Keywords

1 Introduction

In the field of regenerative medicine, scientists aim to re-create or repair living, functional tissue up to the size of complete organs. Complete organs are one of the greatest promises of regenerative medicine, as these organs could potentially replace donor organs and provide a solution for the shortage of available organs for transplantation [1]. Regenerative medicine and other biomedical applications are naturally only concerned with aqueous systems and, therefore, hydrogels are widely investigated, for example, as scaffolds for tissue engineering and as drug delivery vehicles [2–7]. Although polymeric hydrogels have already proven their application in medicine (e.g., as contact lenses), supramolecular hydrogels are still emerging as biomaterials, mainly because of a lack of design rules [8, 9].

Supramolecular hydrogels consist of supramolecular polymers that are able to form freestanding three-dimensional (3D) networks when swollen in water. Supramolecular polymers consist of monomers or oligomers linked together by noncovalent, well-defined, directional interactions such as hydrogen bonding arrays, steric interactions, metal–ligand complexes, hydrophobic interactions, van der Waals forces, π–π stacking, and other electrostatic effects [10]. Via these supramolecular interactions, both long polymers and physically cross-linked networks can be formed [11]. Supramolecular polymers have found use in many other polymer applications such as adhesives, cosmetics, and coatings [12].

Supramolecular systems offer distinct advantages as biomaterials over chemically cross-linked gels. First, because of their noncovalent nature, supramolecular systems are inherently responsive to stimuli such as variations in temperature and pH. Second, the noncovalent bonding allows a unique mix-and-match principle to be used for tuning properties. Third, supramolecular materials biodegrade faster than chemically cross-linked gels as a result of the small molecular precursors that these materials are made of. Furthermore, these supramolecular systems are proposed to be able to display dynamic reciprocal behavior, as found in the natural environment of cells [13]. Supramolecular biomaterials are proposed to be able to spatiotemporally adapt to changes exerted by cells and their natural environment. To fulfill this promise, important features of supramolecular systems are their hierarchical structure/assembly, their dynamic and nonlinear behavior, and their biochemical properties.

The use of supramolecular hydrogels in regenerative medicine is inspired by the extracellular matrix (ECM), which provides the natural scaffolding for cells inside the body [14]. Mimicking the ECM is a major objective for tissue engineering, in order to create a regenerating niche for cells [15]. The natural ECM is inherently dynamic, and the structural support is given by macromolecules such as proteoglycans and fibrous proteins; therefore, supramolecular hydrogels might be an ideal replacement [16]. Furthermore, the ECM provides cells with handles for attachment and signaling, and it regulates the transport and presentation of soluble components such as growth factors. These three roles of the ECM – mechanical support, cell signaling, and soluble factor transport/presentation – need to be fulfilled by an ideal ECM mimic.

The mechanical properties of the ECM in natural tissue show a highly variable stiffness, ranging from hundreds of pascals in the mammary gland to megapascals for articular cartilage [17] (Fig. 1). These tissues resist tensile forces using fibrillar proteins such as collagens, whereas hydrated proteins and glycosaminoglycans resist compressive forces [15]. Cells can sense this mechanical microenvironment and respond to the ECM by changing their own mechanical properties via the cytoskeletal network [18]. Therefore, for adherence, proliferation, and differentiation of cells, the mechanical components of designed cell matrices are essential [19–21]. However, hydrogels generally possess inferior mechanical properties, such as mechanical toughness and resistance to friction, compared with natural tissues [5]. Therefore, it is important to improve the mechanical properties of hydrogels and tailor them for their specific application.

Fig. 1
figure 1

Supramolecular systems described in this chapter. Top: The extracellular matrix inspires most of the work and consists of fibers to resist tensile stresses such as collagens, fibers to resist compressive stresses such as the glycosaminoglycans, and soluble factors for cell signalling. Bottom: Natural systems such as collagen, coiled-coil structures, β-sheet peptides such as peptide amphiphiles and multidomain peptides, hybrid systems such as streptavadin–biotin cross-linked microparticles, and synthetic systems such as the UPy-based hydrogelators

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Besides the physical factors, cell–matrix interactions are also regulated by biological factors. Thus, hydrogels should contain bioactive signals to facilitate and enhance these cell–matrix interactions. Furthermore, they also need to fulfill special functions such as the attraction of specific cells and controlled release of bioactives to regulate cell behavior. In the development of supramolecular hydrogels for biomedical applications, it is necessary to take into account the tuning and optimization of the mechanical properties of these hydrogels as well as their functionalization with bioactive components.

Numerous reviews describe supramolecular hydrogels in detail [22–28]; therefore, this article is focused on those supramolecular hydrogel systems designed for applications in regenerative medicine and tested in vivo, as translational research is essential to show the applicability of these materials. We discuss the chemical structures of the systems, their mechanical properties, and the introduction of bioactivity. Additionally, we show examples of various applications in regenerative medicine. Components derived from the ECM have been used extensively as hydrogelators, and we first discuss some of these natural components together with other natural supramolecular polymers that have been used in vivo (Sect. 2) (Fig. 1). Next, we discuss biosynthetic compounds, where man-made peptides are used to mimic the fibrous proteins found in the ECM (Sects. 3 and 4). Then, we discuss hybrid hydrogels, which combine natural structures with synthetic compounds (Sect. 5). In Sect. 6, fully synthetic compounds, inspired by the ECM but showing no resemblance in the chemical structure, are discussed.

2 Supramolecular Hydrogels Based on Natural Polymers

The use of natural polymers for the preparation of hydrogels is inspired by the ECM, which consists of many small and large natural polymers such as collagens and polysaccharides. The use of these hydrogelators and other natural polymers for tissue engineering has been reviewed extensively, for example, the use of elastins, glycosaminoglycans, hyaluronic acids, collagens, gelatins, and keratins [24, 29–33]. Often, chemical cross-linking is essential to obtain stable gels. Here, we describe natural polymers that form 3D networks via supramolecular interactions.

Collagen is the main structural component of the interstitial ECM and basement membrane. Different types of collagens exist, but all collagens share the repeated –Gly–Xaa–Yaa– sequence and have a high proline and hydroxyproline content [34]. Three peptides containing this repeating sequence form a triple helical chain that subsequently bundles together with other chains. These bundles align to form long collagen fibrils (Fig. 2a). Self-supporting gels can be formed via end cross-linking of these fibers. They show shear moduli of tens of pascals under physiological conditions, compressive moduli of tens of kilopascals, and tensile moduli of hundreds of kilopascals. This shows the importance of collagens in resisting tensile forces in the ECM [15, 35, 36]. The properties of collagen depend heavily on temperature and pH, and they can be altered by simply changing the pH during polymerization [37, 38]. By increasing the pH, fibrils of collagen become longer and thinner, leading to increased moduli and failure stresses. However, under physiological conditions, no orders of magnitude difference in moduli can be obtained. Collagens contain cellular binding epitopes, for instance the CNYYSNSYSFWLASLNPER sequence in collagen type IV [39]. Cell binding can also be mediated by mixing glycoproteins with the collagen substrate [40] or by mixing in drugs. Collagenous scaffolds have been used extensively for biomedical applications because of their inherent biocompatibility. They have often been applied as skin grafts or together with other cells and biopolymers as disks, meshes, or foams in tissue such as cartilage and bone [41, 42]. As early as 1990, aliquots of collagenous hydrogels were used as a drug delivery system for sustained delivery of an anticancer drug in the mouse brain [43].

Fig. 2
figure 2

Supramolecular natural polymeric hydrogels discussed in this chapter. (a) Chemical structure of the most-repeated sequence in collagen, forming the α-chain that folds in a three-stranded superhelix [135]. These superhelices bundle to form the collagen fiber. (b) Representative chemical structure of fibroin and the antiparallel β-sheet formation connected by hydrophilic linkers. (c) Chemical structure of alginic acid, cross-linked by calcium ions (highlighted). (d) Left: Top view of two α-helixes of keratin forming a coiled coil by hydrophobic interactions. Right: Overview of subsequent formation of the fibril. The left part is adapted from [57] with permission of The Royal Society of Chemistry

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Natural polymers not present inside the body have also been tested for regenerative medicine. Silks are structural proteins, and the fibers from the Bombyx mori silkworm are the most-used silks for biomedical applications [44]. Silk mainly consists of fibroin, a protein rich in glycine and alanine that forms β-sheets and becomes insoluble in water, forming cross-links for the gel (Fig. 2b) [45]. Several techniques have been used to prepare silk scaffolds, such as freeze-drying, salt leaching, and gas foaming, which influence the mechanical properties of the silk fibroin gels [46]. Sonication and vortexing have also been used to aid gelation. The properties of silk fibroin hydrogels depend on the processing conditions. Gel formation is irreversible, and compressive moduli are largely determined by the pore sizes in the gel, which can be influenced by silk fibroin concentration, gelation temperature, and calcium ion concentration [47, 48]. In general, compressive moduli of 30–3,000 kPa are obtained. Increasing the concentration or temperature leads to accelerated formation of the physical cross-links. Without the presence of ions, this leads to increasing compressive strength and moduli. Bioactivity can be obtained by mixing in growth factors or other types of bioactive natural polymers, such as collagen, gelatin, and glycosaminoglycans [31]. Several examples of the use of these silk fibroin hydrogels in regenerative medicine exist. Silk fibroin hydrogels aided bone remodeling and maturation in a critical size defect in rabbits, but degraded within 12 weeks [49]. By encapsulation of vascular endothelial growth factor (VEGF) and bone morphogenetic protein 7 (BMP-7) in a sonication-induced silk fibroin hydrogel, bone regeneration was promoted in rabbits [50].

Alginate is a polysaccharide derived from algae that has many carboxylic acid groups and therefore can be cross-linked by calcium ions (Fig. 2c). The mechanical properties of hydrogels formed by alginates are determined by molecular weight when there is a low cross-linking density and by the guluronic acid content; hence, at high cross-linking density the properties depend on the amount of cross-links on the alginate [51]. By varying the polymer and salt concentration, compressive moduli of 5–200 kPa can be obtained. Often, alginate hydrogels are combined with other natural polymers such as collagens to influence cell attachment, or with bioactive signaling proteins such as growth factors to influence cell signaling. In one example, alginate gels were loaded with angiogenic growth factors and placed subcutaneously in the dorsal area of rats [52]. Sequential delivery of three growth factors [VEGF, platelet-derived growth factor, and transforming growth factor beta (TGF-β)] led to enhanced formation of stable and mature blood vessels after 3 months.

The use of human-hair-based keratin has also been explored, but only in a limited number of regenerative-medicine applications [53]. Keratin consists of a family of structural proteins, of which we have already discussed one member, fibroin. In the examples below, we discuss keratin from the human hair, α-keratin, which consists of two α-helices that form a coiled coil as a result of hydrophobic interactions (Fig. 2d). The process of converting hair into a hydrogel consists of oxidizing the keratin and subsequent breaking of the disulfide linkages. After cooling to room temperature, the remaining disulfide and hydrogen bonds form the gel [30]. Temperature, concentration, and vortexing can be used to tune the shear moduli of the gel between 1 Pa and 1 MPa [54]. About 78 % of human hair keratins contain the cell-binding sequence RGD and, therefore, cells easily attach to keratin [55]. Keratin hydrogels have been used in a mouse tibia nerve injury model, where the hydrogel caused improved axon regeneration and functional recovery [55]. In another study, keratin gels were used to stop bleeding in a lethal liver injury in rabbits [56]. Enhanced 24-h survival rates were shown for rabbits treated with keratin hydrogels, and the keratin gels performed better than commercial hemostatic agents used to stop bleeding.

Natural materials offer many advantages, such as inherent biocompatibility and the natural presence of bioactive sequences. However, problems can arise with batch-to-batch variability, contamination by pathogens (when derived from animal origin, for example, collagen I), and contamination by other non-compatible components. For example, silk contains, in addition to the structural protein fibroin, a glue between the fibers, sericin, which needs to be removed because it can lead to problems with biocompatibility and hypersensitivity [44]. Furthermore, tuning the mechanical properties of natural hydrogels is only possible in a limited range. To be able to adapt materials for different applications with the full range of mechanical properties found in natural tissue, the mechanical tailoring possibilities of synthetic materials can be beneficial.

3 Biosynthetic Supramolecular Hydrogels Based on Coiled Coils

Biosynthetic hydrogels are synthetic systems with binding motifs copied from nature and allow exact control over the chemical structure. Therefore, designed supramolecular hydrogelators can be chemically modified to alter the mechanical properties and to include bioactive sequences.

A large body of literature concentrates on the use of coiled coils as cross-linkers for the development of biomaterials [57, 58]. Coiled coils are formed by two interacting α-helices, where, once folded, the subsequent orientation of hydrophobic amino acids leads to the formation of dimers in a superhelical fashion. These coiled coils are some of the most abundant oligomer folding motifs in nature. Tirrell and coworkers developed a large artificial protein consisting of a water-soluble [(AG)3PEG]10 random coil block flanked by two leonine-zipper blocks, all prepared by bacterial expression [59]. These end blocks can be selectively folded into an α-helix, forming coiled coils and subsequent gel networks with shear moduli of 200 Pa (Fig. 3a).

Fig. 3
figure 3

Coiled-coil hydrogelators. (a) Secondary structure of a coiled coil with soluble linkers (left), ultimately leading to a 3D network (right). (b) Engineered two-component hydrogel, where bioactivity is introduced in the hydrophilic spacer via the RGDS sequence. Adapted with permission from [60]

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These leucine-zippers require the use of a trigger such as pH or temperature, for example, to encapsulate cells. Therefore, a two-component system was developed using recombinant protein engineering [60]. The first component consists of a small peptide spacer extended with WW domains (with conserved tryptophan residues, abbreviated by the single-letter amino acid symbol W), which binds to proline-rich peptides incorporated in the second component (Fig. 3b). Protein engineering gives full control over the sequence and, hence, over binding affinity and subsequent bulk properties. This two-component material forms a weak gel with shear moduli of 10–50 Pa, depending on the type of proline-rich sequence used. The binding affinity of the complex can be increased by an order of magnitude by changing the WW component or by changing the amount of repeating units of both binding domains and, thereby, changing the gel point and modulating the plateau modulus slightly [60]. These two-component gels show shear-thinning and self-healing behavior and are therefore suitable as injectable material. Because the material only forms a gel after mixing, simultaneous addition of cells allows mild incorporation of these cells into the material.

Because these peptides are made via protein engineering, cell-adhesion RGD peptide sequences can be encoded in the spacer between the binding motifs, thereby adding bioactivity to the hydrogels. By incorporation of the RGD-motif, PC-12 cells, human umbilical vein endothelial cells (HUVECs), and neural stem cells can be conveniently cultured in 3D. Implanted subcutaneously in mice, this material shows improved retention of adipose-derived stem cells and is therefore promising as a cell carrier in stem cell injection therapies [61].

Poly(ethylene glycol) (PEG) has also been end-functionalized with these α-helical coiled-coil structures to form hydrogels with storage moduli of 1 kPa [62]. Additionally, it was shown that the self-assembly can only induce an immunogenic response in mice as a result of formation of oligomeric aggregates. The individual components do not induce a response, showing that the effect of the self-assembly can be different from the effect of its individual components and should be taken into account when studying supramolecular biomaterials [63].

Coiled coils are interesting as designable cross-linkers because of the well-understood relation between the peptide sequence, the folded and dimerized structure, and the binding affinity. However, as described above, hydrogels based on coiled coils are generally weak, with shear moduli only up to 1 kPa. Furthermore, the use of recombinant protein expression or solid-phase peptide synthesis limits the scale on which these materials can be prepared. However, this approach allows tuning both the cross-links as well as the backbone to regulate the mechanical properties and incorporate bioactivity.

4 Biosynthetic Supramolecular Hydrogels Based on β-Sheets

Peptide-based hydrogelators are a special class of small molecular building blocks in which short peptide sequences are not only used to induce bioactivity, but also to induce self-assembly into a hydrogel. Several groups have investigated this class of hydrogels for application in tissue engineering, inspired by the self-assembly found in nature and because of the inherent biocompatibility and biodegradability of peptides. Several β-sheet-forming peptides have been developed for use in regenerative medicine and have shown promising potential in vitro; nevertheless, to the best of our knowledge, they have not yet been applied in vivo. For example, Schneider, Pochan, and coworkers developed the MAX8 sequence, an amphiphilic β-hairpin peptide of 20 amino acids of alternating hydrophobic valine and hydrophilic lysine or glutamic acid, folded by prolines in the center, yielding gels with shear moduli of a few hundred pascals [64]. The laboratory of Collier has developed a β-sheet-forming peptide, Ac-QQKFQFQFEQQ-Am (Q11), which slowly assembles into gels via long single fibrils in aqueous environments, resulting in gels with moduli of around 10 Pa [65–67].

Peptide amphiphiles (PAs) are well known to form functional supramolecular materials [68]. PAs consists of (i) a short hydrophobic domain, often an alkyl chain, linked to (ii) an oligopeptide that induces and guides self-assembly by the formation of β-sheets, and (iii) an oligopeptide containing a bioactive domain [69]. In water, these molecules assemble into high-aspect-ratio nanofibers, which entangle and form self-supporting hydrogels. This fibrillar structure mimics the ECM components, including the signaling peptides at the periphery.

The chemical structure of the PAs guides self-assembly into nanofibers and ultimately determines the mechanical properties of the hydrogel (Fig. 4a). The self-assembly of these amphiphiles is triggered by screening of the charges on the epitope by counter-ions or changes in pH [70]. These counter-ions also induce a change in the molecular geometry, which, together with the van der Waals forces, hydrogen bonding, and hydrophobic forces, induces self-assembly into cylindrical micelles. Within these nanofibers, parts of the peptide form β-sheets, leading to high-aspect-ratio objects (Fig. 4b). Mechanical properties originate from the strength and stiffness of these nanofibers and the density and strength of cross-links. Counter-ions have an effect on the fiber stiffness, through which the moduli can be modulated, leading to gels with shear moduli of 100 Pa to 10 kPa. The incorporation of carboxylic acids at the periphery of the PA allows cross-linking of fibers with divalent cations such as calcium, giving rise to different gelation kinetics and networks that can withstand more strain, although the moduli are not influenced [71].

Fig. 4
figure 4

Self-assembling hydrogelators based on β-sheets. (a) Representative chemical structure of a peptide amphiphile, here without charged residues and with a heparin binding domain. (b) Peptide amphiphile with bioactive epitopes (left) and its assembly leading to formation of 1D fibers (right). Reprinted from [136], Copyright 2010, with permission from Elsevier. (c) Graph showing the enhanced functional recovery for animals treated with the peptide amphiphile, as assessed via the BBB score. Adapted with permission from [78]. Copyright 2008 Society for Neuroscience. (d) RADA16 peptide, the formation of a β-sheet, and SEM image of fibers in a gel. Adapted from [84]. (e) (1–4) Functional recovery of the vision of animals treated with a RADA16 hydrogel, as shown by the response to a light trigger. Adapted from [89]. Copyright (2006) National Academy of Sciences, USA. (f) Chemical structure of an amphiphilic PSFCFLFEP peptide. (g) Chemical structure of an ABA multidomain peptide, here with an additional bioactive sequence attached

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Gels made from PAs have been used in several regenerative-medicine applications such as the induction of angiogenesis [72], bone regeneration [73], cartilage regeneration [74], enamel regeneration [75], and artery engineering [76]. We discuss three examples here, showing the different ways of introducing bioactivity into these PA systems.

Neural progenitor cells (NPCs) can be encapsulated inside PA gels [77]. By incorporating the laminin-mimicking IKVAV sequence, which promotes neurite sprouting and growth, cells were rapidly differentiated into neurons. Bioactivity was obtained by displaying a peptide-mimicking sequence on the periphery of the nanofibers. Furthermore, PAs in solution gel upon contact with tissue. In vivo experiments using a mouse spinal cord injury model showed promising results in the regeneration of the central nervous system [78]. Mice treated with IKVAV-PA showed enhanced regeneration of both descending motor axons and ascending sensory axons and showed functional recovery (Fig. 4c). This IKVAV-PA gel was used to deliver embryonic stem cells into the auditory nerves of rats [79].

Incorporation of a TGF-binding domain on the epitope of PA leads to capture and display of TGF-β1 in the hydrogels [74]. These materials were used to treat a full thickness chondral defect in a rabbit model. They promote the regeneration of articular cartilage by capturing growth factors from the host animal.

As a third example, inclusion of a heparin-binding domain on the periphery of PAs leads to self-assembled nanofibers, where heparin can be displayed on the nanofiber surface (Fig. 4a) [80]. Subsequently, this bound heparin can be used to capture angiogenic growth factors such as VEGF and fibroblast growth factor 2 (FGF-2) [81]. Implanted in the rat cornea, these materials induced significant blood vessel growth compared with the growth factors alone, showing the added value of a structured delivery system.

Another class of peptide hydrogel formed by β-sheets are hydrogels that consist of ionic alternating hydrophobic and hydrophilic residues, such as the materials based on a β-sheet-forming segment found in a yeast protein developed by Zhang and coworkers [82]. This EAK16 (AEAEAKAKAEAEAKAK) peptide forms β-sheets distinct polar and hydrophobic surfaces, enabling further self-assembly into 20-nm fibers, leading to stable hydrogels in water. In these types of systems, the length of the peptide can be used to tune the mechanical properties, as demonstrated by increasing KFE8 to KFE12 with an accompanying change in the shear moduli from 400 Pa to 2 kPa. This effect can be used together with changing the concentration to obtain gels with shear moduli of 400 Pa to 15 kPa [83].

The RAD self-assembling peptide (Fig. 4d) has been used to locally deliver growth factors such as EGF to accelerate wound-healing in an in vitro model [84]. Both the EAK and RAD systems have been further developed to increase cell adhesion [85], and functional motifs such as laminin and collagen-mimicking sequences have been attached and optimized to culture cells in a 3D environment [86, 87]. These peptides have been used for the delivery of genes to cells [88], and they have proven their feasibility in regenerative medicine by enabling reconnection of brain tissue in hamsters [89]. The peptide scaffold was used to knit together tissue after an acute injury and resulted in increased recovery of the central nervous system and recovery of functional behavioral (Fig. 4e).

Ruan et al. designed peptide-forming nanofibers via β-turn secondary folding [90]. The sequence of the peptide consists of the ionic amino acids lysine and glutamic acid on one end and two phenyl groups and a cysteine on the other end, creating an amphiphile (Fig. 4f). This amphiphilic PSFCFLFEP peptide also includes a proline to increase the formation of fibrils and to induce the β-turn. Hydrogels prepared with this peptide showed shear moduli of around 20 Pa, and addition of sodium chloride increased the moduli approximately tenfold [91]. The fibrils were able to gradually release hydrophobic model compounds such as pyrenes. The peptide hydrogel functioned as a hemostat, stopping bleeding of an injured surface on the rat liver faster than other hemostats.

Another example of a hydrogel based on β-sheet-forming peptides consists of a multidomain peptide (MDP) with an ABA structure [92]. The A-block consists of charged residues to keep the aggregates in solution, attached on both sides to a B-block consisting of alternating glutamine and leucine residues to create a facial amphiphile. Subsequent self-assembly of the B-block leads to poorly soluble aggregates in aqueous environments. This allows formation of a hydrogel. Both blocks of these MDPs can be changed to tune the mechanical properties of these gels, giving moduli ranging from 10 to 500 Pa [93]. Chemical cross-linking via disulfides even leads to gels with moduli of 6 kPa. The gels undergo rapid shear thinning and recovery, and have therefore been used as injectable drug carriers [94]. Bioactive sequences such as the RGD sequence can be incorporated into the material during the solid-phase peptide synthesis of the MDPs (Fig. 4g). Gels were pre-loaded with the secretome from stem cells, which contains more than 36 secreted proteins, and subsequently used to study the effect on a lipopolysaccharide model of acute kidney injury in mice and an ischemia-reperfusion model of kidney injury in mice [95]. The preconditioned nanofibers protect the kidney against injury by delivery of the secretome of stem cells, which is proposed to be the cause of the improvement resulting from stem cell treatment of acute kidney injury.

Many excellent examples of the biomedical application of peptide-based hydrogelators demonstrate their promise in regenerative medicine. The relation between the design of the peptide and the secondary structure is well understood and helps in designing these types of materials. Furthermore, bioactive sequences are attached to these materials in the same way as the self-assembling domain is synthesized. However, the scale on which the solid-phase peptide synthesis used for the preparation of these systems can be performed is limited. The mechanical properties of these materials can be tuned to obtain moduli between 10 Pa and 10 kPa, giving a larger working range than natural hydrogels and hydrogels based on natural motifs. The maximally obtainable moduli are only one order of magnitude lower than those of tissues found in the body.

5 Hybrid Supramolecular Hydrogel Systems

Systems have been developed that combine both natural and naturally derived compounds with synthetic structures. Many groups have adapted peptide-based structures or have used natural cross-linkers to develop hydrogels. Inspired by the collagens in the ECM, collagen mimetic peptides (CMPs) have been developed [96]. The development of solid-phase peptide synthesis has enabled synthesis of more stable collagen-like peptides such as (proline–hydroxyproline–glycine)10, (POG10), which forms stable short helixes. To obtain hydrogels, the focus has been on the self-assembly of these structures into nanofibers and 3D structures. The POG10 structure has been modified with three different metal-binding groups: histidine, nitrolotriacetic acid, and a bipyridyl moiety [97]. Subsequently, these structures were cross-linked using nickel ions. However, the mechanical properties of the resulting gels have not been determined. Bioactive sequences can be coupled covalently to promote cell adhesion [98]. By functionalization of growth factors with a His tag, these growth factors can also be supramolecularly complexed to the matrix using nickel ions.

Another interesting approach is the use of tubulin, which by itself forms long cylinders [99]. Forming a stable 3D gel using these cylindrical fibers requires cross-linking of the tubulin fibers with PEG polymers, yielding gels with shear moduli of 1.8 kPa. This cross-linking can only be performed when the tubulins are assembled into fibers to prevent blocking the supramolecular polymerization site of the tubulins. The gel still shows reversible gelation after cross-linking, demonstrating the supramolecular nature of this system. Furthermore, Wieduwild et al. prepared star-shaped PEGs end-functionalized with heparin-binding peptides, creating physical cross-links and subsequent gel formation upon addition of heparin; the obtained materials had shear moduli of around 2 kPa [100].

A different approach involves functionalization of chitosan with α-cyclodextrins and use of PEG as a thread through the α-cyclodextrins to form polypseudorotaxanes [101]. Cells can be encapsulated inside these gels, which showed fast degradation as a result of dethreading of the PEG. However, although these approaches show the promise of hybrid systems, the materials have not yet been tested in vivo.

Yang and coworkers developed a hydrogel based on two amino acid derivatives and bisphosphonate, which self-assemble into hydrogels with shear moduli of 20 Pa to 20 kPa via nanofibers [102]. The amino acid derivative contained the N-(fluorenyl-9-methoxycarbonyl) (Fmoc) group, which induces self-assembly via π–π stacking (Fig. 5a). A more systematic study of the use of aromatic–aromatic interactions in the formation of supramolecular peptide hydrogels has been performed, showing the universality of this approach [103]. Incorporation of a terminal carboxylic acid group allows tuning and cross-linking of fibers using different types of salts, and subsequent tuning of the gel moduli with salt and concentration to give moduli between 25 Pa and 75 kPa [104]. The method of preparation of the hydrogels also has an influence on their mechanical properties [21]. Slowly adjusting the pH to form homogenous gels elevates the storage moduli of the gels from 5 to 184 kPa. Simple functionalization of RGD peptides with an aromatic Fmoc group leads to direct cell adhesion to these 3D scaffolds [105]. Besides functioning as a 3D support for the gel, these small molecules can also reduce inflammation [102]: The biphosphonate component of the hydrogel developed by Yang and colleagues lowers into uranyl oxide toxicity. Topical application of the hydrogel to wound sites on skin of mice, previously administered with uranyl oxide, showed effective recovery of the skin to normal, showing the therapeutic effect of these gels.

Fig. 5
figure 5

Hybrid supramolecular hydrogel systems. (a) Chemical structures of the three components of Fmoc-gels that assemble via π–π stacking. (b) Chemical structures of hydrophobically modified glucosamines. (c) Self-assembly of a hydrogel in the presence of cells [107]. Copyright 2003 Wiley-VCH Verlag GmbH&Co. KGaA, Weinheim. (d) Hyaluronic acid cross-linking by host–guest interaction of cucurbituril and amines, and the incorporation of cells and dyes. Reprinted with permission from [108]. Copyright 2012 American Chemical Society

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Yang et al. also showed that glucosamine (a naturally occurring precursor for glycosaminoglycans, proteoglycans, and glycolipids) modified with naphthalene can form a hydrogel (Fig. 5b) [106]. After cooling a solution of the modified glucosamine to room temperature, ribbons are formed. These subsequently form large bundles and tangle to make a 3D hydrogel with a storage modulus of 1 kPa. This hydrogel assisted in wound healing and prevented scar formation, as shown in a mouse animal model.

Gels were also shown to form when biotinylated poly(lactic acid) (PLA)–PEG microparticles were cross-linked using avidin (Fig. 5c) [107]. Gels can be formed that can encapsulate cells, although in rheology the gels do not behave as viscoelastic solids and show a large frequency dependence of their shear moduli. The microparticles can be used as conventional carriers for drugs or signaling molecules, providing a well-understood sustained release. Bone defect in femurs of chicken embryos were filled with these materials and the viability of the bone could be shown, proving applicability in regenerative medicine.

Hyaluronic acid has been modified with cucurbit[6]uril (CB[6]) and polyamines to form stable hydrogels with shear moduli of around 3 kPa when mixed [108]. The CB[6] can then also be used to include functional molecules in the hydrogel in a supramolecular fashion (as demonstrated with dyes to allow in vivo imaging) without influencing the mechanical properties of the hydrogel (Fig. 5d). Subsequently, RGD was incorporated by conjugation to CB[6] to promote cell adhesion of encapsulated cells in these gels, showing that this gel is a promising candidate for in vivo cell delivery. The gel can be formed under the skin of mice and functionalized in a supramolecular fashion inside the body.

The hybrid materials discussed here are still in their infancy and only initial experiments in regenerative medicine have been performed using them. Furthermore, not much knowledge has been gained yet on the mechanical tuning of these materials, except for their small peptide fragments. So far, the use of aromatic–aromatic interactions with small peptide fragments has shown promising properties for regenerative medicine, but more work is needed to demonstrate the applicability of these materials.

6 Synthetic Supramolecular Hydrogels

Besides functionalization of synthetic polymers with peptides or functionalization of peptides with synthetic cross-linkers, fully synthetic supramolecular systems have also been developed for use in regenerative medicine, although they have hardly been tested in vivo. Nature provides good inspiration for new synthetic biomaterials. We envision that many benefits can be gained by using completely synthetic systems, because such systems are infection-free, cheap, fully tunable, and scalable [5]. Before discussing the ureido-pyrimidinone (Upy) hydrogelators developed in our group, we discuss two classes of synthetic systems that are presently being developed for application in regenerative medicine but have not yet been tested in vivo. The first class consists of low molecular weight hydrogelators; the second class consists of hydrogels made from macromolecular monomers that are cross-linked via supramolecular interactions.

Low molecular weight hydrogelators form one-dimensional (1D) stacks and, subsequently, gel networks in water. The behavior of these molecules in organic solvents is well known, and the ability to tune these materials lies in the power of the chemistry to be scalable. Therefore, we consider these materials promising for application in regenerative medicine and give a few examples.

van Esch and coworkers developed hydrogelators based on 1,3,5-triamide cyclohexane with amino acid derivatives and water-soluble chains in the periphery [109]. These molecules interact via hydrophobic interactions and hydrogen bonds to form long 1D stacks, which at low concentration form hydrogels. A similar C3-symmetric molecule, benzene-tricarboxamide, has been developed that forms a hydrogel via a similar mechanism [110]. Lee and coworkers synthesized coordination polymers by functionalization of pyridine ligands with dendritic tri(ethylene glycol) chains, which lead to helical chains after the addition of silver ions, forming a 3D network [111].

The second class consists of polymers that are cross-linked by different moieties. Cyclodextrin (CD) has been widely studied as a supramolecular cross-linker for hydrogels [112]. The cavity of β-CD can include hydrophobic molecules as guests. For example, adamantane can be grafted onto poly(acrylic acid), which can form a complex with poly(acrylic acid) functionalized with β-CD to yield a supramolecular hydrogel [113]. This versatile approach allows the use of different polymers and different guests. Some polymers that have been cross-linked to form hydrogels using this complex are chitosan [114], poly(acrylamide)s [115], and polyacrylates [116]. Other guests that have been used include cholesterol [117], alkyl chains [118], azobenzenes [119], and ferrocenes [120], to name a few. However, not many systems containing CD have been used in vivo because of toxicity and reduced performance [112].

Another popular supramolecular host used to form supramolecular hydrogels is cucurbituril (CB), as described above. The big CB[8] molecule allows encapsulation of two hydrophobic guests. For instance, naphthol and methyl-viologen attached to PEG can be complexed inside CB[8] to form cross-links via a ternary complex, leading to gels with plateau modulus of around 10–1,000 Pa [121]. Also a short butane linker has been used to connect napthol and methyl-viologen to prepare supramolecular hydrogels [122]. Although the CB[8] and the guest show limited water solubility, encapsulation of the guests leads to enhanced water solubility and, subsequently, to formation of a hydrogel.

The synthetic systems described above have not been tested in vivo and their mechanical properties have not been determined. In our group, bifunctional telechelic hydrogelators were developed based on PEG cross-linked with UPy groups at both ends of the polymer (Fig. 6a) [123]. In dilute solution, stacking of UPy dimers results in the formation of nanofibrous structures, aided by additional urea hydrogen bonds in a hydrophobic pocket formed by long alkyl spacers. Above a critical concentration, bundling causes formation of a transient network (Fig. 6b).

Fig. 6
figure 6

Ureido-pyrimidinone (UPy)-based hydrogelators. (a) Chemical structure of telechelic UPy-functionalized PEG. (b) Chemical structure of monofunctional UPy-functional PEG monomethylether. (c) (1–4) Formation of fibers and subsequent bundling, leading to a 3D hydrogel [131]. Copyright (c) Wiley-VCH Verlag GmbH & Co. KGaA, Weinheim. (d) Chemical structures of chain-extended UPy- modified PEGs [127]. (e) Chemical structure of RGD-peptide functionalized with the UPy moiety [128]

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Analysis of the telechelic hydrogel with rheology and small angle X-ray scattering (SAXS) has shown that gel formation coincides with the formation of hydrophobic domains, where the fibers formed by the UPy stacks probably bundle to form cross-links [124]. The storage modulus and the relaxation time depend on concentration, temperature, and pH. By increasing the concentration, higher moduli and slower relaxation times are obtained, whereas increasing the temperature induces a switch from a hydrogel to a viscous liquid. The change from a viscous liquid to a hydrogel by increasing the pH allows the use of this system as an injectable delivery system with shear moduli of 100 Pa to 10 kPa. Interestingly, the addition of small monofunctional oligo(ethylene glycol) chains (modified only on one side with a UPy group) to the bifunctional UPy-hydrogelator gives a stronger gel than that formed from the individual components (Fig. 6c). In this mixture, the transient network is vitrified [125]. The synergetic network formed shows a storage modulus of 20 kPa, corresponding to a tenfold increase over that of the bifunctional hydrogelators at the same concentration. After mixing, the gel loses its responsiveness to temperature but is still responsive to pH. By incorporating UPy groups in the main chain of a PEG polymer, the hydrogels can be extended further while retaining the storage moduli (Fig. 6d) [126, 127]. These hydrogels show tensile moduli of 480–920 kPa, reaching almost the modulus of articular cartilage. This opens the opportunity for carefully designed gels by mixing-and-matching of the components for specific applications.

Although not yet applied in hydrogel materials, several UPy-modified peptides have been synthesized, such as UPy–RGD and a UPy–collagen I binding peptide, which in principle can be used for bioactivation of these hydrogels (Fig. 6e) [128]. These UPy–peptide conjugates have already proven their value in activating solid supramolecular polycaprolactone–UPy-based materials for regenerative medicine [129, 130].

Bifunctional UPy–PEG hydrogels have been applied to deliver the growth factor BMP7 under the kidney capsule in rats, and are able to promote fibrous tissue formation by diminishing the amount of infiltrating inflammatory cells and myofibroblasts. The bifunctional hydrogels have also been used for minimally invasive delivery of growth-factor proteins IGF and HGF in the infarcted myocardium in a porcine model using a long thin catheter. This was possible as a result of the pH responsiveness of the hydrogelators. At basic conditions with a pH of approximately 8.5–9, the UPy-hydrogelator formed a solution, whereas at the site of injection, the hydrogelator solution was neutralized to form a hydrogel drug delivery depot in situ [131]. Delivery of the growth-factor proteins stimulated endogeneous cardiac regeneration, that is, diminished collagen deposition, improved cardiac function, reduced hypertrophy, improved cell proliferation, and promoted formation of more capillaries [131, 132].

Although synthetic supramolecular hydrogelators offer several advantages over biomimetic systems and could be an interesting addition to the field, not many fully synthetic systems have been tested in vivo. This might be partially due to the easy availability of non-supramolecular conventional synthetic hydrogels, which have a proven track-record in medicine. However, as new systems are being developed, cheaper and scalable hydrogelators that more closely mimic the ECM can become available. We have shown that, with supramolecular hydrogels, it is possible to obtain good control over the mechanical properties (both strength and dynamics). Furthermore, bioactivity can be conveniently introduced into supramolecular hydrogels by using a mix-and-match approach.

7 Perspective

This chapter demonstrates that supramolecular hydrogels are promising materials for use in regenerative medicine. The examples show how a modular approach allows convenient tuning of mechanical properties and controlled addition of bioactive compounds. The use of the resulting soft dynamic and bioactive materials can provide a more natural-like environment for cells interacting with the biomaterial. Mimicking the native environment, and more specifically the ECM, has inspired a large body of the work performed in this field. However, the ECM is much more complex than the currently used biomaterials. There is not a single supramolecular biomaterial yet that is able to perform in a similar way as the complex set of components in the ECM. For example, the ECM consists of different components that are involved in resisting tensile, compressive, and shear forces, whereas in general only one of these is measured and considered for the hydrogels discussed in this chapter. Furthermore, many biological tissues show strain stiffening, which is essential for their physiological function [133], but nonlinear behavior is observed and studied only in a limited number of synthetic systems, such as polyisocyanopeptides hydrogels [134]. In addition, although short peptide sequences such as cell-adhesive RGD peptides offer an easy way to incorporate cell-attachment sites in biomaterials, these peptides are not very specific for inducing cell signaling. Fortunately, the use of supramolecular systems allows mixing of different (bioactive) components, both naturally derived and synthetically produced. This can lead to more complex systems offering all the structural and bioactive components required for specific functions. Although many systems are in the developmental phase, more translational research is required to show the true applicability of these systems in vivo.