Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https://creativecommons.org/licenses/by/4.0/).
Natural polymers, based on proteins or polysaccharides, have attracted increasing interest in recent years due to their broad potential uses in biomedicine. The chemical stability, structural versatility, biocompatibility and high availability of these materials lend them to diverse applications in areas such as tissue engineering, drug delivery and wound healing. Biomaterials purified from animal or plant sources have also been engineered to improve their structural properties or promote interactions with surrounding cells and tissues for improved in vivo performance, leading to novel applications as implantable devices, in controlled drug release and as surface coatings. This review describes biomaterials derived from and inspired by natural proteins and polysaccharides and highlights their promise across diverse biomedical fields. We outline current therapeutic applications of these nature-based materials and consider expected future developments in identifying and utilising innovative biomaterials in new biomedical applications.
Keywords: biomaterial, scaffold, tissue engineering, drug delivery, collagen, gelatine, silk, cellulose, chitosan, alginate
Traditional biomaterials used in biomedicine, such as gelatine, silk and collagen, were derived from natural sources [1], with their first clinical applications dating to the 1950s [2]. While they have had enormous impact on patient quality of life to date, they are being continuously modified, exploiting advances in the fields of molecular and cellular biology and polymer chemistry, to improve their material properties, bioactivities and suitability for therapeutic applications [3]. Furthermore, as biomaterials expand into new applications such as drug delivery, tissue engineering, scaffolds and bioprinting [4], new and modified materials are being developed that can remain in intimate and productive contact with tissues in the body for long periods [5].
Biologically inert materials were originally favoured for biomedical applications on the basis of safety and stability. Years of clinical use have identified that even inert mate-rials may elicit damaging cellular and immunological responses, however [6,7]. As a re-sult, biomaterials must now, at a minimum, interact with their surrounding tissues, while in more advanced applications, they may be designed to interact with surrounding cells and tissues to promote tissue healing and regeneration. Biomaterials that are biologically inert and passive are augmented with drugs, growth factors or gene delivery vectors to manipulate cellular responses in vivo for greater therapeutic effect [8,9].
This review describes the broad range of current biomaterials that are derived from, or inspired by, natural proteins and polysaccharides. We summarise the diverse biomedical and biotechnological roles successes of these materials to date, in fields such as regenerative medicine and therapeutics. Finally, we consider potential future directions for the development and modification of novel biomaterials with broader and more effective biomedical applications.
Choosing a suitable biomaterial for, e.g., scaffold construction is critical for its success. In vivo target sites at which biomaterials are used differ greatly and include soft (ligament, skin, cartilage, muscle, nerve, tendon, vascular sites) [10] and hard (bone and teeth) [11] tissues with very different biological and physicochemical properties [12]. Traditional alloys, metals and ceramics with limited functionality are increasingly being replaced by more versatile materials, while polymeric biomaterials are replacing permanent prosthetics due to concerns about the lower biocompatibility of the latter and the need for revision surgeries [13]. Increasing numbers of polymeric biomaterials are emerging from synthetic and natural sources, with diverse applications in tissue engineering and regeneration, as well as in specialist fields such as drug delivery, nanotechnology and gene therapy [14].
Biodegradable polymeric biomaterials can be divided into natural and synthetic materials, based on their origin and whether they contain naturally occurring extracellular matrix (ECM) [15]. Commonly used synthetic polymers, such as polyglycolide (PGA), polylactic acid (PLA) and poly(lactic-co-glycolic acid) (PLGA), are often cheaper to produce at scale and have more homogeneous structures, mechanical strengths and degradation rates [16]. They lack natural binding sites for cells, however, resulting in lower bioactivities and increased in vivo rejection rates than natural biomaterials [17]. Natural biomaterials are sub-classified based on their composition and include protein-based materials such as collagen, gelatine, silk and fibrin, and polysaccharide-based materials such as cellulose, chitosan and alginate [5,15], as described in detail below.
Scaffolds that act as templates for tissue regeneration and guide the development of new tissues can be produced from a diverse range of materials ( Table 1 ) [18]. Natural ECM can act as a porous 3D microenvironment “scaffold”, with its multitude of growth factors, effector molecules, enzymes and cellular adhesion motifs influencing cell proliferation, gene expression and intracellular signalling [19,20,21]. Due to the complexity and variability of ECM, however, scaffolds used in tissue engineering are typically more tailored to particular functions, such as promoting cell adhesion or differentiation [22,23]. Regardless of the clinical objective or environment, all biomaterial scaffolds share the following essential properties.
Summary of the different protein and polysaccharide-based biomaterials, their sources, main properties, structural forms used in biomedicine and biomedical applications.
| Material | Source | Properties | Structures | Biomedical Applications |
|---|---|---|---|---|
| Collagen | Natural ECM or recombinant | Weakly immunogenic Cross-linked to increase strength, stability Cell binding | Scaffolds Sponges Hydrogels Films/membranes Bioinks | Tissue repair [34,35,36,37,38] Wound care [39,40,41] Drug delivery [42,43,44,45] |
| Gelatine | Bovine or porcine collagen | Biocompatible Biodegradable Cross-linked to increase strength, stability Cell binding | Microparticles | Drug delivery [46,47,48,49,50,51,52] Tissue regeneration [53] |
| Silk | Butterflies/moths, spiders or recombinant | High strength-to-density Insoluble in water Highly stable | Films Woven meshes | Wound dressings [22,54,55] Suturing [56,57] Device coatings [58,59] |
| Cellulose | Plants, bacteria | Biocompatible Combine stiffness and flexibility Tuneable properties | Nanofibres Gels Nanocrystals | Tissue engineering [60,61] Artificial blood vessels [62,63,64] Drug delivery [65,66,67] Wound repair [68,69,70,71,72,73,74,75,76,77] |
| Chitosan | Exoskeleton of crustaceans; plant cell envelopes | Rigid structure Insoluble in water Biodegradable Antimicrobial Versatile | Films Fibres Scaffolds Hydrogels Nanoparticles | Wound healing [78] Anti-microbial coatings [79,80] Drug delivery [81,82,83,84] |
| Alginate | Brown algae | Widely available Inexpensive Biodegradable Excellent gelating | Hydrogels Sponges Films Microparticles | Wound healing [85,86,87,88,89] Drug delivery [90,91,92,93,94,95] Tissue engineering [94,96,97,98,99,100] |
Scaffolds should allow or promote cells to adhere, proliferate and spread before creating a new matrix, and must not elicit an inflammatory response that could lead to infection, longer healing times or patient discomfort [18].
Scaffolds should ideally be temporary templates which will be replaced by newly regenerated tissue [24]. Therefore, they should be biodegradable, resulting in non-toxic breakdown products which are safely excreted without interfering with normal bodily functions [18].
Scaffolds must be highly porous to promote cell migration, waste dispersal, scaffold-tissue interaction and nutrient and fluid permeability [25]. Cell-binding ligands may be naturally present in ECM-derived scaffolds or incorporated into synthetic materials [18]. Pores must be large enough to allow infiltration of cells but small enough to establish a suitable cell density attached to the scaffold [18,26].
Scaffolds must have sufficient mechanical strength to maintain their structural integrity—including during transport, surgical handling and implantation [27,28,29]. Engineered scaffolds should typically mimic the mechanical properties of their target tissue [27] and therefore vary greatly between, e.g., soft tissue applications and bone or cartilage scaffolds [30,31].
Scaffolds must be cost-effective to produce and easily scalable from laboratory production. Production must also be compatible with good manufacturing practice (GMP) standards [32] for translation from the laboratory to the clinic [33].
In nature, an array of proteins play vital structural roles in living organisms, which has led to their incorporation in recent years into protein/polypeptide-based biomaterials based on their structural chemistry, cellular interactions or cell communication properties [101]. Non-structural proteins are also gaining attention due to their ability to modulate the functional properties of biomaterials. In this section, we review protein-based materials derived from natural sources and consider their increasing impact in biomedicine.
Collagen is the most abundant structural protein in humans and animals. It makes up approximately 30% of all mammalian proteins and is an essential component of the ECM [102]. By virtue of its characteristic fibrillar structure, it provides structural support to hard and soft tissues, including cartilage, tendon, bone, ligament and blood vessels [103].
The collagen family consists of 29 distinct collagen types which are divided into four classes based on their composition and structural properties [102]. All types exhibit a characteristic triple helix structure, consisting of three α-chains comprised of more than 1000 amino acids and a repeating Gly-X-Y sequence. The glycine residues allow tight inter-molecular packaging of the α-chains while the X and Y positions are typically filled by proline and 4-hydroxyproline, respectively [104]. Of the 29 types, only types I, II, III, V and XI are known to form collagen fibres [103] and these are favoured in collagen-based biomaterials [105].
Collagen is weakly immunogenic, with fibrillose collagen exhibiting lower immunogenicity than smaller molecules due to the burial of potentially antigenic sites during its auto-polymerisation [103]. Removal of non-helical regions and cross-linking of collagen chains can be carried out to further reduce antibody recognition [106]. Cross-linking also provides stability and increases resistance to collagenase activity—properties which contribute to its biocompatibility and degradability in biomedical applications [103,107].
Specific peptide sequences in collagen bind four different types of cell-surface receptor. GPO (Gly-X-Y) motifs, where X and Y can be any amino acid residue but are most commonly proline [108], bind type 1 receptors, which includes glycoprotein VI [109]. GFO (Gly-Phe-Hyp) inter-act with type 2 receptors, which consist of collagen-binding integrin proteins and discoidin domain 1 and 2 [110,111]. Small molecule-binding cryptic domains in collagen [112] bind various integrins in type 3 receptors, while type 4 cell receptors bind non-collagenous domains within the protein [113]. Indirect cell-collagen interactions also promote cellular adhesion to the ECM, often via fibronectin (an ECM glycoprotein). Fibronectin was the first molecule on which the integrin-binding sequence RGD (Arg-Gly-Asp) was identified, which has since been found in many protein types, explaining their ability to bind collagen [103,114]. Receptor-binding motifs make an important contribution to the success of collagen as a biomaterial scaffold by aiding in cell seeding, adhesion and promoting cell differentiation and migration.
Collagen can be obtained for biomedical applications from mammalian sources such as cows, pigs, rats and sheep [105,115], as well as human peripheral nerve tissue [116] or human placenta [117]. It is purified by decellularisation, in which cellular antigens from the natural collagen matrix are removed while maintaining the ECM shape and structure, or by an extraction, purification and polymerisation approach which yields more refined scaffolds [103]. Decellularisation can be achieved by physical, chemical and enzymatic processes [118]. Physical decellularisation disrupts cell membranes and promotes cell lysis by the use of rapid freezing or high pressure approaches, and may be combined with chemical steps to increase tissue penetration [118,119]. Chemical decellularisation involves the addition of alkalines, acids, detergents and chelating agents to remove the cellular components of ECM while enzymatic decellularisation utilises the proteolytic enzyme trypsin, nuclease enzymes and ethylenediaminetetraacetic acid (EDTA) to remove proteins and DNA and RNA [103,118]. As none of the methods yields an ECM entirely free of cellular material, several techniques are typically combined to produce a pure, decellularised ECM [103] suitable for use as ligament prostheses or cardiac valves.
Extraction and purification of collagen from natural sources can also be carried out by solubilisation and purification. Due to its covalent cross-linking, collagen exhibits low solubility in organic solvents and is typically solubilised in acidic (0.5 M acetic acid), neutral salt (0.15–0.20 M NaCl) or proteolytic solutions.
Unlike decellularised collagen which has been naturally polymerised in vivo, extracted collagen must be cross-linked to increase its mechanical strength and resistance to enzymatic degradation. There are several well-established methods of cross-linking, including physical processes such as ultra-violet (UV) or thermal treatment [120], via chemicals such as formaldehyde and glutaraldehyde [121], and by the use of cross-linking enzymes such as transglutaminase [122]. The addition of biomolecules such as elastin [123], chitosan [124] and glycosaminoglycans (GAG) [125,126] during cross-linking can be used to improve the cell differentiation, migration, or proliferation characteristics, or mechanical properties of the resultant scaffolds [103,107].
Whether purified via decellularisation or solubilisation, extraction and cross-linking, collagen must be sterilised for in vivo use. Due to its relatively fragile, temperature-sensitive structure, sterilisation methods can alter its molecular properties. Even low-dose gamma irradiation reduces its enzymatic resistance and mechanical strength, though the addition of glucose can mitigate these effects by mediating cross-linking [127]. β- and electron-beam irradiation, while less damaging than gamma irradiation, have still been reported to cause structural degradation, and reduced mechanical strength and resistance to enzymatic degradation [103]. Immersion of collagen in low concentrations of sterilant is becoming a popular sterilisation approach, with low concentration pancreatic acid effective for decellularised collagen, and an ethanol/antibiotic mix for extracted, cross-linked collagen. As no approach can completely avoid altering the collagen structure, however, it is essential to thoroughly investigate the effect of each method on the properties of the resultant material relative to its intended application [103].
As collagen is primarily obtained from animal sources, the transmission of infectious disease is a concern in therapeutic applications. The recent COVID-19 pandemic exemplifies the dangers of zoonotic infections, while cases of bovine spongiform encephalopathy have resulted from the use of prion-contaminated bovine scaffolds [128]. Meanwhile, religious constraints surround the use of porcine and bovine materials, and up to 2%–4% of the world’s population may be allergic to porcine and bovine-derived collagen [129]. These factors, as well as the heterogeneity of natural collagen preparations, have led to the development of approaches to produce recombinant human collagen (rhCOL) [130,131].
Escherichia coli is the best-established expression system for recombinant proteins due to its ease of genetic manipulation, rapid growth rate, track record in protein production and suitability to scale-up [132]. It can produce a protein similar to human collagen which contains the characteristic Gly-X-Y sequences but differs from the native collagen through its lack of proline and lysine hydroxylation [133]. The failure of E. coli to carry out these post-translational modifications (PTMs) results in a collagen with reduced thermostability and limited fibre assembly, thereby restricting its usefulness in tissue engineering [131]. Cloning of prolyl and lysyl hydroxylase genes from the aquatic giant Mimiviridae virus family into E. coli has enabled the successful production of molecules with hydroxyproline and hydroxylysine patterns that are characteristic of human collagen and capable of supporting the growth of human endothelial cells [134], but despite this success, eukaryotic systems continue to dominate recombinant production of collagen.
Since the expression of human interferon in Saccharomyces cerevisiae in the 1980s [135], yeasts have been very successfully used in recombinant protein production. S. cerevisiae and Pichia pastoris are the two most commonly used yeast hosts [136] and their utility is due to their eukaryotic protein folding mechanisms and ability to carry out PTMs required for many proteins’ functioning. Like prokaryotic systems, they lack native lysyl and propyl hydroxylases [132] but co-expression of human hydroxylases enables them to produce rhCOL closer in structure and properties to native human collagen that that expressed in E. coli [132]. rhCOL from both S. cerevisiae and P. pastoris has been used to produce hydrogels for wound healing applications [131].
A variety of mammalian systems have also been investigated for the accurate production of human collagen. Chinese hamster ovary (CHO) cell-derived rhCOL was shown to reverse the disease phenotype of dystrophic epidermolysis bullosa (characterised by collagen deficiency within the skin) when administered intravenously, without eliciting an immune response in mice models of the disease [137]. Human HeLa cells [138] and embryonic kidney cells [139] have also been used to produce rhCOL types I, V and VII identical to native human collagen produced in vivo. Yields are much lower than from other expression systems, however, so non-human animal platforms are typically preferred. rhCOL has also been produced in transgenic animals, with mouse embryos transfected with a COL1A1 gene found to secrete correctly folded rhCOL type I through their mammary glands [140]. Transgenic animal production systems are considerably more specialised and expensive than cell-based systems, however [132].
Upon obtaining collagen, scaffolds of pure collagen, a collagen/natural polymer blend or a collagen/synthetic polymer blend can be produced [141].
Collagen types I, II and III have been electrospun into fibres at submicron to nanometre scale that replicate the biological properties of ECM. Collagen fibres with a 67-nm binding pattern, similar to native collagen, have been formed from electrospinning collagen type I [142], leading to the production of biomimetic scaffolds with tuneable porosity, mechanical strength and fibre alignment to topographically guide tissue formation. Electrospun collagen scaffolds have also been demonstrated to support cellular growth [142,143,144]. Pure collagen scaffolds have weak structural stability and mechanical strength, however. While this can be improved by cross-linking fibres using UV irradiation or dehydrothermal treatment (high temperature exposure under a vacuum), care must be taken to ensure no residues of toxic cross-linking chemicals such as glutaraldehyde remain in the scaffold [141].
Scaffolds can also be constructed by blending natural or synthetic polymers with collagen for improved mechanical properties. Natural polymers already established in tissue engineering include fibroin, chitosan and silk. Chitosan is a non-immunogenic, biodegradable, positively charged polymer which has been combined with collagen to form scaffolds with excellent mechanical and biological properties, as well as excellent compatibility with a range of seeded cell types [145,146,147,148]. Synthetic polymers utilised in blended collagen scaffolds include PLA and polyethylene glycol (PEG) [141]. In this scenario, the synthetic polymer typically improves the mechanical properties and structure of the scaffold while collagen provides cell signalling and binding sites crucial for tissue repair.
Collagen sponges are produced from insoluble collagens extracted from cows or pigs. The scaffold is created by freeze-thawing alkali and aqueous acid collagen solutions containing up to 5% dry matter, with the rate and temperature of freezing determining the pore size and structure: rapid freezing at extremely low temperatures cause cracking to occur in the collagen, resulting in small channels and a highly fibrous structure, while slower freezing and higher temperatures causes the collagen to have large, non-uniform pores and more continuous channels [103].
Sponges are ideal for use in wound care as they adhere smoothly to the wound bed, are capable of absorbing large volumes of exudate, maintain a moist environment and shield against physical trauma and bacterial infection [104]. Cross-linking with glutaraldehyde or other polymers can be used to increase their mechanical strength. As collagen promotes invasion of immune/inflammatory cells such as neutrophils, sponges have potential uses in treating burns, diabetic ulcers and at donor sites [39]. Loading of sponges with exogenous growth factors can also be used to improve wound healing, such as platelet-derived growth factor (PDGF) and fibroblast growth factor (FGF) to promote capillary formation and epidermal wound healing, respectively [40,41]. Collagen sponges have also been used to provide sustained delivery of antibiotics such as vancomycin [42] and gentamycin [43] to treat sepsis, and intra-vaginal delivery of retinoic acid to avoid systemic effects in the treatment of cervical dysplasia [44].
A hydrogel is a 3D network of polymers which can hold significant volumes of fluids. The main therapeutically-relevant properties of polymeric hydrogels are their water retention, due to the hydrophilic functional groups on their polymeric backbone, their resistance to dissolution due to their cross-linking and their similar flexibility to natural ECM [149]. Due to their structural similarity to tissue, collagen hydrogels are frequently investigated as biomimetic 3D scaffolds to support cell growth [150].
The amphoteric (adsorbs to both anions and cations) nature of collagen type I fibres enables them to form hydrophobic, dipole-dipole, electrostatic and hydrogen interactions which lead to gel formation in aqueous systems [150]. The gels can typically be dissociated by collagenases or changes in temperature or pH [151]. While the natural cross-linking of collagen confers proteolytic resistance and mechanical strength, additional physical or chemical cross-linking can be necessary to prevent enzymatic degradation [152]. Of these approaches, glutaraldehyde cross-linking via lysine and hydroxy-lysine residues [153], sometimes used in combination with 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) cross-linking of carboxyl and amine groups [150], yields the most stable hydrogels and can be designed to vary the mechanical properties and function of the hydrogel [150], though some cytotoxicity has been reported [154].
Figure 1 illustrates non-cross-linked collagen compared with glutaraldehyde cross-linked and EDC cross-linked collagens.
Collagen hydrogels: (a) non-cross-linked, (b) glutaraldehyde cross-linked and (c) EDC/NHS cross-linked [150].
Collagen hydrogels are attractive scaffolds in tissue engineering due to their retention of cells and bioactive molecules. They have found widespread application in cartilage and bone tissue engineering, such as acting as carriers for bovine chondrocytes [34] to provide structural support and pain-free articulation of cartilage [35]. As chondrocytes produce cartilage ECM, their transplantation in collagen hydrogels can be used to treat a variety of articular cartilage defects [36] or to engineer bone tissue [37]. Osteoblasts derived from calf metacarpus periosteum, have also been demonstrated to proliferate and migrate within a 3D collagen hydrogel without any loss of viability over three weeks, as well forming a bone-like ECM containing osteonectin, osteocalcin and new collagen type I [37]. Collagen type I hydrogel-mediated treatment of bone defects in rat dorsal nasal bones was demonstrated to lead to the growth of a thin layer of bone after six weeks [155]. Blending of synthetic polymers such as polyvinyl alcohol (PVA) and polyacrylic acid (PAC) with collagen in hydrogels can be used to improve the mechanical strength of the natural polymers and the biocompatibility of the synthetic molecules [1]. In one such example, PVA was blended with collagen and used to form sponges, films and hydrogels, which were loaded with growth hormone. Release of the hormone was monitored in vitro by enzyme-linked immunosorbent assay and could be tuned by altering the collagen content of the gels [156].
Collagen type I is commonly used as in bioprinting but its slow gelation rate at physiological temperatures means that it must typically be used in combination with biomolecules that improve its structural integrity [4] ( Figure 2 ). Collagen-alginate bioinks have been demonstrated to have increased mechanical strength and to accelerate the proliferation of human chondrocytes for articular cartilage repair, [38] while printed collagen-alginate hydrogels have also demonstrated sustained release of antibacterial drugs [45].
Bioprinting of natural polymers, frequently in combination with cells and/or biomolecules to fine-tune or increase in vivo activity, has potential to provide carefully designed, highly structured materials for tissue and organ engineering applications.
Collagen films of 0.01–0.5 mm thickness can be produced by drying bovine collagen that has had its telopeptides (nonhelical regions that flank collagen’s triple helix) removed, followed by a series of enzymatic and chemical cross-linking steps [39,157]. Their main applications are as barriers to protect wounds or ulcers, while simultaneously releasing therapeutic drugs. Drugs can be loaded onto films via covalent or hydrogen bonding or through simple entrapment and the films can be easily sterilised without affecting their mechanical properties [39]. Collagen films are well established in wound dressing applications, as well as reinforcing compromised tissues and guiding tissue regeneration. They have also been used to deliver antibiotics [158,159] and, as collagen-coated polyurethane (PU) films, to promote attachment and proliferation of fibroblasts [160], which stimulates further collagen synthesis and the formation of new connective tissue and ECM. Individual films can also be easily combined into multiple-layered membranes which can release molecules such as PDGF at constant rates for up to 100 h in vivo to aid wound healing and tissue regeneration [107], or human growth factors to support healing of diabetic ulcers in murine models [161]. Collagen-based scaffolds have also been utilised as a resorbable template, alone or with additional components such as hyaluronic acid [162] or synthetic polymers [163,164], to regenerate the meniscal template of the knee, with promising in vivo outcomes [165].
Collagen-based materials are currently to the fore amongst biomaterials used in regenerative medicine and tissue engineering, based largely on their low immunogenicity, high biocompatibility and structural versatility. Advances in extraction and scaffold formulation have led to increasingly diverse applications of collagen in fields such as wound healing, drug delivery and tissue regeneration. Future research is likely to focus on improving the mechanical strength, drug delivery capabilities and biodegradability of collagen-based scaffolds in order to enhance their in vivo efficacies.
Gelatine is a well-characterised, biocompatible and biodegradable polymer which is formed by disintegration and denaturation of natural collagen, typically of bovine or porcine origin [166]. It is commonly utilised in food and cosmetic production as a cheaper alternative to collagen and has extensive pharmaceutical applications. It is a derivative of type I fibrillar collagen and contains up to 92% pure protein, as well as mineral salts and water [46]. It exhibits several advantages over its parent collagen in therapeutic applications, including reduced immunogenicity [167], increased solubility in aqueous systems and ease of transition from solution to gel at temperatures of 30 °C [168].
The structure of gelatine depends on the source and method of denaturation of the parent collagen [168]. This leads to heterogeneity of gelatine and batch-to-batch variability in its molecular weight from a thousand to greater than a million Dalton [166,169]. Denaturing collagen forms a gelatine solution with very low viscosity, which forms a gel at temperatures below 37 °C. This reversible process of thermo-responsive gelation results from peptide coils transitioning into helices due to the high numbers of Gly-Pro bonds within the structure [170]. As the gels return to a liquid state at temperatures above 37 °C, this makes natural gelatine unstable for in vivo use, though numerous cross-linking approaches have been developed to stabilise the macromolecular structure and avoid its rapid degradation in host tissues [169], as outlined below.
Mammalian gelatine is preferred for in vivo applications due to its high concentration of cell-binding domains, which create an excellent substrate for recruitment and attachment of adherent cells. Meanwhile, gelatine is less antigenic than collagen due to its lower composition of phenylalanine and absence of tyrosine and tryptophan which form aromatic rings and radicals that can promote an antigenic response [171].
Gelatine is produced by thermal denaturation and hydrolysis of collagen. Heating to only 40 °C disrupts the interior structure of newly formed or highly soluble collagen [169], whereas non-soluble collagen requires an additional hydrolysis treatment using acid or alkali solutions [169]. Soaking in dilute acidic solutions yields gelatine type A while submersion in alkali solutions produces gelatine type B, with the former more similar to collagen in isoelectric point and amino acid composition [172].
As with collagen, gelatine can be cross-linked using chemical, physical or enzymatic approaches. Physical cross-linking involves physical gelation via heat and pH changes as described above. To avoid the reversibility of physical gelation upon temperature or pH changes, chemical methods have been developed to produce a more stable polymer. Glutaraldehyde cross-linking is typically the preferred strategy as glutaraldehyde is easily accessible, cheap and rapidly boosts the mechanical strength of gelatine [151]. Chemical cross-linking can leave behind traces of potentially toxic chemical agents, however, leading to investigation of natural cross-linking agents such as caffeic acid, tannic acid, genipin and grape seed proanthocyanidin. Genipin, derived from the fruits of Gardenia jasminoides [173], is far less toxic than glutaraldehyde and achieves comparable mechanical properties in the cross-linked gelatine, but is limited by its high cost and formation of a dark blue pigment which can constrain its biomedical use [46,174]. Enzymatic approaches have been used to synthesise highly stable gelatine structures, with transglutaminases in particular favoured due to their abundance in nature and ability to produce a mechanically strong product [175].
Gelatine is favoured in biomedical applications due to its commercial availability, low price, high solubility, biocompatibility, biodegradability, the presence of cell-binding domains and its lack of antigenicity or toxicity to cells [46]. Its disadvantages, however, surround its poor mechanical properties, lack of thermal stability, greater susceptibility to some proteases than collagen and faster degradation [176]. To overcome these disadvantages, advanced manufacturing techniques and cross-linking approaches are used to improve the thermal and mechanical stability, biocompatibility and overall bioactivity of native gelatine.
The ability of gelatine to form a gel, its biocompatibility and its biodegradability make it ideal for the production of microparticles. Gelatine microparticles are extensively used as drug carriers due to their ease of production, stability and lack of toxicity, as well as their ability to interact with multiple bioactive compounds [177]. While smaller particles are used to protect and control the release of bioactive molecules in vivo [47], such as growth factors to stimulate cell proliferation and differentiation [46,48], larger microparticles, with modified surfaces for improved cell attachment and differentiation, can be used as “microcarriers” of cells, e.g., delivery of embryonic stem cells to aid bone regeneration [53].
Gelatine microparticles are traditionally produced by techniques such as solvent evaporation, spray drying and precipitation, though these may cause denaturation [47] or leave solvent traces in the final product [178] ( Figure 3 ). Therefore, improved production methods such as water-in-water emulsification have been developed [177].
Loading of gelatine microparticles with growth factors and/or cells to direct cellular differentiation (based on [46]).
