Bioactive Polymer Scaffolds for Tissue Engineering
Bioactive Materials
Vol. 1, No. 2 December 2016 pp. 93-108
View original article here.
Author link overlay panel Scott Strattonac1 Sangamesh G. Kumbarabc
https://doi.org/10.1016/j.bioactmat.2016.11.001
Highlights
• The requirements necessary for an ideal biomaterial-based scaffold are highlighted.
• The relationship between mechanical properties and porosity is explained.
• The advantages and disadvantages of several scaffold fabrication techniques are discussed.
• Examples of tissue engineering strategies
Abstract
Various artificial scaffolds for tissue engineering have been created using polymers, ceramics, and their composites.
Biomimicry is employed in the majority of three-dimensional (3D) scaffold designs in terms of physicochemical properties and bioactivity for superior tissue regeneration. Scaffolds fabricated by salt leaching, particle sintering, hydrogels, and lithography have been successful in in vitro cell proliferation and in vivo tissue regeneration. Scaffold systems derived from the decellularization of organs and tissues have gained popularity due to their reliable biocompatibility and bioactivity.
Conventional scaffold fabrication techniques were often unable to create complex structures with higher resolution, had low reproducibility, and required multiple steps. 3D printing technology has overcome some of the limitations of conventional techniques, allowing for the easy creation of micro-nanostructured scaffolds and devices for tissue engineering and drug delivery by employing several thermoplastics and hydrogels.
This review focuses on scaffold fabrication methodologies with an emphasis on optimizing scaffold performance through matrix pores, bioactivity, and degradation rates to enable tissue regeneration.
This review highlights several examples of bioactive scaffolds that mediate the regeneration of nerves, muscles, tendons/ligaments, and bone. Regardless of the efforts required for optimization, if some of the methods covered in this review become more rationalized,it is expected that 3D scaffolds will transition from the laboratory to daily life in the near future.
Graphical Abstract
Keywords
Bioactive biomaterial scaffolds Porous Biodegradable Tissue regeneration
1. Introduction
Bioactivity refers to the ability of a material to influence its biological environment.
Langer and Vacanti first defined the term "tissue engineering" in the 1990s. Since then, three-dimensional (3D) structural biomaterial-based scaffolds have been used to provide a bioactive environment for cells to adhere and proliferate [1]. Researchers hypothesized that scaffolds could provide structural stability and an environment for cell regeneration, potentially mimicking native tissue functionally, thus paving the way for tissue regeneration. Since then, 3D scaffolds have been evaluated for various applications, including bone regeneration, nerve regeneration, muscle regeneration, and tendon/ligament regeneration [2], [3], [4].To realize these scaffolds, synthetic and natural polymers have become popular biomaterials due to their diverse properties and bioactivity [5], [6], [7].Natural polymers were among the first biodegradable scaffold materials used clinically due to their excellent interaction with various types of cells and lack of immune response. However, it was later found that synthetic polymers were cheaper and had better functionality than natural polymers, although there is a potential for immune response and toxicity, especially with certain polymer combinations [8]. Among synthetic polymers, poly(L-lactic acid) (PLLA), polyglycolic acid (PGA), polycaprolactone (PCL), and poly(lactic-co-glycolic acid) (PLGA) are currently the most commonly used to create 3D structures in the form of scaffolds [9], [10], [11]. These polymers are also used in combination with natural polymers to improve unfavorable issues related to hydrophilicity, cell adhesion, and biodegradability. Furthermore, scaffold surfaces are functionalized using specific ligands, such as protein molecules, to help enhance cellular responses.
Three-dimensional scaffolds such as nanofibers, hydrogels, and sintered microparticles have been widely studied using synthetic and natural biomaterials [8], [12]. These three-dimensional highly porous scaffolds are used to generate a local bioactive environment upon implantation to regenerate damaged or lost tissue.In 3D scaffolds, porosity has been shown to be a critical determinant for second-generation tissue engineering, andthe need for angiogenesis and cell infiltration into the pores within the scaffold is further emphasized [13]. Previous studies have shown that cell proliferation and adhesion are highly dependent on both the size and density of pores within the scaffold, and specific parameters must be carefully manipulated depending on the material and application [14], [15].
One of the main reasons why porosity is important is that cell networks rely on interconnected pathways for nutrient transport, cell signaling, and proliferation, structurally mimicking the native extracellular matrix (ECM) environment. However, the porosity of the scaffold, and consequently the surface-to-volume ratio, should not be so large as to weaken its mechanical strength [16]. 3D scaffolds for tissue engineering applications face two major constraints: the scaffold must not be too porous (as mechanical strength decreases) nor significantly lack porosity (as cell infiltration, angiogenesis, and signaling become impossible). This trade-off is shown in Fig. 1, indicating that greater cell infiltration requires higher porosity, and porosity is generally inversely proportional to the mechanical strength of a given scaffold. This trade-off is generally one of the fundamental concepts of tissue engineering and must always be considered when fabricating biomaterials.
Fig. 1.
Three-dimensional scaffolds were very small when they first began to be used, but they have gained explosive popularity as researchers around the world seek to maximize their potential. In the meantime, various guidelines for the creation of 3D scaffolds have been established.Currently, 3D scaffolds are expected to exhibit at least three properties: functional space for cell and nutrient transport and adhesion, mechanical properties suitable for the intended application, and biocompatibility to prevent unwanted immune responses [17].
A major advantage of using 3D scaffolds for tissue engineering is that they can be functionalized for mechanical strength, degradation rate, and cell adhesion. There are various different techniques for functionalizing scaffolds, most of which depend on the material of interest. Surface modification to add molecules necessary for cell proliferation has been performed with various polymeric materials. For example, copolymerization using the same monomers has been used to functionalize polymeric scaffolds to promote cell adhesion to overcome the limitations of the parent polymer [18]. Another advantage from a functionality perspective lies in the manufacturing techniques, as 3D printing, electrospinning, microparticles, and hydrogels can all provide very different mechanical properties, degradation rates, and cell adhesion.
3D structures created by electrospinning in the form of nanofibers are one of the most widely used scaffold types and have been shown to be particularly useful for mimicking the extracellular environment due to their high surface-to-volume ratio, excellent mechanical properties, high porosity, and pore size distribution [19]. Furthermore, the diameter and orientation of the fibers in these scaffolds can be manipulated to obtain different cellular responses depending on the application [20].3D hydrogel-based scaffolds are also widely used due to their unique and stimulus-responsive properties and their unique ability to maintain their original structure well. Furthermore,by using hydrogel scaffolds, active agents such as growth factors can be released at a specific required rate due to their ability to encapsulate bioactive agents [21], [22].In addition to these approaches, scaffolds fabricated using microparticles of PLLA, PLGA, and their blends with natural polymers are popular due to their ability to reduce the degradation of various encapsulated biomolecules and also release them over a long period of time [23], [24], [25], [26].
Relatively new techniques for fabricating 3D scaffolds includedecellularization and 3D printing. Decellularization is a process of creating and functionalizing natural 3D scaffolds, where organs are obtained from animals (xenogeneic), all cells are removed using detergents, and then stem cells from the desired host candidate are re-implanted. Growth factors are sometimes added to decellularized scaffolds to promote differentiation [27]. Recently, whole organs have been regenerated and cellularized using donor stem cells. A well-known example is the recellularized heart, which has been shown to begin functioning again and successfully recover functionality [28]. In addition to the heart, several organs such as the lungs and bladder have been recellularized in vitro [29], [30]. However, decellularization has been shown to have many significant drawbacks, such as heterogeneous distribution of cells, difficulty in completely retaining the extracellular matrix, and immunogenicity if trace amounts of material remain before decellularization. Excessive or abnormal immune responses require high-risk, long-term immunosuppressive drug therapy [16].
Among these techniques,3D printing has the advantage of having nanoscale precision in scaffold dimensions compared to other conventional manufacturing methods. Countless bioactive materials have already been processed into scaffold materials using 3D printing.Even hydrogels are being 3D printed to create specific 3D scaffolds [31].This technology has been theoretically proven to surpass conventional techniques for 3D scaffold fabrication such as porogen leaching. The latter part of this article discusses various aspects of 3D scaffold fabrication using synthetic and natural biomaterials.
2. Synthetic and Natural Biomaterials for Tissue Engineering
2.1. Polycaprolactone (PCL)
Polycaprolactone is a famous polyester material first synthesized in the 1930s and is widely used in the fabrication of 3D scaffolds for tissue engineering applications. PCL is an elastic material consisting of non-polar methylene groups and semi-polar ester groups. PCL is used in various forms, including films, fibers, and microparticles. To improve bioactivity, it is used in various tissue regeneration applications in the form of composites with other polymers such as gelatin and chitosan [32], [33], [34], [35], [36], [37], [38]. In addition to tissue engineering applications, PCL is also widely used for various drug delivery applications and has received FDA approval for many different products [39], [40]. Many of these applications are explained in more detail in the applications section. Because PCL is relatively elastic compared to other polyesters, its mechanical use has become widespread [41]. PCL has the advantages of very high drug permeability, relatively slow degradation rate, and fewer acidic byproducts compared to other polyesters, making it useful for applications in drug delivery and, in recent years, tissue engineering [39], [42]. Polycaprolactone has been shown to have high solubility and blend compatibility with other biomaterials [43], [44], [45]. However, its relatively slow degradation rate (2-4 years) acts as a drawback, making it not an ideal scaffold for short-term drug or growth factor delivery applications [46]. Furthermore, PCL has poor cell adhesion by itself without some form of functionalization [47]. To overcome this drawback, numerous approaches have been used to improve its bioactivity, such as copolymerization, surface functionalization, and blend formation [48], [49]. For example, Chang et al. demonstrated in vitro that poly(epsilon-caprolactone)-grafted type II collagen-grafted-chondroitin sulfate (PCL-g-COL-g-CS) scaffolds fabricated by microparticle leaching and surface modification of PCL allow for chondrocyte proliferation when loaded with type II collagen and chondroitin sulfate [50]. Significant cell proliferation was observed during the 4-week culture period. Histological staining revealed that large amounts of collagen, a key cell viability marker, were being secreted. It was demonstrated that cells could maintain the exact same phenotype as chondrocytes within native cartilage tissue within porous PCL scaffolds, proving the viability of porous PCL scaffolds for tissue engineering applications and their ability to mimic the native tissue environment [51]. This finding indicates that the functionalization capability of PCL scaffolds is very high and is important in overcoming the lack of bioactivity of unmodified PCL.
2.2. Poly(L-lactic acid)
Poly(L-lactic acid) is a biodegradable synthetic polyester formed from the polymerization of l-lactide obtained from renewable resources such as starch, and has a wide range of applications including sutures, drug delivery vehicles, prosthetics, artificial blood vessels, bone screws, skin regeneration scaffolds, and fixation pins [52]. An example of a PLLA product already approved by the FDA is Sculptra™, which is an injectable currently used to treat facial atrophy [53]. Poly(L-lactic acid) decomposes into almost non-toxic byproducts and can be easily mixed with other materials, making its use widespread [54]. Although PLLA has a faster degradation rate due to bulk degradation compared to PCL, it is still considered relatively slow compared to other polymeric biomaterials used for tissue engineering scaffolds [55]. Furthermore, PLLA has high crystallinity of degradable fragments and can cause inflammation in the body, so it is sometimes blended with other polymers to form 3D scaffolds [56]. To solve this problem and enable higher bioactivity, Fukushima and Kimura demonstrated that PLLA can be fabricated as a combination of L-lactic acid and D, L-lactic acid. This is because D, L-lactic acid degrades more rapidly and lacks the inflammation associated with high crystallinity [57]. An example of a PLLA composite scaffold for inflammation suppression is shown in Fig. 2. This shows the surface morphology of a PLLA/Rg3 nanofiber matrix using SEM microscopy. The fibers shown represent pure PLLA (a), PLLA containing 2% Rg3 (b), PLLA containing 6% Rg3 (c), and PLLA containing 10% Rg3 (d). The purpose of Rg3, known as a scar formation inhibitory compound, is to counteract the inflammation associated with PLLA and allow for faster skin regeneration [58]. It can be seen that because the fibers are relatively uniform, cell infiltration into the pores of the scaffold is possible. In addition to electrospinning, various techniques such as 3D printing and solvent casting may be used to obtain blended PLLA scaffolds with enhanced non-toxic bioactivity.
Figure 2. The surface morphology of typical PLLA nanofiber scaffolds can be viewed using SEM microscopy [58]. The fibers represent pure PLLA (a), 2% Rg3-containing PLLA (b), 6% Rg3-containing PLLA (c), and 10% Rg3-containing PLLA (d). Rg3 is used to enhance the bioactivity of the scaffold, and since this compound plays an important role in reducing scarring, such materials are more useful for skin regeneration applications. It can also be seen that the fiber structure is relatively uniform, allowing for cell infiltration into the scaffold.
2.3. Poly(lactic-co-glycolic acid)
PLGA is a combination of the polyester polymers PLLA and PGA, and is one of the most commonly used biodegradable synthetic polymers for tissue engineering applications [59]. It is expected that the higher the ratio of PGA within the PLGA scaffold, the faster the PLGA will degrade. Its degradation byproducts, lactic acid and glycolic acid, are non-toxic [60]. PLGA is popular for various reasons, including its biodegradability, adaptability and customizability for various types of formulations, and surface modification for targeted drug delivery [59], [61]. An example of an FDA-approved PLGA scaffold is Osteofoam™ for bone regeneration applications, which has been shown to demonstrate a 3D morphology similar to human cancellous bone and allow for cell colonization [62], [63], [64]. Unfortunately, PLGA has one major drawback that limits its bioactivity. Its degradation byproducts are highly acidic, and if produced in large quantities, they become very difficult for the human body to metabolize quickly [65]. This can be a problem in the presence of acid-sensitive drugs, especially in drug delivery applications. Due to such a central drawback, some researchers have attempted to negate the effects of PLGA degradation byproducts. One conventionally well-known method is simply to change the ratio of PGA to PLLA, increasing the amount of PGA to slow down the degradation rate and release fewer acidic byproducts at once. However, it has recently been shown that in the presence of certain salts, the pH of PLGA byproducts can be raised, resulting in overall greater bioactivity, especially from the perspective of delivery applications [66]. PLGA is used in the production of nanoparticles, microparticles, and 3D scaffolds for drug delivery and tissue engineering applications, which will be explained in more detail in further sections of this paper [67], [68], [69], [70], [71].
2.4. Silk
Unlike the other polymers discussed so far, silk is a naturally derived macromolecular protein extruded from insects and worms. Also called silk fibroin, this biomaterial, which is biocompatible due to the protein components of silk, has been used not only for textile applications but also in the field of tissue engineering, particularly as a scaffold material. In particular, it is used in the field of tissue engineering due to its remarkable cell adhesion properties. Silk first requires the relatively time-consuming removal of sericin, a protein component with secondary toxicity, but the remaining fibroin component has been found to have relatively high tensile strength in addition to biocompatibility. For these reasons, it is widely used in the form of gels, sponges, and films for the regeneration of cartilage, bone, tendons, nerves, and ligaments [72], [73], [74], [75], [76], [77], [78]. To obtain properties such as elasticity, a constant degradation rate, and porosity, silk composite materials have been commonly used. For example, chitosan and silk fibroin nanofibers for wound dressing applications are very common and are produced by electrospinning [79]. In addition, a complex of silk and hydroxyapatite is used as a wound dressing by adding hydroxyapatite powder to silk fibroin and gelling it [80]. Freeze-drying, electrospinning, and 3D printing have been shown to be effective techniques for creating 3D scaffolds of silk biomaterials [73], [81], [82]. Fabrication techniques and applications using silk-based biomaterials will be explained in detail in later sections.
2.5. Collagen
There are approximately 29 known types of collagen, all of which have different properties and can be extracted from almost every biological species on Earth, including humans. Collagen exists in the extracellular matrix and bones in the form of fibers and gels, supporting tissues [83]. Due to its abundant nature and its role in organ growth and support, it has been used as a 3D scaffold for various tissue engineering applications: a wide range of applications from hard tissues such as bone to soft tissue regeneration such as cartilage, blood vessels, and nerves [84], [85], [86], [87], [88], [89]. Various types such as collagen I, II, III, V, and XI have been tested for tissue engineering applications. Of these, type I collagen is called the "gold standard" by many scholars due to its low associated immunoreactivity [83]. Furthermore, because the differences in collagen properties between different host species are very small, unintended variations can occur within a specific manufactured scaffold.
Furthermore, collagen scaffolds can unfortunately be relatively difficult to synthesize without making significant changes to the integrity of the intended structure, as is the case with proteins [83]. Collagen scaffolds typically exhibit a relatively rough surface morphology, which contributes to their fibrous nature and the structural porosity of the sample. Figure 3A shows an overview of a typical collagen scaffold and its dimensions, and Figure 3B shows the surface and cross-sectional views. Using SEM microscopy techniques, we are looking at the porosity with an average pore size of approximately 80 μm. This figure also highlights the rough surface morphology often associated with collagen. The scaffolds were fabricated by freeze-drying techniques to create the indicated porous structure.
Figure 3. A typical collagen scaffold is shown above. (A) is a scaffold with a diameter of 8 mm and a thickness of 2 mm, and (B) is a surface view (left) and cross-sectional view (right) using SEM techniques [91]. The scaffold has an average pore size of approximately 80 μm and was fabricated using a freeze-dryer to create the indicated porous structure. The rough surface morphology associated with many collagen type II scaffolds can be seen.
Regarding manufacturing techniques, aligned fibrous collagen scaffolds have been created by standard electrospinning techniques, and it has been shown that rabbit conjunctival fibroblasts proliferate more rapidly on aligned collagen fibrous scaffolds than on random collagen fibrous scaffolds, highlighting the role of the manufacturing technique used [92], [93]. Porous collagen-based scaffolds have also been created by solvent casting-particulate leaching, phase separation, gas foaming, emulsion freeze-drying, and fiber mesh. Many of these methods fail to ensure cell adhesion due to changes in surface topography. Therefore, solid freeform fabrication is used and has been shown to be more effective in terms of cell adhesion [94]. Since cell adhesion to collagen scaffolds is highly dependent on the surface morphology of the biomaterial, it is important to carefully consider the fabrication method used and its optimization.
2.6. Hyaluronic Acid (HA)
Hyaluronic acid is a type of non-adhesive glycosaminoglycan, a natural biodegradable material found abundantly in connective, epithelial, and nervous tissues [95]. HA scaffolds are used for the regeneration of both hard and soft tissues, which will be discussed in more detail in future sections. However, one of the frequent uses of HA for tissue engineering applications is in the form of hydrogel materials. This is due to its swelling capacity and its ability to encapsulate cells and other materials for delivery applications. Cell viability in HA hydrogel constructs is shown in Fig. 4. This figure shows cell viability on HA scaffolds as indicated by specific cell viability markers important for proliferation, namely the secretion of collagen and aggrecan. Collagen and aggrecan staining increased, and the increase was dependent on the concentration of the cross-linking agent used in the hydrogel material. Interestingly, at the edges of the construct, both collagen and aggrecan were present in higher concentrations than in the center. Even if mechanical properties require adjustment through chemical modification to resemble native tissue, HA has similar physical and biological functions [96], [97], [98].
Figure 4. Cell staining for collagen type II and aggrecan markers shows cell viability on HA scaffolds photocrosslinked in vitro for cartilage tissue engineering applications [103]. Collagen type II and aggrecan are important byproducts for indicating cell viability in many applications. As the amount of cross-linking throughout the scaffold increases, the secretion of collagen and aggrecan markers can be observed.
It is also observed that collagen type II and aggrecan markers are concentrated at the edges rather than in the center of the construct. In addition to its role as a hydrogel, there are less common methods for fabricating HA scaffolds. HA scaffolds can be made by electrospinning to form fibrous scaffolds [99] or blended with other biomaterials to create porous scaffolds by salt particle leaching [100]. More recently, "wet spinning," a new spray-assisted layer-by-layer assembly technique, has been used to deposit various polyelectrolyte membranes onto porous HA scaffolds, making cell adhesion and proliferation of human keratinocytes in vitro more possible [101]. Finally, a promising technique for the fabrication of HA scaffolds appears to be 3D printing, which allows for microscale precision of scaffold parameters and may lead to higher cell performance in animal models [102]. The popularity of HA seems to be largely due to a wide range of diverse manufacturing techniques that pave the way for using this biomaterial in almost all tissue regeneration applications.
2.7. Chitosan
Chitosan is a linear biodegradable polysaccharide obtained from the partial deacetylation of chitin by chemical hydrolysis. It has been confirmed that 3D scaffolds of chitosan have functional and structural properties similar to glycosaminoglycans, which are lubricants present in the human body. Unlike some of the biomaterials already mentioned, thispromotes natural cell adhesion without further functionalization, as well asproving its bioactivity [104], [105], [106]. Therefore,3D scaffolds of chitosan alone or in combination with other natural polymers are used in many tissue engineering applications in the form of gels, sponges, or fibers [7], [107]. Chitosan as a biomaterial has limited solubility at physiological pH, which isadvantageous for long-term use [108], [109]. Therefore, chitosan is considereda viable biomaterial for tissue engineering applications in the form of 3D scaffolds. Furthermore, due to its multifunctional structure and cross-linking ability, chitosan is often blended with other biomaterials to modify the properties of the scaffold. Figure 5 shows SEM images of chitosan-collagen composite microparticles.
Note that these scaffolds exist in a uniform porous structure and have adequate mechanical stability. These structures also show that blending chitosan with other well-known biomaterials yields scaffolds with desirable bioactive properties. For example, HA and chitosan have been shown to have exceptional structural performance, with positive cartilage staining for collagen type II and GAG [110]. Also, chitosan/PCL scaffolds used for tissue engineering applications show that bovine articular chondrocytes adhere and proliferate on such 3D scaffolds in vitro after 21 days [111].In applications for nerve regeneration, collagen-chitosan scaffolds with RGD were demonstrated to stimulate linear axonal growth in rats with 15 mm long nerve defects after 4 months [112]. These applications will be explained in more detail in future sections.
Figure 5. The microstructure of a typical chitosan-collagen composite is shown above using SEM techniques [113]. Part (A) is a magnified image of the structure, and (B) shows the sphericity of the microparticles. (C) shows the honeycomb structure, and (D), (E), and (F) show the unique orientation of the microchannels of the entire structure formed by the microparticles bonding to each other.
3. Types of Bioactive 3D Structures
3.1. 3D Printed Structures
Recent advances in 3D bioprinting have revolutionized the field of tissue engineering. Using various complex algorithms, 2D images on a computer screen can be processed into 3D equivalents in real time. When biocompatible 3D scaffolds must be printed, a series of different steps are required [114]. Precision ranges from millimeters to the nanoscale, and the higher the precision, the higher the cost. Figure 6 shows the precise similarity between the original scan and the 3D printed model. In (A), only the MRI image is used, and in (B), an image digitally rendered with dedicated software is constructed. The final result is (C). The leaves are finely connected, indicating the high precision of 3D printing.
Figure 6. Shows a 3D printed brain model [115]. (A) is a sagittal view of the brain from an MRI scan, (B) is a rendered digital image, (C) is the 3D printed result, and (D) shows the model being electrically stimulated. This model can be electrically stimulated in vitro for cell studies. However, it seems it will still take a considerable amount of time before it can fully reproduce the functions of the human brain.
Similarly, several different types of tissues, including multilayer skin, bone, artificial blood vessels, tracheal splints, heart tissue, and cartilage structures, have already been manufactured and transplanted by 3D printing [116]. 3D bioprinting allows for the creation of more complex geometric parameters, which is a major advantage compared to conventional methods such as porogen leaching. Furthermore, it can improve productivity and be more cost-effective with a more rational approach [117]. 3D printed scaffolds have been manufactured and applied for engineering applications of both hard and soft tissues, such as bone, cartilage, nerves, and muscles [118], [119], [120], [121]. For example, human chondrocytes bioprinted in nanocellulose alginate bioink showed 73% viability after 1 day and 86% viability after 7 days of culturing cells on the scaffold in vitro, demonstrating the viability of 3D printed scaffolds for tissue engineering applications [119]. Similarly, collagen scaffolds can also be 3D printed, and it was demonstrated that hMSCs show attachment and proliferation in vitro over 4 weeks on 3D bioprinted collagen scaffolds with a predefined capillary network [122]. It is worth noting that while such scaffolds with cell viability similar to those obtained by conventional methods have been created by 3D printing technology, there are also drawbacks associated with 3D printing technology. For example, the biomaterial must be in a liquid state to print droplets, and it is difficult to achieve biologically appropriate cell densities [116]. As mentioned earlier, one of the major advantages of 3D printing in tissue engineering applications is that the spatial parameters of the network can be set extremely accurately.
3.2. Nanofibers
Nanofibers have become popular primarily because, as shown in Figure 7, they can mimic the inherent properties of the extracellular matrix at such a small, precise scale [123], [124]. Figure 7 shows a uniform nanofiber matrix. Part (a) of the figure shows a very fine diameter from a close-up view, demonstrating the uniformity of individual fibers. (b) and (c) are enlarged views, showing that the uniformity of the fibers is still visualized. Such fibers, typically produced by electrospinning, are frequently used as strong reinforcements in nanocomposites [125], [126]. In addition to their ability to mimic the native extracellular matrix, nanofiber scaffolds exhibit a higher surface-to-volume ratio, leading to greater cell attachment than larger fibers [127]. Furthermore, nanofiber-reinforced composites have been shown to have higher mechanical strength than conventional unfilled or carbon/glass fiber-filled composites, adding a significant advantage to scaffolds requiring high mechanical strength [128].
Figure 7. SEM showing nanofiber morphology [129]. A) shows the ultrafine diameter of the fibers, and (B) and (C) show the morphology before and after immersion in an aqueous solution, respectively. Polystyrene (PS) was used as the material. Through inspection, it can be visually confirmed that the fiber diameter, which is important for cell compatibility, is uniform.
Various materials, such as PCL and chitosan, have been used to produce nanofibers [38], [130], [131], [132], [133]. Such fibers are either aligned or randomly oriented. Aligned nanofibers have very specific applications and can be used to specify the direction of tissue growth. It has been reported that when neurites obtained from dorsal root ganglion explants are seeded onto uniformly aligned electrospun PLA nanofibers, they grow outward from the ganglion in the direction of the fibers in vitro [134]. It has also been shown that a polyvinyl alcohol random fibrous scaffold supplemented with chondroitin sulfate promoted chondrogenic differentiation of mesenchymal stem cells in vitro and promoted the growth of osteogenic defects in vivo in rats, indicating that nanofibers are useful for cartilage tissue engineering applications [135]. However, one major advantage of random fibers is that when subjected to mechanical testing, their mechanical properties can be altered to improve stiffness and resistance in all directions [136], [137], [138], [139]. This is a significant advantage compared to aligned nanofibers, which exhibit strong mechanical properties only in the direction of the fibers. Therefore, fine-tuning of mechanical properties is paramount when creating nanofiber scaffolds, creating a general rule of thumb: aligned fibers exhibit greater strength in one direction and are therefore more useful for tendon and ligament regeneration applications, while random nanofibers exhibit similar mechanical properties in all directions and are therefore more useful for skin and cartilage regeneration applications.
3.3. Microparticles
Microparticles were initially developed as carriers for anticancer drugs, but they have now also entered the field of tissue engineering. The unique aspect of microparticles is that they can deliver growth factors and soluble drugs in a slow, controlled manner and can be engineered to allow for some degree of site-specific targeting [140]. These properties have been shown to allow for stable and cumulative release of growth factors and drugs when embedded in 3D scaffolds [141]. In addition to serving as a delivery vehicle by being embedded in 3D scaffolds, microparticles have many applications, and are particularly popular for use as injectable scaffolds because they overlap with drug delivery applications. A 3D construct seeded with genipin-crosslinked chitosan microparticles and goat bone marrow stromal cells (GBMC) demonstrated feasibility as an injectable scaffold after evaluating cell viability at 7 and 14 days in vitro, paving the way for injectable applications [25]. More recently, injectable microspheres have been demonstrated in vivo in rats. Lovastatin microparticles combined with polyurethane (PUR) scaffolds resulted in the sustained release of lovastatin over 14 days, stimulating the expression of BMP-2 growth factor in osteoblasts at the defect site [142]. Porous PLLA microparticle scaffolds containing PVA and treated with serum showed superior cell adhesion properties compared to other types of 3D microparticle scaffolds in vitro [24]. Such findings indicate that microparticle-based scaffolds provide a unique overlap between tissue engineering and drug delivery applications. Atomization, spray drying, and sintering have traditionally been used for the fabrication and functionalization of microparticles [143]. Oil-in-water dispersion techniques have also been relatively commonly used for the production of microparticles [144]. Today, a very common functionalization technique is perhaps laser sintering, which can impart further bioactivity to the microparticle surface, paving the way for the attachment of various molecules as well as cell adhesion properties [145]. In vitro tests have demonstrated that sintered 3D chitosan/PLGA microspheres are useful for bone tissue engineering applications because MC3T3-E1 osteoblast-like cells adhere to and proliferate on the material surface [146].
3.4. Hydrogels
Hydrogels are water-absorbing polymer materials made using synthetic or natural polymers, which are mostly hydrophilic and very flexible. Hydrogels have many advantages that make them popular as 3D scaffolds for tissue engineering applications, such as their structural similarity to native ECM and the potential to deliver drugs and growth factors in a non-invasive manner [147]. However, due to their high water content (sometimes up to 90%), hydrogels are usually mechanically very weak and inefficient for loading cells, so further improvements are needed to increase their bioactivity [148]. Many different synthetic and natural materials are used in the form of hydrogels, with HA and PEG being two examples already discussed in the previous section [98]. Other examples of natural polymers include collagen, gelatin, fibrin, alginate, and agarose [22], [149], [150]. Synthetic polymers that can be used in hydrogel form include poly(acrylic acid) (PAA), poly(ethylene oxide) (PEO), PVA,
polyphosphazenes, and various polypeptides [151]. Hydrogels have been fabricated using physical and chemical methods. Warming polymer solutions to form gels, crosslinking in aqueous solutions, lowering the pH of aqueous solutions, mixing solutions to form coacervate gels, gelling polyelectrolyte solutions using oppositely charged polyvalent ions, and crosslinking specific polymers in the solid state with different types of radiation or chemical crosslinkers are the most commonly used techniques for hydrogel fabrication [148].One of the relatively recent applications of these hydrogels is cell encapsulation, which provides a hydrated environment for cells to proliferate. The degradation of such encapsulated hydrogels has been shown to depend on the segments of the hydrogel itself, and natural biopolymers that can be degraded by enzymes can be used [152]. Hydrogels can also befunctionalized with peptides such as RGDS for further bioactivity, and have been shown toenable cell attachment capabilities far exceeding those of typical non-functionalized hydrogel scaffolds[153]. Similarly, hydrogels have also beenused to encapsulate growth factors for drug delivery applications, which can be released slowly over time [150]. Thus, hydrogels are similar to the microspheres discussed in the previous section in that they are commonly used as encapsulation materials, and therefore have a significant overlap with drug delivery. However, they are more commonly used to load cells rather than being embedded within a scaffold to release growth factors or other secondary substances.
4. Applications
4.1. Nerve Regeneration
The nervous system is divided into the central nervous system (CNS) and the peripheral nervous system (PNS). It is important to distinguish between the two in that the CNS consists of the brain and spinal cord, while the PNS consists of ganglia and nerve tissue other than the CNS. Nerve regeneration applications in tissue engineering have traditionally dealt with both systems [154], [155], [156]. However, there are various unresolved challenges, such as the fact that mouse nerve axons are smaller and shorter than human ones, making in vivo models not work well, and that atrophy of target tissues makes functional recovery difficult [156]. Therefore, to bridge the gap between theory and demonstration, further in vivo testing in large animal models closer to humans is needed. Schwann cells are being studied at the forefront of PNS regeneration applications because they contribute to structural support and myelination of axons. Therefore, 3D scaffolds used for nerve regeneration applications must have bioactivity that promotes myelination, axonal elongation, and structural support via Schwann cells. Among the 3D structures mentioned earlier, nanofibers appear to play an important role in nerve regeneration, especially when dealing with nerve guidance conduits (NGCs) designed to induce axonal elongation of neurites. NGCs have been reported to have multiple applications, including the ability to exhibit multifunctional properties aimed at guiding axon growth from proximal nerve ends, secreting growth factors that aid tissue regeneration, and reducing inhibitory scar tissue at the injury site [157], [158]. Figure 8 shows a digital image of a typical NGC. Part (A) shows a digitally constructed image of an NGC, part (B) shows two severed nerve ends, and part (C) shows the typical placement of an NGC joining the severed nerve from both ends. Visually, it can be seen that the NGC is intended to bridge the gap between the two severed nerve ends and to extend the nerve itself. This has greatly increased the potential for recovery from nerve injury in recent years.
Figure 8. A typical PGA/collagen NGC is shown in the figure above [159]. A) is a digital image of the conduit, and (B) and (C) show how it fits into the nerve gap. The ruptured nerve has two cut ends, and the tube fits directly between them and can extend further toward the gap itself.
In a laboratory setting, NGCs are often the subject of novel customization. Xie et al. reported that a novel combination of both bilayer aligned PCL nanofibers and randomly oriented electrospun PCL nanofibers was useful for guiding neurite elongation in vitro when pre-seeded with Schwann cells, and enabled moderate functional recovery in an in vivo 14 mm rat sciatic nerve injury model [160]. This novel combination of a bilayer NGC, consisting of randomly oriented nanofibers in the outer layer and aligned nanofibers in the inner layer, was also observed to be much more tear-resistant during surgical procedures compared to NGCs consisting only of conventional aligned nanofibers. Such scaffolds, which are much more robust and tear-resistant in vivo, demonstrate their real-world capabilities. Interestingly, while nanofibers remain very popular as NGCs, it was recently discovered that protein films consisting of blended silk fibroin and coated human tropoelastin protein exhibit significant amounts of neurite elongation and Schwann cell area growth in vitro [161]. This new biodegradable scaffold gained both a robust biomaterial component from silk and a 2.4-fold improvement in neurite elongation capability from tropoelastin protein compared to standard poly-d-lysine film coatings. NGCs are often functionalized with proteins such as collagen and laminin to further stimulate nerve axon elongation and similarly improve nerve functional recovery [162]. Recently,hydrogels have also been adopted as scaffolds for nervous system regeneration, as findings suggest that scaffolds of this nature may play a role in aiding Schwann cell-based axonal recovery. Tseng et al. recently demonstrated thata chitosan-based hydrogel (1.5 kPa stiffness) containing proliferating and differentiating neural progenitors can be injected in vivo for central nervous system regeneration using a zebrafish injury model [163].The novelty of this study stems from the fact that it is one of the few examples whereself-healing hydrogels, which are expected to become widely used in the future, have been shown to be effective in vivo for nerve regeneration applications. Althoughsilk fibroin-based hydrogels show little cytotoxicity and enable significant nerve regeneration when used with Schwann cell cultures in vitro, they are unfortunately still rarely used for nerve regeneration applications [164]. Much of the success of hydrogel scaffolds is related to stiffness. It has been reported that neural stem cells tend to differentiate into astrocytes and neurons when cultured on soft hydrogels, and into oligodendrocytes on stiff hydrogels [165]. For these reasons, stiffness is usually emphasized, as in the study by Tseng et al. mentioned earlier.
4.2. Bone Regeneration
Synthetic polymer scaffolds for tissue engineering applications have begun to eliminate the need for bone grafts, which have traditionally been used to treat osteogenic defects [166]. Recently, there has also been a move toward nanostructured materials for bone regeneration because they can react at the cellular level [167]. Hydroxyapatite, β-tricalcium phosphate (β-TCP), and bioactive glass are well-known as scaffold materials for bone regeneration due to their structural similarity [168]. The advantages of inorganic materials like these appear to be their high compressive strength and potential for osteoconductivity [169].
These materials are often blended with other biodegradable polymers to obtain higher bioactivity. Recently, a new type of mixed scaffold was announced, and Zhang et al. proved that a porous nano-hydroxyapatite/PCL spiral scaffold (1:4 ratio), fabricated using a modified salt-leaching technique and seeded with human fetal osteoblasts (hFOBs) for 14 days, significantly increases the amount of mineralized extracellular matrix material. Bone mineralization markers, including bone sialoprotein (BSP), osteonectin (ON), osteocalcin (OC), and type I collagen, were measured using reverse transcription polymerase chain reaction analysis [170].
In tissue engineering for bone regeneration, the grain size of the scaffold is one of the most important points. Grain size refers to the size of the individual pieces of material, or "grains." It has been shown that smaller grain sizes are advantageous for cell attachment and proliferation, as well as for the differentiation of most osteogenic lineages [171]. The cause of this phenomenon is controversial and unknown, as studies have not yet elucidated the mechanism. Nevertheless, due to the observed cell attachment profiles, manufacturing processes are often executed to produce smaller grain sizes. It has also been demonstrated using hydroxyapatite scaffolds that grain size depends on sintering temperature. A temperature of 1325°C was found to be the ideal sintering temperature to achieve an ideal small grain size without compromising porosity [172]. Such findings are an important breakthrough for researchers who desire standardized procedures for sintering these biomaterials.
Apart from grain size, another important factor essential for bone regeneration is mechanical strength that mimics the native tissue. For this reason, polymer materials with high mechanical strength are often used as scaffolds for bone tissue engineering applications, including silk proteins, PLLA, chitosan nanofibers, and bioactive glass materials [139], [173], [174], [175]. Metal scaffolds such as titanium are widely popular due to their high compressive strength, porosity, and fatigue resistance [176]. Recently, Guneta et al. attempted a new breakthrough by demonstrating that 3D-printed titanium scaffolds are effective for bone tissue engineering applications [177]. By varying the sintering temperature of these titanium scaffolds from 1250°C to 1370°C, pore sizes in the range of 17 μm to 24 μm were achieved, theoretically proving that mechanical properties can be optimized from a combination of both 3D printing specifications and variable sintering temperatures.
This is consistent with other findings mentioned earlier that pore size changes depending on sintering temperature. Unfortunately, the mechanical strength and angiogenic potential of scaffolds currently applicable to bone tissue engineering are still insufficient, despite the wide variety of scaffolds available [178]. However, the research on 3D-printed scaffolds by Guneta et al. perhaps provides a viable alternative for the future.
Although various solutions have been presented in recent years, angiogenesis remains one of the key challenges in bone tissue engineering. It is known that insufficient angiogenesis leads to a strong deficiency of nutrients essential for cell survival within the scaffold, leading to unexpected and dangerous irregularities in differentiation [179]. One approach to stimulate angiogenesis is to introduce certain growth factors into the scaffold. Growth factors such as vascular endothelial growth factor (VEGF) have been demonstrated by Wernike et al. to significantly increase vascular density and bring osteogenic cells to the defect site in a mouse model in vivo [180]. Angiogenesis over 28 days was reported using intravital microscopy. Interestingly, it was shown that the concentration of VEGF was a determinant of angiogenesis, but only at limited concentrations, as locally high concentrations of growth factor unfortunately created malformed blood vessels. Growth factors are also sometimes combined within scaffolds to perform multiple functions. The combination of growth factors BMP-2 (loaded in PLGA microspheres for bone regeneration) and VEGF (loaded in gelatin hydrogel for angiogenesis) embedded in a polypropylene (PP) scaffold has been shown by Kempen et al. to promote osteogenesis and angiogenesis in an in vivo rat bone defect model [181]. Interestingly, despite the traditional role of VEGF primarily involving angiogenesis, the combination of VEGF and BMP-2 was found to increase osteogenesis over 56 days more than BMP-2 alone. This demonstrates that in most cases, a combination of growth factors is desired. Figure 9 further demonstrates this phenomenon shown by Kempen et al. and discussed earlier, as it can be seen that the incorporation of VEGF and BMP-2 simultaneously induces both angiogenesis and further bone growth. Compared to parts (a) and (b) of the figure without growth factors, the growth factors in parts (c) and (d) result in greater cell proliferation and differentiation related to angiogenesis and osteogenesis.
Figure 9. Silk fibroin scaffolds were stained with GFP at 8 weeks post-surgery in vivo [170]. a) is pure silk, (b) is cell proliferation on the scaffold, (c) shows silk with VEGF and BMP-2 growth factors for angiogenesis, and (d) shows cell proliferation again. It can be seen that new blood vessels are formed by VEGF stimulation.
To satisfy both mechanical strength and an environment suitable for angiogenesis, PLGA scaffolds are considered strong candidates for bone tissue engineering, especially when blended with other bioactive materials to enhance cell adhesion [64], [182]. Sheik et al. recently announced a novel PLGA/silk hybrid scaffold for bone tissue engineering applications that combines the degradation rate of PLGA with hydrophilic silk polymers and is further combined with hydroxyapatite nanoparticles to improve biocompatibility, and its effectiveness was evaluated both in vitro and in vivo [183]. Variable pressure field emission scanning electron microscopy (VP-FE-SEM) was used to demonstrate the porosity of the scaffold, and contact angle measurements showed that the silk component imparts additional hydrophilicity to the scaffold. These findings represent an important step in showing how PLGA scaffolds can be modified or blended for greater cell infiltration and overall bioactivity. Osteoblasts were cultured on the scaffolds for 14 days in vitro, and subsequent MTT assays showed cell attachment and proliferation during the culture period. Furthermore, when implanted into rat calvarial defects for 4 weeks, bone formation was confirmed by hematoxylin and eosin (H&E) staining. Thus, in the field of bone regeneration, concepts such as grain size, mechanical strength, and angiogenesis remain dilemmas, and research continues to optimize these parameters.
4.3. Muscle Regeneration
Muscle tissue regeneration has posed a very difficult challenge, perhaps even more so than nerve or bone tissue regeneration. These challenges are primarily due to the fact that the scaffold must have structural integrity while simultaneously being able to induce both strong contraction and force regeneration [184]. It should be noted that there are three types of muscle: cardiac muscle, smooth muscle, and skeletal muscle. Cardiac muscle tissue is primarily found in the walls of the heart, smooth muscle tissue is primarily found in the walls of several other organs, and skeletal muscle fibers are specifically attached to the skeleton. Of these, cardiac muscle regeneration is particularly important because cardiac muscle tissue generally has extremely limited natural regenerative capacity in most mammals [184]. At the forefront of cardiac muscle regeneration are cardiomyocytes, which can differentiate to form the most basic structure of cardiac muscle tissue. Cardiomyocytes are often seeded onto PGA, gelatin, alginate, or collagen scaffolds, but it has been proposed that stacking cell sheets in a 3D structure without an artificial scaffold may provide higher bioactivity [185].
When peptides such as RGD are attached to a scaffold, cell functionality is improved. Previously, it was shown that neonatal rat heart cells seeded onto RGD-immobilized macroporous alginate scaffolds promoted cell adhesion more than unmodified alginate scaffolds and promoted in vitro cardiac tissue regeneration in cardiac patches [186]. Expression of proteins essential for cell activity, such as α-actinin, N-cadherin, and connexin-43, was confirmed by Western blotting. Furthermore, cell apoptosis was significantly reduced in the RGD-immobilized macroporous alginate scaffolds compared to the control group. Also, regarding the overall functionality of the scaffold, particularly cell adhesion, cell adhesion molecules such as vitronectin and fibronectin play a similar role to RGD [187].
Smooth muscle cells are unique in that they can transition reversibly between a quiescent contractile phenotype and a synthetic phenotype, posing challenges for tissue engineering applications, especially when reversing to a contractile phenotype [188]. Nanofibers such as PCL and collagen have been shown to be particularly useful as scaffolds for smooth muscle tissue engineering because growth is guided by the orientation of the nanofibers, allowing cells to maintain their typical phenotypic shape [189]. However, crosslinking can also be utilized for favorable mechanical properties and growth in multiple directions. Recently, crosslinked multilayer electrospun gelatin nanofibers (±45° orientation) have been demonstrated to be useful biomaterials for the in vitro culture of human umbilical vein smooth muscle cells (HUVSMCs), with cell viability reaching 80%–92% after 9 days and cell proliferation occurring along various fibers [190].
Skeletal muscle has a large, already incorporated capacity to regenerate without external intervention because such tissue is constantly being destroyed, repaired, and remodeled. However, in the case of very severe injury, this regenerative capacity can be limited or completely lost. During the repair and regeneration phases, satellite cells are known to play a very important role in migrating to the defect site, proliferating, and differentiating to restore functional properties [191]. Collins et al. demonstrated the importance of satellite cells in skeletal muscle tissue engineering applications, showing that seeding as few as 7 of these cells into the muscle tissue of irradiated mice could produce over 100 muscle fibers in vivo [192]. Thousands of muscle nuclei were reported. Hydrogels, fibrous meshes, and patterned substrates are commonly used as scaffolds for seeding stem cells to regenerate skeletal muscle tissue [193].
Furthermore, electrical stimulation has been shown to be a novel catalyst with the potential to expand many myogenic progenitor cells on 3D scaffolds in vitro, and may become an important method in the future for the expansion of other cell types such as satellite cells [158]. Although the reason is unknown, as in the paper by Serena et al., routine electrophysiological stimulation has been shown to improve the differentiation potential of muscle progenitor cells (MPCs) both in vivo and in vitro [194]. MPCs were seeded onto 3D collagen scaffolds and subjected to routine electrical stimulation. Electrical stimulation did not affect cell viability, but it was noted that NO(x), a satellite cell activator, increased with a 65% greater release rate. Other myogenic markers such as desmin also increased compared to the control group. Serena et al. transplanted the new scaffolds with electrical stimulation in vivo into the tibialis anterior muscle of mice and confirmed the formation of new muscle fibers after 10 days.
From this section, we can conclude that muscle regeneration is unique in that it targets a more diverse range of tissues, including cardiac, smooth, and skeletal muscle. Each tissue type has its own key players, and it is highly likely that there is no single scaffold useful for the regeneration of all tissue types.
4.4. Tendon and Ligament Regeneration
Compared to the skeletal muscle tissue mentioned earlier, tendons do not regenerate well naturally after injury, and even minor injuries can make the entire healing process difficult. For this reason, 3D scaffolds play a crucial role in tendon regeneration. Conventional methods such as grafts cannot restore the mechanical and structural properties of the original tendon, nor can they stimulate cell proliferation. Many solutions have been proposed for this dilemma over the years. Since the tendon ECM is composed mostly of type I collagen and exhibits an intricately intertwined structure, it is relatively difficult to reproduce a 3D biomimetic environment.
Achilles tendon regeneration is a major hurdle due to the constant mechanical loading [195]. For this regenerative application, collagen has become popular as a scaffold material. Juncosa-Melvin et al. reported that a 2 cm long rabbit AT defect was reduced to 85% of the original maximum stress and elastic modulus after 12 weeks by transplantation of a type I collagen gel loaded with MSCs (cell-to-collagen ratio 0.08 M/mg) in vivo [196]. Juncosa-Melvin et al. concluded that it is necessary to test at high stiffness even at lower cell densities to avoid excessive contraction and weak mechanical properties that cause tearing during culture.
Juncosa-Melvin et al. concluded that it is necessary to test at high stiffness even at lower cell densities to avoid excessive contraction and weakening of mechanical properties that cause tearing during culture. For all tendon regeneration applications, polymeric materials exhibiting very strong mechanical properties, such as silk and PLGA, have also been used as scaffolds. PLGA/silk fibrous scaffolds have been reported to be used as bFGF-releasing devices to stimulate the differentiation and adhesion of mesenchymal progenitor cells (MPCs) in vitro [197]. PLGA fibers were used for bFGF encapsulation and release, and microfibrous silk was used as a reinforcing material. The results showed that gene expression of common ligament and tendon ECM proteins increased, and collagen production increased. Other past studies, such as those by Ouyang et al., reported that knitted PLGA loaded with bone marrow stromal cells (bMSCs) stimulated type I collagen production in 10 mm long rabbit AT defects in vivo, helping to restore the native tendon environment [198]. Compared to the control group where knitted PLGA alone was used or the defect was left untreated, no lymphocyte infiltration was reported in the PLGA scaffolds carrying bMSCs. Immunohistochemical analysis showed strong expression of collagen type I and collagen type III, indicating that the native ECM environment was reconstructed.
Since the anterior cruciate ligament plays an important role in knee stabilization, ligaments have attracted attention in the field of tissue engineering. Because the ACL has low healing capacity and requires excellent mechanical stability, polymeric scaffolds such as PLLA and silk fibers have been used to deliver growth factors and mimic the mechanical properties of the native tissue environment [199]. Recently, a new scaffold for ACL regeneration has been reported, consisting of a mechanically strong extruded PLLA nanofiber and a flexible shell of electrospun PCL nanofibers, which can incorporate bFGF and platelet-derived growth factor (PDGF) in a controlled-release manner, and can further proliferate hMSCs in vitro [200].
Both PLLA and PCL fibers were fabricated by electrospinning. Gene expression confirmed that key ligament markers necessary for cell survival, such as type I collagen, type III collagen, tenascin C, and toughness, increased over 21 days. AT and ACL are considered the two major hurdles to overcome in tendon and ligament regeneration due to their complexity and high frequency of injury. Therefore, it is believed that ultimately, it will be possible to mimic the complexity of native tissue through a combination of bioactive substances.
5.5. Conclusions and Perspectives
While there are still many challenges, 3D structured biomaterial-based scaffolds hold a promising future in tissue engineering applications. Since the emergence of tissue engineering in the 1990s, what was once considered science fiction is now becoming a reality. Scaffolds have been fabricated using a variety of synthetic and natural biomimetic materials. The main challenges of current technology include inducing sufficient vascularization, especially in bone tissue engineering, finding ways to efficiently manufacture such scaffolds in non-laboratory environments, fine-tuning degradation rates for specific applications, and creating scaffolds with biomimetic mechanical and structural properties. Of particular importance in this review is a manufacturing technique called 3D printing, which perhaps more than any other method, gives hope for the future mass production capability of 3D scaffolds. Growth factors are also at the forefront of scaffold materials, especially regarding the hurdle of vascularization, as they can bring about cell differentiation and proliferation at a faster rate. Thus, it is certain that there are various methods. Ultimately, by combining different manufacturing methods, different growth factors, and different biomaterials, it is highly likely that they will become excellent scaffolds suitable for very specific applications. While NGCs for nerve regeneration are certainly among the most complex scaffold structures mentioned in this review, the trend toward further complexity is expected to continue as more and more research combines different materials and methods. With new discoveries constantly being made, the next generation of tissue regeneration scaffolds will likely utilize more functionalization techniques to enhance bioactivity, in contrast to the previous generation of tissue engineering that relied on the properties of unmodified biomaterials themselves.
Acknowledgments
We also acknowledge funding from the National Institute of Health-5R03NS058595, Connecticut Regenerative Medicine Research Fund-15-RMBUCHC-08, National Science Foundation (Award Numbers IIP-1311907, IIP-1355327; EFRI-1332329), and Department of Defense (OR120140).
References
[1]
R. Langer, J.P. Vacanti
Tissue Engineering
Science, 260 (1993), pp. 920-926
View PDF CrossRef Google Scholar
[2]
P. Lee, K. Tran, W. Chang, Y.-L. Fang, G. Zhou, R. Junka, et al.
Bioactive polymeric scaffolds for osteochondral tissue engineering: In vitro evaluation of the effect of culture medium on bone marrow stromal cells
Polym. Adv. Technol., 26 (2015), pp. 1476-1485.
View PDF CrossRef View record in Scopus Google Scholar
[3]
P. Lee, K. Tran, G. Zhou, A. Bedi, N.B. Shelke, X. Yu, et al.
Induced differentiation of bone marrow stromal cells on scaffolds co-cultured with cartilage and bone
Soft Matter, 11 (2015), pp. 7648-7655
View PDF CrossRef View record in Scopus Google Scholar
[4]
P. Lee, K. Tran, W. Chang, N.B. Shelke, S.G. Kumbar, X. Yu
Effect of the presence and arrangement of chondroitin sulfate and hyaluronic acid in nanofibers on bone marrow stromal cells: Cartilage regeneration
J. Biomed. Nanotechnol., 10 (2014), pp. 1469-1479
View PDF CrossRef View record in Scopus Google Scholar
[5]
E. Piskin
Biodegradable polymers as biomaterials
J. Biomater. Sci. Polym. Ed., 6 (1995), pp. 775-795
View record in Scopus Google Scholar
[6]
N. Shelke, R. Nagalare, S. Kumbar
Chapter 7 - Polyurethane
S.G. Kumbar, C. Laurencin, M. Meng Deng (Eds.), Natural and Synthetic Biomedical Polymers (2014), pp. 123-144.
Article Download PDF View record in Scopus Google Scholar
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[Abstract] Injectable self-healing hydrogels for repairing the central nervous system (chitosan-based hydrogels)
Ting-Chen Tseng, Lei Tao, Fu-Yu Hsieh, Yen Wei, Ing-Ming Chiu, Shan-hui Hsu
First published: May 07, 2015
https://doi.org/10.1002/adma.201500762
Abstract
An injectable self-healing hydrogel (approx. 1.5 kPa) has been developed to heal defects in the nervous system. Neurosphere-like progenitor cells proliferate within the hydrogel and differentiate into neuron-like cells. In a zebrafish injury model, injection of the hydrogel partially restored central nervous system function, and injection of the hydrogel containing neurospheres was found to significantly restore it. Thus, self-healing hydrogels have the potential to repair the central nervous system.
