPNs induced a more rapid and pronounced spherical to ellipsoidal transition in MACs (Figure 2C). This spherical-to-ellipsoidal transition was shown to overlap with increased proliferation rate and larger cell bodies, indicators of increased metastatic potential. These studies imply that mechanical confinement may trigger an increase in cell proliferation.

In the human body, cells proliferate and differentiate when exposed to spatiotemporal biomechanical loading (Mammoto and Ingber, 2010). Furthermore, mechanotransduction also plays a vital role in organogenesis (Wozniak and Chen, 2009). While single-polymer biologically-derived hydrogels typically fracture when exposed to biological loading conditions, IPN hydrogels can be formulated to withstand these loads and thus enable the study of mechanotransduction pathways that elicit cell proliferation and differentiation. For example, alginate-only hydrogels permanently deformed while alginate-GelMA IPN hydrogels retained their original structure at 50% strain and repeated cyclic loading (20,000 cycles at 0.5 Hz of 10% strain) (Jeon et al., 2014). MSCs exposed to such cyclic loading proliferated and tended to differentiate towards an osteogenic lineage, as demonstrated by an increase in alkaline phosphatase activity. Similarly, increasing the stiffness of alginate-fibrin IPNs cross-linked with calcium silicate encouraged rat bone marrow stromal cells to favor osteogenic differentiation, as demonstrated by the upregulation of RUNX-2 and ALP osteogenic genes, and an increase in calcification, visualized by a Von Kossa staining (Zhang et al., 2002). In summary, since IPN polymers consist of two or more polymers, the stiffness of the bulk scaffold can be tuned to physiological parameters by systematically varying the properties/cross-linking mechanisms of one or both polymers and by changing the density of the constituent polymers.

IPN hydrogels derived from biological and synthetic polymers

The decision to incorporate a synthetic polymer into a biological hydrogel is typically driven by the desire to fine-tune the emergent material properties of the composite IPN hydrogel. Synthetic polymers, generally, are more hydrophobic, more resistant to deformation, and less degradable than their biological counterparts (Gyles et al., 2017). Therefore, the incorporation of synthetic polymers often leads to a dramatic increase in the strength and stability of biologically-derived hydrogels. For IPN hydrogels consisting of both biologically-derived and synthetic polymers (“biological-synthetic IPNs”), we have identified four main areas in which biological-synthetic IPNs are being used to improve matrix properties in tissue engineering: bulk mechanical properties, stiffness, electrical responsivity, and degradability.

Enhancing the bulk mechanical properties of tissue-engineered scaffolds

Synthetic polymers are commonly used to improve the mechanical strength of single-polymer biologically-derived hydrogels. The addition of a synthetic polymer can improve the bulk material properties of the scaffold without adversely affecting the bioactivity of the biological polymer (e.g., cell attachment). Specifically, synthetic polymers have been incorporated into biologically-derived polymer hydrogels to reinforce scaffolds used for dermal, bone, and cartilage tissue engineering.

Synthetic-biological IPN hydrogels have been synthesized as scaffolds for dermal tissue engineering. For example, Gsib et al. created a fibrin-polyethylene glycol dimethacrylate (PEGDM) IPN hydrogel as a suitable scaffold for skin healing (Gsib et al., 2018). Fibrin, a significant component of skin ECM (Gsib et al., 2017), was physically cross-linked, forming a weak biologically-derived single-polymer hydrogel. The addition of commercially available synthetic PEGDM, cross-linked with UV light in the presence of photoinitiator (Irgacure 2959, Ciba Specialty Chemicals), increased the elastic modulus of the resulting IPN hydrogel up to 4 kPa, which, unlike a pure fibrin gel, is stiff enough to be deployed by a surgeon in a clinical setting. The IPN hydrogel was able to support cultured human keratinocytes for 21 days with 90% cell viability. Similarly, Zhou and colleagues incorporated methacrylated poly(vinyl alcohol) (PVA) into maleilated chitosan hydrogels, which increased the Young’s modulus of the resulting IPN hydrogel from 380 Pa to 8000 Pa (Zhou et al., 2018). When combined with biocompatibility studies, these hydrogels were shown to have the mechanical integrity and bioactivity to serve as viable skin-thickness wound dressings.

Tissue-engineered bone needs to display physiological mechanical strength because (1) bone is subjected to cyclic, high-intensity loading, and (2) the stiffness of bone tissue regulates osteoblast differentiation. To satisfy these two physiological requirements, Chen et al. fabricated a PEG diacrylate-collagen IPN hydrogel intending to improve the mechanical properties of pure collagen hydrogels (Chen et al., 2017). This IPN hydrogel exhibited an average compressive fracture strength of 0.68 MPa, which was 9–12 times greater than that of pure PEG and 36–48 times that of pure collagen. Although this value remains significantly lower than the compressive strength of human cortical bone (141.6 MPa for males and 118.91 MPa for females, (Havaldar et al., 2014)), it is nevertheless a significant improvement compared to pure collagen or PEG hydrogels.

Cartilage tissue must be able to withstand high-pressure cyclic loading while demonstrating minimal deformation. To this end, anisotropy in cartilaginous tissue is essential because it allows cartilage to absorb and distribute loads on the axial (compression) and transverse (shear) axes. An agarose-PEG IPN hydrogel was embedded within a 3D poly(caprolactone) (PCL) scaffold to incorporate anisotropy in cartilage tissue replacements (Figure 3A) (Moffat et al., 2018). The combination of the three polymers created differential microenvironments, thus providing anisotropy to the scaffold to mimic the nonlinear stress-strain response of cartilage. Miao et al. fabricated a cartilage tissue model by incorporating porosity into mechanical enhanced PVA-gelatin IPN hydrogels (Miao et al., 2015). Initially, the hydrogel consisted of PVA-gelatin and PEG; the PEG aggregated in large domains and phase-separated from the PVA. When the PEG porogen was removed via dialysis, the PVA-gelatin IPN hydrogel was left with large pores that allowed for cell attachment and viability. The incorporation of PVA improved the storage modulus compared to pure gelatin from less than 1 Pa to over 1 kPa. Interestingly, the use of the PEG porogen further increased the storage modulus to 5.3 kPa, yet the authors did not explain this phenomenon. Despite these improvements, these engineered tissues remain in need of further reinforcement, as human articular cartilage has a storage modulus several orders of magnitude higher. In summary, the addition of a synthetic polymer lends considerable strength to single-polymer biological hydrogels without adversely affecting biocompatibility and cell viability.

Figure 3:

Biological-synthetic IPN hydrogels in tissue engineering (A) MSC-seeded biological-synthetic IPN hydrogels consisting of agarose and PEG-DA were interwoven with PCL fibers to create anisotropic tissue scaffolds capable of increasing chondrogenesis in engineered cartilage. Reprinted with permission from (Moffat et al., 2018). (B) Uncross-linked GelMA (turquoise panel) was physically cross-linked with an existing PEG network cross-linked with thiol-yne click chemistry to create cytocompatible IPN hydrogels. To further increase the stiffness of these scaffolds, gelatin was chemically cross-linked with thiol-ene click chemistry (magenta panel). Reprinted with permission from (Daniele et al., 2014). (C) An IPN hydrogel synthesized from fibrin and PAA underwent flexion in response to an applied electrical field, which, in turn, encouraged the migration and alignment of smooth muscle cells. Reprinted with permission from (Rahimi et al., 2012). Copyright (2012) American Chemical Society.

Decoupling stiffness and density in synthetic-biological hydrogels

Varying the gelation time or cross-linking density allows for the tuning of matrix stiffness without needing to change the concentration of polymers in the IPN, i.e., decouples the stiffness from the bulk properties of the material. This decoupling allows for detailed studies on how substrate stiffness affects tissue development. For example, Fares et al. created a set of IPNs and sIPNs with GelMA and pectin-g-PCL (Fares et al., 2018). A range of matrix stiffnesses was obtained by varying the concentration of photoinitiator (Irgacure 2959), and Ca²⁺ to cross-link the GelMA and pectin-g-PCL, respectively. sIPNs were created by only cross-linking the GelMA with the photoinitiator. Varying the cross-linking density resulted in hydrogels with compressive moduli ranging from 3.1 to 10.4 kPa for the sIPN hydrogels and 39 kPa to 5 MPa for the IPN hydrogels. The IPN hydrogels were able to support encapsulated MC3T3-E1 preosteoblasts for five days with cell viability greater than 80%. However, the sIPNs were not appropriate for long term cell culture since they degraded in 3 days after cell encapsulation. Alternatively, a gelatin-PEG IPN hydrogel was synthesized by cross-linking the two polymers with thiol-ene and thiol-yne click chemistry, producing non-cytotoxic and homogenous cross-links (Figure 3B) (Daniele et al., 2014). Adjusting cross-linking to either covalently or physically cross-link gelatin resulted in a set of hydrogels with stiffnesses ranging from 10.8 kPa and 327.7 kPa. Although both the physically and chemically cross-linked IPN hydrogels were cytocompatible and supported the proliferation of endothelial cells, the stiffer chemically cross-linked hydrogels encouraged cell spreading and higher proliferation rates compared to the physically cross-linked hydrogel.

Synthetic-biological IPN hydrogels with tunable mechanical properties have been deployed in drug screening studies. For example, Lee et al. created gelatin-PEG IPNs in order to model tumor microenvironments and thereby probe the effect of matrix stiffness on drug resistance (Lee et al., 2019). PEG does not degrade in vitro; therefore, the IPN hydrogel can maintain its mechanical properties as the encapsulated cells degrade the gelatin and lay down their own ECM. Matrix stiffness was changed by increasing or decreasing gelation time since both polymers were physically cross-linked. These studies revealed that matrix stiffness could impact drug resistance: cell viability after exposure to the chemotherapy drug 5-FU dropped from 80% in the low elastic modulus hydrogels (270 Pa) to less than 60% in the high elastic modulus hydrogels (1300 Pa). It is important to note that the impact of stiffness on cell morphology and proliferation can appear contradictory, as high stiffness in different hydrogels can either promote metastatic growth or encourage a rounded, stunted morphology. The effect of stiffness on cells in 3D may show a Gaussian-like dependency: cells extend and grow on substrates/hydrogels of intermediate stiffness, and are adversely affected by microenvironments that are significantly more or less stiff than physiological tissue (Crosby and Zoldan, 2019). Furthermore, these studies demonstrate the substrate stiffness, decoupled from bulk polymer concentration, impacts cell maintenance and proliferation; further work needs to be conducted to leverage these findings for therapeutic benefit effectively.

Synthetic polymers aid in the creation of electrically conductive tissue scaffolds

The creation of electrically-conductive tissue scaffolds could lead to significant breakthroughs in the engineering of cardiac, muscle, nervous, and dermal tissue, as well as provide additional insights into molecular-level physiological mechanisms (Hardy et al., 2013). However, conductive, synthetic polymers are non-degradable or lack cell adhesion sites; these limitations have consequently hindered these polymers’ tissue engineering applications. Adding conductive polymers to hydrogels that already contain a cross-linked bioactive polymer creates IPN hydrogels that can support cells and dynamically respond to electrical flux in the local microenvironment. Scaffolds that are responsive to electrical signaling could be beneficial for growth and function of cell types, such as neurons, that rely on the electrochemical coupling to enable physiological function. For example, poly(3,4-ethylene dioxythiophene) (PEDOT) is an electrically conductive polymer that can enhance neuronal growth via conductivity. However, PEDOT’s high stiffness and non-degradability make this polymer ill-suited for in vivo applications. To deploy the promising electrical properties of this polymer in a neural tissue engineering setting, Xu et al. created an IPN hydrogel with PEDOT and chitosan to retain PEDOT’s conductivity while lowering the overall stiffness of the matrix and imparting degradability (Xu et al., 2018). Xu and colleagues found that mouse neural crest cells attached and proliferated more on matrices containing PEDOT. Likewise, Rahimi et al. synthesized an electrically conductive, bioactive IPN hydrogel for vascular tissue engineering applications by combining polyacrylic acid (PAA) and fibrin (Figure 3C) (Rahimi et al., 2012). An applied electric field stimulated smooth muscle cells to migrate from the surface of the hydrogel to the interior, with cells seeded on electrically stimulated-hydrogels penetrating up to 700 μm into the hydrogel within one week compared to 200 μm in the unstimulated condition. A follow-up study from the same research laboratory noted that the electric field stimulated smooth muscle cells to preferentially align in the direction of the applied, deposit collagen, and increase matrix metalloproteinase-2 (MMP-2) activity, all of which are vital for the creation of biomimetic vascular grafts (Rahimi et al., 2014). Therefore, both PEDOT and PAA have tremendous potential in IPN hydrogels that encapsulate cells, such as those found in nervous and muscle tissue, that rely on electrical signaling to reproduce physiological function. Electrical conductivity may be especially critical in the manufacture of functional cardiac tissue, as it is increasingly realized that electrical stimulation in 3D environments (e.g., electrically conductive hydrogels) may be necessary to encourage fetal-like cardiomyocytes to mature and show characteristics of an adult phenotype (Tu et al., 2018).

Degradation encourages ECM deposition and releases therapeutic agents

The degradation of IPN hydrogels encourages transplanted cells to deposit matrix proteins while the IPN degrades, to migrate through an otherwise constraining microenvironment, and to undergo morphological changes. This deposition, migration, and morphological change, in turn, stimulates various cell-signaling pathways. For example, to encourage hMSC osteogenesis, cellulose nanowhiskers (CNW) were cross-linked with poly(caprolactone) (PCL) and poly(2 hydroxyethyl methacrylate) (PHEMA) (Shahrousvand et al., 2019). CNWs, which are inexpensive and naturally abundant (Fragal et al., 2016), were added to increase cell viability in the otherwise synthetic-synthetic scaffold, PHEMA increased the tensile strength of the scaffold, and PCL represented the biodegradable component of the IPN. A three-step degradation study (incubation, induction, and erosion) suggested that an increasing concentration of CNW increased the degradability of the scaffold. Another study described the synthesis of a co-network of polyvinyl alcohol (PVA) and serum albumin (SA) with fibrin. PVA increased the mechanical integrity of the scaffold by an order of magnitude, and, when cross-linked with SA, created a network that was sensitive to enzymatic degradation (Bidault et al., 2015). It was demonstrated that PVA-SA-fibrin hydrogels maintained the viability of fibroblasts for over one week of culture and approximately doubled their proliferation rate. Since the PVA adds additional mechanical support, these scaffolds may avoid the compaction seen in many traditional fibrin-based systems. This system is one of the first IPN hydrogels that degrades through fragmentation and elimination, making these hydrogels excellent candidates for tissue engineering.

Similarly, poly(ethylene glycol) (PEG) diacrylate macromers were combined with HA in an IPN hydrogel to create a scaffold sensitive to two modes of degradation, hydrolysis and enzymatic (Kutty et al., 2007). Another study detailed the synthesis of a differentially degradable HA IPN hydrogel that contained the anti-inflammatory drug dexamethasone. Dexamethasone was activated upon HA degradation and subsequently induced hMSC osteogenic differentiation, which could lead to the application of this IPN hydrogel in bone tissue engineering applications (Bae et al., 2015). Skaalure et al. created a PEG-HA sIPN cartilage tissue replacement to demonstrate that increasing HA concentration increased the levels of matrix deposition by encapsulated chondrocytes. Furthermore, lower molecular weight HA increased the concentration of matrix-degrading enzymes, which could provide a novel way to tune these IPNs (Skaalure et al., 2014). In summary, degradability is a critical paradigm in tissue engineering, and IPN hydrogels consisting of synthetic and biological can exhibit unique, tunable degradation profiles that approximate physiological activity and can be adjusted to fit the needs of the engineered tissue.

IPN hydrogels tissue scaffolds derived from synthetic polymers

In the previous section, we reviewed how the integration of a synthetic polymer network can improve the material properties of a biologically-derived single-polymer hydrogel without significantly compromising the bioactivity of the construct. The monomers and side chains of synthetic polymers can be precisely engineered to confer mechanical strength and viscoelasticity to the final hydrogel. However, the purification, immunogenicity, and the possible transmission of pathogens continue to hinder the translation of many animal-derived polymers to tissue-engineered constructs, increasing the need to develop biomimetic synthetic polymers (Lutolf and Hubbell, 2005). Yet, in the last six years, few synthetic IPN hydrogel composites have been developed for tissue engineering applications, with two notable exceptions. In 2015, Parke-Houben et al. combined PEG and PAA with polystyrene beads to create high-strength hydrophilic scaffolds for corneal tissue engineering applications. The polystyrene beads were then removed from the construct, resulting in differential pore sizes that were directly related to the polystyrene bead packing efficiency. Since both PEG and PAA are typically inert in biological settings, type I collagen and fibronectin were tethered to PAA via 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC)/N-Hydroxysuccinimide (NHS) coupling chemistry to allow for primary corneal epithelial cell attachment and proliferation (Parke-Houben et al., 2015). PAA was also photopolymerized with a copolymer of PNIPAAm and PEGDMA to create a more mechanically stable hydrogel for multiple bone fragment surgery. Since PNIPAAm and PEGDMA are also inert in biological settings, two ceramic additives, tricalcium phosphate (TCP) and nanoparticles of hydroxyapatite (nano-HAP), were incorporated into the final cross-linked hydrogel and were demonstrated to slightly increase 3T3 fibroblast metabolic activity and viability (de Lima et al., 2019). In both studies, additional polymers and ceramic particulates were added to lend bioactivity to the final construct. Therefore, the native inertness of many synthetic polymers remains one of the significant obstacles to the proliferation of synthetic polymer engineering techniques in tissue engineering and regenerative medicine.

Future applications of IPN hydrogels in tissue engineering

In 2016, Khademhosseini and Langer outlined six grand challenges to future tissue engineers: (1) clinical translation (2) organs-on-a-chip/disease modeling (3) biorobotics/bioactuators (4) engineered meat and (5) leather, and (6) cryopreservation/fast delivery (Khademhosseini and Langer, 2016). IPN hydrogels, and most biomaterials to date, have yet to be applied in five of these six new directions. One critical exception is the use of IPN hydrogels in organs-on-a-chip and disease modeling. Briefly, advances in microfabrication technologies paired with breakthroughs in materials science and stem cell biology have led to the development of the organ-on-a-chip, a new microfluidic-based platform, that promises to mimic tissue barrier function, parenchymal tissue function, and multi-organ systems (Zhang et al., 2018a). While this recent technology still faces significant translational hurdles before it can be manufactured by large pharmaceutical companies (Haddrick and Simpson, 2019), the possibility of creating personalized drug screening platforms that mimic physiological drug transport is undeniably an exciting prospect. In 2018, Qiu and colleagues were one of the first research laboratories to deploy an IPN hydrogel within an organ-on-a-chip platform (Qiu et al., 2018). In this breakthrough study, it was determined that an alginate-agarose IPN hydrogel could provide the necessary stability to examine microvascular obstruction and endothelial permeability in response to pathological conditions for more than two weeks. Furthermore, the stiffness of the alginate-agarose composites could be tuned to the physiological range of vessels (approximately 20 kPa). Though this remains the only example of an IPN hydrogel applied in an organ-on-a-chip platform, the tunability and bioactivity of IPN hydrogels will continue to make these tissue engineering scaffolds attractive options for incorporation into such devices. Another technology that may prove pivotal in the high-throughput utilization of organ-on-a-chip devices is the widespread adoption of bioprinting (Nie et al., 2020).

As mentioned in the introduction of this review, synthesizing functional tissues and organs has proved challenging and largely remains an unmet goal. One hurdle that has hindered this translation is the lack of cell-friendly fabrication technology to precisely engineer hierarchical architecture at multiple length-scales (Williams et al., 2018). 3D bioprinting, which uses established additive manufacturing practices to precisely control the fabrication of freeform solid tissue constructs, has been increasingly deployed in tissue engineering to better mimic tissue complexity. Though bioprinters differ in their construction, they all rely on the use of bioinks. Bioinks, by definition, are cell formulations that can be processed by automated biofabrication technology (e.g., bio-printers) (Groll et al., 2016). In extrusion bioprinters, bioinks flow in response to shear stress (i.e., are shear-thinning) to allow for controlled deposition and are cross-linked, physically or chemically, shortly after extrusion. Biomaterial inks, in contrast, do not inherently contain cells and are often employed as sacrificial materials to support the integrity of the newly-printed tissue (Groll et al., 2019). With few exceptions, almost all instances of IPN hydrogels as bioinks or biomaterials inks have included alginate as a constituent because this biologically-derived polysaccharide cross-links within seconds upon exposure to a solution containing a sufficiently high concentration of calcium ions. For example, in a study that is representative of this rapidly expanding field, Chen and colleagues synthesized two sIPNs from gelatin and alginate and demonstrated that neuroblastoma SH-SY5Y cells remained viable for up to two weeks of in vitro culture (Chen et al., 2020). Krishnamoorthy et al. also used an alginate-gelatin IPN while further optimizing the nozzle speed and dispensing pressure of a micro-droplet 3D bioprinter to generate bioactive millimeter-scale bifurcated 3D printed constructs (Krishnamoorthy et al., 2019). In a study demonstrating that IPN hydrogel bioink might be synthesized from renewable marine sources, alginate-(fish-derived) GelMA bioink displayed higher mechanical strength and decreased swelling and degradation in extended culture (Figure 4A) (Zhang et al., 2018b). It was further demonstrated that NIH 3T3 fibroblasts remained viable for up to a week in alginate-f-GelMA 3D-printed constructs.

Figure 4:

IPN hydrogels in bioprinting (A) Gelatin derived from fish was combined with alginate to create a biological-biological bioink that could be printed into large scaffolds that encouraged 3T3 fibroblast viability. (B) An alginate-PLA nanofiber IPN hydrogel was combined with MSCs to form a novel bioink. After being extruded through a syringe-based bioprinter, the resulting cartilaginous tissue constructs encouraged MSC viability and chondrogenesis. Reprinted (adapted) with permission from (Narayanan et al., 2016). Copyright (2016) American Chemical Society.

Recently, IPN-based bioinks have been utilized to tissue engineer viable connective tissue replacements. For example, Narayanan et al. encapsulated human adipose-derived stem cells (hASCs) in an alginate hydrogel reinforced with polylactic acid nanofibers to create a nanofibrous bioink that could reconstitute the native ECM of a meniscus. The authors 3D-printed a sample meniscus, generated from a magnetic resonance imaging (MRI) scan of a real patient, to demonstrate the clinical applicability of their methodology (Figure 4B) (Narayanan et al., 2016). To develop materials to aid intervertebral disc regeneration, Hu et al. combined a gellan gum-poly (ethylene glycol) diacrylate (GG-PEGDA) IPN hydrogel with a robust PLA scaffold (Hu et al., 2018). The resulting scaffold displayed enhanced mechanical properties and encouraged the viability and proliferation of bone marrow stromal cells.

Conclusions and Recommendation

In conclusion, we have shown that tissue engineers have increasingly utilized IPN hydrogels in the past five years to biomanufacture complex tissue replacements. The first challenge outlined by previous reviews has been partially resolved: there are many IPN hydrogels for tissue engineering applications that show improved strength, degradability, and bioactivity. However, a detailed mathematical model to predict and validate the parameters extracted from mechanical tests is still lacking, which will significantly advance further progress in the field. IPN hydrogels have been effectively deployed in bioprinting and organ-on-a-chip devices, and the use of IPN hydrogels in other applications such as the bio-manufacture of meat substitutes and the creation of biomimetic leather is beginning to appear on the horizon. We argue that IPN hydrogels should continue to be a critical component in the tissue engineer’s toolkit and will enable the long-awaited translation of engineered tissues and organs to the clinic.