Introduction

According to the Organ Procurement and Transplantation Network, there are over 120,000 patients in the USA awaiting organ transplantation (as of 15 November 2015). Transplanting a tissue from one location to another within the same patient (an autograft) or from one patient to another patient (an allograft) is effective in many instances for replacing organs with functional outcomes. However, many problems exist with both procedures. Autograft techniques are costly, result in an additional site of injury, increase the risk of infection or hemorrhages, and are limited by anatomically incorrect replacements from other regions of the body [1]. Allografts on the other hand, are often rejected by the immune system (requiring immunosuppressant therapies), and pose risks of infection or transfer of diseases between patients. As an alternative, there has been growing interest in designing replacement organs that use mechanical, physical, and/or biological components to restore or replace function. In particular, tissue engineering historically has involved the concept of isolating cells from a patient, expanding them in vitro, and seeding them onto a biomaterial that could be implanted in vivo in the site of injury [2]. The formation of artificial organs with tissue engineering techniques is necessary to address the growing disparity between the supply and need for donor organs, as well as the limitations associated with autograft and allograft techniques.

For use in tissue engineering regenerative applications, biomaterials should be biocompatible, porous (to allow cellular infiltration, nutrient transport, and waste removal), mechanically tunable (to maintain the intrinsic mechanical properties of the tissue through the healing process), biodegradable (to allow the tissue to develop as the material degrades), reproducible, easily prepared, and cell/tissue compatible. However, optimizing these biomaterial design criteria can result in distinct changes in regenerative potential (Table 1). Therefore, this review will focus on various biomaterial design considerations and their effect on regenerative outcomes (Fig. 1), keeping in mind that tissue-specific considerations [7] will also be required to choose the optimal biomaterial system.

Table 1 Effects of biomaterial characteristics on regenerative outcomes
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Fig. 1
figure 1

By adjusting biomaterial designs, including: the porosity and mechanical properties, the surface chemistry, the degradation kinetics, and combining biomaterials with cell- and tissue-specific factors (including soluble factors), intrinsic properties of tissue constructs can be controlled to enhance remodeling and functional outcomes

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Biomaterial Sources

Biomaterials that emulate the native extracellular microenvironment can support or direct the proper cellular phenotypic state. Generally, there are three different approaches for generating the scaffolds required to engineer a tissue (examples, see Table 2): (1) construct scaffolds from natural or synthetic sources [2, 10, 11, 14], (2) decellularize the tissue of interest leaving the extracellular matrix as the scaffolding material [12], or (3) stimulate cells to generate their own matrix [13]. Each method has advantages and disadvantages that make them more appropriate for certain circumstances based on their inherent biocompatibility, reproducibility, and ease of preparation.

Table 2 Examples of biomaterials used in regenerative applications
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Intuitively, the optimal biomaterial would be the native extracellular matrix of the tissue of interest [7]. Both decellularized tissues and cell-generated materials have specific ligands for cell adhesion, can be degraded and remodeled readily by cells, and have other tissue-specific cues consistent with the native organ, such as signaling epitopes and/or sequestered growth factors [12]. These matrices are naturally biocompatible, however are not easily prepared or reproducible, with relatively low yields of material. Therefore, generating scaffolds from natural or synthetic sources has some advantages over decellularizing tissues or stimulating cells to generate their own matrix. For instance, there is a greater supply of natural and synthetic sources from which to generate the scaffolds. Due to control of processing methods, there is also greater control over the resulting shape and related physical and chemical features of the scaffolds when compared with decelluarized options. Further, the process time required to obtain the material is often faster. Comparing synthetic to biological materials, synthetic materials have minimal lot-to-lot variability, are highly tunable (chemical and mechanical properties), are easier to obtain a purified form, and are generally more reproducible then natural sources [14]. However, natural sources of scaffold materials have advantages such as potentially enhanced biocompatibility, ease of chemical functionalization or sequestering of growth factors, in some cases biological recognition, and susceptibility to cell triggered proteolytic degradation and remodeling [14].

Tuning Mechanics

The appropriate mechanical characteristics of a biomaterial are highly dependent on the regenerative application. For example, mesenchymal stem cells differentiate into different lineages, such as neurons, myoblasts, and osteoblasts, on progressively stiffer substrates [15]. There are many ways of changing the mechanics of a biomaterial (depending on what type it is). Some common approaches include: crosslinking materials [16], controlling crystallinity during processing [17], and using inorganic reinforcing fillers (to enhance the stiffness of the matrix [18]). Greater precision in scaffold processing such as microfabrication techniques [19] and employing mechanical gradients within the same scaffold system (for soft-to-hard interfaces [20••]) are also being used to more closely mimic in situ architecture.

Porous Structures

Pores are necessary to allow cells and the nutrients they require to migrate and infiltrate into the bulk of the scaffolds. Pores can be formed by many different processing techniques, including: freeze drying, porogen leaching, gas foaming, microfabrication, and three-dimensional (3D) printers (for a review on processing techniques that result in differences in pore sizes see reference [21••]). While increasing the pore size of a scaffold increases infiltration of cells and improves transport, there is a trade-off with mechanical properties. For example, when biomaterials are implanted in vivo to promote bone ingrowth, higher porosity, and pore sizes result in greater osteo-integration with a reduction of mechanical strength [3]. Larger pore sizes also result in decreased integrin/ligand interactions between cells and the biomaterial [1]. On the other hand, decreasing the pore size significantly limits cell migration (surface growth versus bulk infiltration) and can lead to cellular encapsulation of the scaffold [1]. While larger pores and enhanced porosity increases nutrient transfer, this can also be achieved by adding channels to scaffolds [22–26]. In some tissue engineering constructs, these channels are lined with endothelial cells [27–30] to recreate the network of vasculature that provides nutrients in vivo. It is important to consider the dynamic nature of endothelial networks, however, which is dependent on the type of mural cells and the mechanical and bioactive properties of the matrix [31]. In addition, the interconnectivity of pores and adding channels to the materials, affects the mechanical properties where randomly oriented defects from both processing techniques soften the bulk properties of tissue constructs resulting in problems generating structurally robust tissues (compared with more ordered arrangements). Therefore, careful consideration should be taken when choosing the porous (and/or channel) structure of biomaterials.

Controlled Degradation

One of the goals of tissue engineering is to design the degradation and remodeling kinetics of a biomaterial to: (1) initially support seeded cells (and/or infiltrating cells) to proliferate and secrete matrix, and (2) gradually degrade allowing de novo tissue to fill the empty spaces left by the scaffold. In this way, paired scaffold degradation and tissue remodeling can maintain the intrinsic mechanical properties through the healing process. Unfortunately, each tissue requires unique degradation/resorption kinetics based on differences in matrix properties, cell types, and bioactivity of other factors in the extracellular matrix. An added confounder is that the regenerative capacity of patients varies, for example, children exhibit higher regenerative capacity and will likely require a faster degrading scaffold than adults that have a slower healing capacity [4]. Other variables aside from age include patient nutritional status, disease state, lifestyle, sex, and other related factors.

Materials can be biodegradable, bioresorbable, bioerodible, or bioabsorbable [32]. Biodegradable materials break down due to macromolecular degradation to byproducts that can move away from the implantation site but are not necessarily removed from the body [32, 33]. In contrast, bioresorbable materials break down due to bulk degradation and are eliminated through natural pathways (such as metabolism by cells) with no residual side effects [32, 33]. This is ideal in some circumstances, as complete removal of the biomaterial eliminates any remaining foreign body reaction. Alternatively, bioerodible polymers degrade and resorb only at the surface [32, 33]. Finally, bioabsorbable materials dissolve in body fluids without a decrease in molecular mass (a bioabsorbable material is bioresorbable if the byproducts are excreted). All classifications result in different regenerative outcomes, and are highly dependent on tissue-specific requirements.

An important consideration is whether the degradation byproducts will have toxic effects on cells and the surrounding tissue. To address degradation and remodeling concerns, it is important to consider the source of the material. Many synthetic polymers are biodegradable, for example, the co-polymer poly-d,l-lactic-glycolic acid (PLGA). However, the acidic environment generated by the degradation of PLGA is detrimental to cell mobilization and angiogenesis [8]. Proteins derived from the extracellular matrix, on the other hand, are naturally susceptible to proteolytic degradation and remodeling by cells. The byproducts of degradation are less likely to have toxic effects; however, the byproducts may cause downstream impact. For instance, byproducts of collagen degradation control cell migration of vascular smooth muscle cells which are responsible for de novo synthesis of specific extracellular matrix components (including restenosis after balloon angioplasty) [34].

An interesting approach to controlling degradation in polymers is with biological triggers. In this approach, labile chemical groups are conjugated into the backbone of polymers that can be degraded by physical or chemical signals, such as near-infrared light [35], ultraviolet light [36•], ultrasound [37, 38], redox [4], various enzymes [39, 40], and pH changes [41]. As an example, a peptide-conjugated polymer was generated, which underwent degradation when it was exposed to collagenase [39]. When designing these systems, the physical or chemical cue must be carefully designed to avoid off target effects. For instance, some enzymes, such as collagenase, can cause further damage to the surrounding tissue. Targeting local cell types that release the enzyme of interest (for instance MMPs [40]), or directly injecting biological triggers into the implantation site [4], are approaches that can be explored to reduce injury to the surrounding tissue.

Release of Soluble Factors

To improve biological activity, controlled biomaterial degradation can be paired with the release of soluble factors to enhance regenerative outcomes. Drugs have been incorporated in many biomaterial scaffolds, for example, to promote angiogenesis in ischemic tissues [42]. Another important area of research is the controlled delivery of growth factors, which are important mediators of cellular signaling, migration, proliferation, differentiation, and maturation [2]. To incorporate both drugs and growth factors in regenerative approaches, local and temporal control is required to improve clinical outcomes [43]. To accomplish this task, soluble factors have been delivered from particles (at the micro- and nano- scale), from scaffolds, and also encapsulated to release as the biomaterials are degraded [5]. Binding of growth factors and drugs to biomaterials can occur through covalent [44] or ionic bonds [45] and depends on the properties of the soluble factors and the biomaterial. Encapsulating growth factors or drugs helps to control the release profile and can protect the unstable bioactive molecules from degradation (and loss of activity). However, covalent/ionic bonds and encapsulation alone often lead to an initial burst release of soluble factors. Many approaches aimed at accomplishing sustained release and zero-order release have been proposed to address the burst issue (for reviews see: [46, 47]). For example, in one system [48], bone morphogenetic protein 2 was encapsulated in a nano-sphere which was encapsulated in a micro-sphere scaffolding system to enhance bone regeneration. The hierarchical structure minimized the initial burst from the nano-sphere by confining it within the scaffolding system of the micro-sphere. Other biomaterial approaches control temporal timing of factors to release different factors over time. For example, epidermal growth factor at early time points and mitomycin C at later time points [49].

Other Surface Chemistry Considerations

When cellular binding sites are not present on the material, cell adhesion ligands such as proteins, peptides, and oligosaccharides can be added to enhance cell adhesion. Coating with extracellular matrix proteins which are specific for certain tissues is commonly used to enhance adhesion, differentiation, and/or proliferation of certain cell types. For instance, fibronectin [50], collagen [51], and laminin [52], some of the most common protein coatings used, have all been recently used to enhance tissue-specific regenerative outcomes for periodontal ligament, smooth muscle, and tracheal epithelial tissue, respectively.

Coating with proteins has several drawbacks. Proteins must be isolated and purified from an organism, making large-scale production challenging, or they can be generated by recombinant DNA methods, which adds cost and purification challenges. They may also elicit immune responses and can be proteolytically degraded [6]. Decorating the materials with immmobilized peptides for cell recognition motifs, such as RGD (arginine (R), glycine (G), aspartic acid (D)) therefore can be used to counteract the issues associated with protein coatings since they are easily characterized, cost effective, exhibit higher stability, and can be packed at a higher density on the surface of biomaterials [6]. However, immobilized cell recognition motifs may have lower cell adhesion activity and are specific to only one cell adhesion motif (whereas matrix proteins have many) [53].

Future Considerations for Biomaterials Aiding Tissue Regeneration

Creative approaches to adjust material designs, including pore size and degradation kinetics, in combination with functionalization with cell- and tissue-specific factors will result in enhanced tissue remodeling and functional outcomes. Greater precision in scaffold processing is required to provide control over the structure of biomaterials to recreate tissue-specific features. Advances in 3D printing and nanotechnology development will likely be paramount for this precision. In particular, patient- and tissue-specific architectures and degradation kinetics will be a goal. In this way, scaffolds can be designed to match the geometry of the patient-specific defect and account for location, age, or disease state for degradation properties. In addition, it will be important to modulate specific cues spatially and temporally within a scaffold, for instance by coating specific regions with different extracellular matrix proteins or by pairing degradation with the release of specific soluble factors. This approach will be especially important in cases where multiple cell types or tissues exist within one organ. Likewise, creating gradients and accounting for interfaces (i.e., between soft and hard tissues) with different biomaterial processing techniques will be crucial. Hybrid materials that combine the advantages of different materials will also be essential in many applications. As an example, extracellular matrix/fibrin scaffolds were generated, which combined the highly angiogenic properties of fibrin with cardiac extracellular matrix to promote both vascular and cardiac differentiation within the same scaffold [54•]. This hybrid system also resulted in varying levels of mechanics, which will be an important consideration for biomaterials moving forward. As these new areas of biomaterial development are explored, standardization, good manufacturing practices, and cost-effective methods will be essential for translation toward clinical utility.

Conclusions

Emulating the extracellular microenvironment is essential for establishing proper cellular phenotypes required for tissue regeneration. Careful consideration of biomaterial choices in tissue engineering can improve regenerative outcomes. For instance, by making adjustments to the material design, including pore size and degradation kinetics, intrinsic properties of tissue constructs can be controlled to enhance remodeling and functional outcomes. Additionally, the extracellular matrix is a repository for a variety of growth factors and other signaling ligands, which can be incorporated in biomaterial approaches to improve regenerative outcomes.