Natural and synthetic biodegradable polymers: different scaffolds for cell expansion and tissue formation

Natural and synthetic biodegradable polymers: different scaffolds for cell expansion and tissue formation

Int J Artif Organs 2014; 37(3): 187 - 205

Article Type: REVIEW



Annalia Asti, Luciana Gioglio


The formation of tissue produced by implanted cells is influenced greatly by the scaffold onto which they are seeded. In the long term it is often preferable to use a biodegradable material scaffold so that all the implanted materials will disappear, leaving behind only the generated tissue. Research in this area has identified several natural biodegradable materials. Among them, hydrogels are receiving increasing attention due to their ability to retain a great quantity of water, their good biocompatibility, their low interfacial tension, and the minimal mechanical and frictional irritation that they cause. Biocompatibility is not an intrinsic property of materials; rather it depends on the biological environment and the tolerability that exists with respect to specific polymer-tissue interactions. The most often utilized biodegradable synthetic polymers for 3D scaffolds in tissue engineering are saturated poly-a-hydroxy esters, including poly(lactic acid) (PLA) and poly(glycolic acid) (PGA), as well as poly(lactic-co-lycolide) (PLGA) copolymers. Hard materials provide compressive and torsional strength; hydrogels and other soft composites more effectively promote cell expansion and tissue formation. This review focuses on the future potential for understanding the characteristics of the biomaterials considered evaluated for clinical use in order to repair or to replace a sizable defect by only harvesting a small tissue sample.

Article History


Financial Support: This work was supported by the University of Pavia, Pavia, Italy.
Conflict of Interest: The authors state that there are no conflicts of interest.

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Tissue organ repair has been the ultimate goal of surgery from ancient times to the present: reconstruction using gold in cranial defects dates back to 2 000 BC, and tissue grafting has been used since at least the 1660s (1). Tissue engineering emerged in the early 1990s to address limitations of tissue grafting and alloplastic tissue repair. The first use of the term “tissue engineering” was in reference to an organization of an endothelium-like structure that was observed on the surface of a polymethylmethacrylate (PMMA) ophthalmic prosthesis (2). Two early examples were the growth of chondrocytes on a polyglycolic acid (PGA) mesh (3) and the culture of hepatocytes in hollow fibers (4). During the seven years that have passed since the initial report of these studies, the principles of tissue engineering have been applied to virtually every organ system in the body.

The roots of tissue engineering can be traced to the field of biomaterials; in many cases, the objective was to use materials that were as inert as possible, and therefore not degradable or harmful to the host (5). The concept of tissue engineering is to transplant a biofactor (cells, genes, and/ or proteins) within a porous degradable material known as a scaffold. The biofactors, which include stem-cell and gene therapy approaches, are used to stimulate tissue repair. Far from being a passive component, scaffold material and porous architecture design play a significant role in tissue regeneration by preserving tissue volume, providing temporary mechanical function, and delivering biofactors (1).

Physical characteristics are certainly the most important factors to consider when scaffolds are applied for tissue reconstruction. Scaffold materials can be synthetic or biologic, degradable or non-degradable, depending on the intended use (6). The cellular component is necessary for the generation of new tissue through production of extracellular matrix (ECM) and is responsible for the long-term maintenance of this matrix. The interaction of two components, such as the coordination of polymer degradation rates with cellular biosynthetic rates and the cell-seeding characteristic of the polymer, is critical for the success of an engineered tissue construct. As in native tissue, cells within a construct respond and adapt to the physical and biological stimuli to which they are exposed in vivo.

The second step is to expand the cell population in vitro; the ease or difficulty of expansion is highly dependent on the cell type. In this process, it is important to ensure that the expanded cell population retains its phenotypic function. The issue of phenotype expression and dedifferentiation has led to the investigation of pluripotent stem cells as a source for engineered tissue (5).

The geometric arrangement of the local environment is a critical factor in determining cell behavior. The chemical composition and the subcellular organization of focal adhesion molecules differ for cells grown in a 3D matrix compared with those grown on a flat plate in two dimensions (7, 8). Tissue neogenesis in engineered constructs may involve the same processes present in tissue development during embryogenesis. This is supported by studies in which periosteal cells on a polyglycolic acid (PGLA) scaffold appeared to first generate hypertrophic cartilage prior to mineralized bone formation (9), reminiscent of endochondral ossification observed during skeletal morphogenesis.

The scaffold can be opened or closed. Open cell scaffold systems are implanted in the body and become completely integrated with the host tissue. In a closed construct system, cells are isolated from the body by a membrane that permits nutrient and gas exchange while acting as a barrier for large entities such as antibodies and immune cells (10).

A variety of biodegradable materials have been used for tissue scaffolds, including ceramics and polymers; polymers have seen widespread use as scaffold materials because of their good processing characteristics. Typically, polymer scaffolds are in the form of fibrous meshes, porous sponges or foams or hydrogels. The mechanical properties of the scaffold must be such as to ensure that it does not collapse during handling and during the patient’s normal activities (11-12-13). These materials have a degradation time ranging from very short (days) to quite long (several months).

Scaffolds fabricated by methods such as solvent casting/particulate leaching contain pores that reflect the shape and size of the particulates used (the process is based on the dispersion of a porogen agent such as salt or sugar in a liquid particulate or powder-based material) but do not allow for the predetermination of the internal scaffold architecture or pore connectivity (14).

Electrospinning is a promising method for producing artificial ECM tissues for supporting cells such as MSCs (15). The process is controlled by a high intensity electric field created between two electrodes bearing electric charges with opposite polarities. This technique is technically feasible for the fabrication of filaments ranging from the nanometer to micrometer scale (15). In contrast, rapid prototyping technologies, originally developed for the manufacturing industry, provide exceptional spatial control over the polymer architecture. Other rapid prototyping techniques such as fused deposition modeling (FDM) and stereolithography are also being explored for scaffold fabrication (16). While each method presents distinct advantages and disadvantages, the appropriate technique must be selected to meet the requirements for the specific type of tissue (6).

Porosity is of major importance for materials used in regenerative medicine. Micro-porosity is a less explored field, although a combination of macro-pores and interconnected micro-pores (range of 0.5-10 μm) has been shown to promote ingrowth of osteogenic cells into the micro-pores. This provides bone formation on the cell walls, thereby strengthening the material and forming a mechano-sensory network (17). In addition to all these essential properties, for engineering applications of elastic tissues like blood vessel, skin, heart valves, cartilage, tendon, and bladder, elasticity of scaffolds is regarded as an important design parameter. Cellular interaction with biomaterials and the architecture of the scaffold seem to play an important role in adequate vascularization (18).

There are two types of biodegradable polymers: the natural-based materials, including polysaccharides (starch, alginate, chitosan, hyaluronic acid derivatives) or proteins (soy, collagen, fibrin, gels, silk), polynucleotides (DNA, RNA), and a variety of biofibers. A second category of synthetic biodegradable polymer includes poly lactic acid (PLA), poly-glycolic acid (PGA), poly lactic-co-glicolide (PLGA) copolymers, and polycaprolactone (PCL). PLA exist in three forms: L-PLA (PLLA), D-PLA (PDLA), and a racemic mixture of D,L-PLA (PDLLA). Other important categories of materials are bioactive ceramics such as calcium phosphate and bioactive glasses or glass-ceramics (11). This paper considers relevant biodegradable natural polymers with their properties for cell encapsulation, and synthetic biodegradable polymers such as PLA, PGA, PLGA, and PCL for bone tissue engineering.


The ECM provides a three-dimensional (3D) structure that regulates cell and cell-adhesion and provides signals that direct the cellular processes leading to tissue development. Cells attach to the ECM through integrin receptors and engagement of this receptor initiates multiple intracellular signaling cascades that regulate cell survival, proliferation, and differentiation (19, 20). The composition of ECM (fibronectin, laminin) determines which integrin receptors are involved in binding and impacts the signaling that leads to tissue formation. The link between integrin binding and tissue formation has been difficult to characterize; however, the development of tissue-engineering matrices that provide a 3D support and present defined ECM molecules can regulate the receptors used for cell-matrix adhesion and impact cell-cell cohesion (21). The culture of one or more cell types within these matrices can provide a model system to investigate the signals that promote tissue formation. The continuous cross talk between cells and the surrounding matrix environments directs the critical processes in tissue development and maintenance of the cellular phenotype. Tissue engineering matrices may be useful for the development of systems. For example, the in vitro maturation of ovarian follicles, which are needed to preserve reproductive potential for women faced with infertility resulting from chemotherapy or other ovarian disorders (premature ovarian failure) (22-23-24).

In addition to maintaining the cell-cell connections, cells within three-dimensional environments utilize different integrins and form distinct adhesion with the ECM compared to two-dimensional (2D) systems, (25) which can affect both cellular organization and tissue function. The transformation from the compact form to the extended fibrillar form of fibronectin, a highly regulated process termed fibrillogenesis, requires application of mechanical forces generated by cells. Cells bind and exert forces on fibronectin through transmembrane receptor proteins of the integrin family, which mechanically couple the actin cytoskeleton to the ECM via an elaborate adhesion complex (19, 20, 25).


Natural polymers, such as collagen, glycosaminoglycan, chitosan, starch, hyaluronic acid, and alginate (Tab. I) are comparatively weaker, softer materials than ceramics but offer the advantage of flexibility to adapt their shape to required forms through a variety of molding and casting techniques. Moreover, natural polymers usually contain specific molecular domains that can support and guide cells at various stages of their development and can thus enhance biological interaction of the scaffold with the host tissue (15). Naturally occurring scaffolds composed of ECM proteins offer promising alternatives for tissue repair and regeneration. Important examples are small intestinal submucosa, acellular dermis, the bladder acellular matrix graft, and the amniotic membrane. These types of scaffolds have been shown to promote rapid interaction with the surrounding host tissue, to induce the deposition of cells and additional ECM, and to accelerate the process of angiogenesis (18).


Polymer Biomedical Applications Ref. No
HA = Hyaluronic acid; HAP = hydroxyapatite.
Alginate scaffolds materials engineering, supporting matrix cells, drug delivery techniques of artificial insemination 6, 12, 27-29, 42, 46-52, 53-55, 58-69, 70, 81
Alginate/Chitosan enhance biological performances 10, 56, 76
Chitosan wound healing, drug delivery, tissue engineering, skin tissue regeneration 28, 29, 35-37, 44 71-77, 104, 114
Chitosan/Collagen artificial skin 71
Hyaluronic acid wound repair, angiogenesis, bone and cartilage tissue repair and regeneration 10, 78-82, 37-40, 109, 122
HA/HAP bone tissue engineering 110
Collagen cartilage and bone scaffolds tissue engineering, artificial skin, blood vessel 32, 52, 83-85, 118
Gelatin delivery of growth factor, neovascularization 15, 33, 84


Hydrogels have structural similarity to the macromolecular-based components in the body and are considered biocompatible. Among the materials used for regenerative applications, hydrogels (HGs) are receiving increasing attention due to their ability to retain a great quantity of water, their good biocompatibility, their low interfacial tension, and the minimal mechanical and frictional irritation that they cause (26, 27) all of which are appealing features from the perspective of bioenvironmental mimicking. The structure of a hydrogel is morphologically similar to that of the ECM when used as an encapsulation medium (28).

The term hydrogel is composed of “hydro” (water) and “gel,” and it refers to aqueous (water-containing) gels, or to be more precise, to polymer networks that are insoluble in water, where they swell to an equilibrium volume but retain their shape (26). The hydrophilicity of the network is due to the presence of chemical residues such as hydroxylic (-OH), carboxylic (-COOH), amidic (-CONH-), primary amidic (-CONH2), sulfonic (- SO3H), and others that can be found within the polymer backbone or as a lateral chain. The insolubility of the gel in water is due to the presence of a 3D network, in which an equilibrium exists between forces that are dispersive (acting on hydrated chain) and those that are cohesive (preventing further penetration of water) (27, 29, 30). These structures expand when they are put into contact with water (20); the network properties and swelling characteristics are further related to the mass transport characteristics of hydrogels (27, 29, 31).

Common hydrogel substrates include the copolymers of polyethylene oxide and polypropylene oxide (Pluronic® types), and natural polymers including alginate (6), collagen (32), gelatine (33), fibrin (34), chitosan (35-36-37) and hyaluronate (37-38-39). The properties of hydrogels have made them attractive in the medical field, such as in controlled release drug delivery (29, 40). As an alternative to hydrophobic scaffold materials, hydrogels offer an attractive opportunity to seed living cells and other biological species during the fabrication process of scaffolds. Besides their appropriate mechanical properties and mass transport characteristics, the degradation of hydrogel is essential for many tissue-engineering applications (41, 42). Ionically cross-linked hydrogels, such as alginate gels, normally undergo slow dissolution due to complexation of divalent cations or gradual exchange with monovalent cations present in the environment (43). The preparation of ionic chitosan hydrogels avoids the use of catalyst or toxic reagent; these conditions compatible with the human body have led to the development of injectable solutions that exhibit a sol/gel transition upon injection in the body (43).

The formation of covalent chemically cross-linked hydrogels occurs via covalent bonding between polymer chains. However, this approach requires the chemical modification of the primary structure which could alter its initial properties particularly if amino groups are involved in the reaction. Covalently cross-linked hydrogels are the only systems characterized by a permanent network, due to their irreversible chemical links (44). Therefore, they exhibit good mechanical properties and can overcome dissolution, even in extreme pH conditions, while the other types of hydrogels are more labile. To obtain hydrogels with these interesting characteristics, the use of covalent crosslinkers is necessary. However, most of the crosslinkers used thus far are either known to be relatively toxic, or their fate in the human body is unknown and/or there is a lack of data concerning their biocompatibility (44).


Alginate is a polysaccharide extracted from seaweeds that is composed of (1-2-3-4) linked β-D-mannuronic acid (M units for mannuronic) and α L-guluronic acid (G units for guluronic) monomers along the polymer backbone (Fig. 1), whose sequences and relative percentage within the polymer depends on many different factors. Thus alginate can be cross-linked with divalent ions (calcium, barium, magnesium, strontium) and the obtained gels may be further stabilized by a coat (10-30 μm) of polymer such as poly-L-lysine or chitosan followed by an additional layer of alginate. Calcium is one of the primary ions that influence bead stability (45); it has been used for scaffold and matrix production, cell encapsulation, cytokine, growth factor, gene, and DNA encapsulation and delivery (46). Poly α-(1, 4)-linked L-guluronate-β-(1, 4)-linked D-mannuronate have hydrogen bonds between the carboxyl group on the mannuronate and the 2-0H and 3-0H groups of the subsequent guluronate; however, the differing degrees of freedom of the two residues gives greater overall flexibility than the poly β-(1,4)-linked D-mannuronate chains. The guluronate blocks of one polymer then form junctions with the guluronate blocks of adjacent polymer chains in what is termed the “egg-box model” of cross-linking, resulting in a gel structure (46).

Structural unit. Alginates are linear unbranched polymers containing β-(14)-linked D-mannuronic acid (M) and α-(14)-linked L-guluronic acid (G) residues.

Gelling depends on the ion binding Mg 2+<< Ca 2+< Sr 2+< Ba 2+; high G content produces strong brittle gels with good heat stability (except if present in low molecular weight molecules) (46, 47), whereas high M content produces weaker more elastic gels with good freeze-thaw behavior and high MGMG content zips with Ca 2+ ions. The most common method of incorporating bioactive molecules or cells into alginate matrices is via extrusion, in which an alginate suspension is extruded through a needle to form droplets that fall into a solution containing divalent cations, which cause alginate crosslinking (48-50). The bead shape contributes to the biological and transport properties of alginate beads: a non-spherical bead influences the risk of an adverse foreign body reaction, which could lead to the eventual failure of the system (45, 51). One critical drawback of ionically cross-linked alginate gels is their limited long-term stability in physiological conditions, because these gels can be dissolved due to the release of divalent ions into the surrounding media caused by exchange reactions with monovalent cations. Therefore, covalent cross-linking with various types of molecules and different cross-linking densities have been attempted to precisely control the mechanical and/or swelling properties of alginate gels (52).

Alginate in tissue engineering

Alginate has been widely used for microencapsulation (Figs. 2A and B), for immobilization of enzymes or cells for bioreactors, and also for tissue engineering (53). It has been mainly utilized as a delivery vehicle or supporting matrix cell via encapsulation techniques (54). Various techniques have been demonstrated for producing alginate beads of controlled dimensions for cell encapsulation (55). Alginate degradation is not carried out through mammalian cell digestion; instead divalent calcium ions slowly diffused out of the hydrogel allowing alginate molecules to be excreted in the urine. One of the inconveniences of alginate is that there is no specific interaction between mammalian cell alginate in gel. Alginate has also been mixed with other materials such as chitosan to enhance its biological performance (10, 56).

Scanning electron microscopy (SEM). A) Barium alginate polymer at great magnification, bar = 10 μm; B) Barium alginate hydrogel inner layer; bar = 50 μm; C) Granulosa cells (GCs) inside the alginate bead. GCs cultured in 3D barium alginate retain a rounded form similar to cells of the antral follicle, bar = 10 μm.

Cell encapsulation

Cell encapsulation is a strategy whereby a pool of live cells is entrapped within a semipermeable membrane. The first scientific publication describing the principle of bioencapsulation was by Chang (57); the original method consisting in the generation of a controlled-size droplet (Figs. 3A and B), followed by the process of stabilizing the interface and creating a membrane around the core. The breakthrough in applying Chang’s principles of bioencapsulation came with the work of Lim and Sun (49). For the first time, a natural biocompatible polymer, alginate, was employed for Langherans islets encapsulation with the aim of preventing the rejection of the transplanted cells in type I diabetes-affected rats (58). The first application of cell encapsulation in mammalian reproduction was made by Nebel et al (58), who devised a technique for bovine sperm encapsulation in calcium alginate and polyamines. Nebel (58) introduced the idea of spermatozoa controlled release and the targeting of the female reproductive system for artificial insemination. Starting from these works, the application of cell encapsulation in reproductive biotechnologies has expanded to other fields (59, 60).

Alginate bead and PLA scaffold. A) A capsular-like structure containing living cells. B) Light microscopy of cells inside the alginate bead; alginate layer is visible in the upper part of the picture, mag. 40x. C) PLGA unseeded scaffold obtained using the porogen particle leaching method; the bioresorbable scaffold should present mechanical properties equivalent to those of the host tissue until the bioresorbable scaffold matrix is substituted by the new tissue.


The encapsulation procedure similar to that of Lim and Sun (49) involved three steps: a) the production of a calcium alginate cell-containing matrix; b) the formation of a semipermeable membrane by interfacial polymerization with a multivalent polyamine; c) the liquefaction of the semi-solid matrix by chelation of calcium with sodium citrate. The use of a biocompatible polymer such as alginate (Figs. 2A and B) led to the maintenance of optimal cell viability levels both in vivo and in vitro. The capsule membrane is permeable to small molecules such as glucose and oxygen but impermeable to large molecules such as immunoglobulins (49). The most common method of incorporating bioactive molecules or cells into alginate matrices is via extrusion, in which an alginate suspension is extruded through a needle to form droplets that fall into a solution containing divalent cations, which causes alginate crosslinking (9, 23, 48). In techniques of artificial insemination, capsules protected male gametes from uterine macrophage phagocytosis and promoted bioadhesion of the delivery system, preventing semen reflux (58, 61-62-63).

Alginate beads for granulosa cells and for semen encapsulation

During follicle evolution, oocytes secrete several molecules, commonly known as oocyte-secreted factors (OSFs), that guide the granulosa cells (GCs) towards growth, preventing cell death and inhibiting luteinization via LH receptor suppression (64). In turn, the oocytes induce GC expression of molecules related to oocyte maturation, inhibiting their luteinizing ability (65). GC cultured in 3D in barium alginate capsules retain a rounded form similar to that of the cells in the antral follicle (66) (Fig. 2C). Peptides including the sequence arginine-glycine-aspartic acid (RGD) have been extensively used as model adhesion ligands (46, 67). Researchers (68, 69) have shown that a specifically synthetic matrix composed of Arg-Gly-Asp (RGD)-modified alginate supported granulosa cell adhesion and spreading, and increased estradiol and progesterone secretion.

Torre et al (12) proposed a one-step reverse technique for encapsulating swine spermatozoa in barium alginate. Barium chloride is added to the semen and the resulting suspension is dropped into a sodium alginate solution; barium ions contained in the ejaculate diffuse out of the droplets and when they reach the interface, they react with the alginate chains, leading to the formation of a barium alginate gel membrane around the semen droplet. The thickness of the gel membrane increases until the diffusion of the barium ions through the semen droplets ends (68).

Alginate in drug delivery

Controlled-release drug delivery systems are designed to give a reproducible and kinetically-predictable release of a drug substance. Alginates may be utilized in dosage forms designed for either type of drug release; the formulations employ a chemical or physical “barrier” to provide a controlled release of the drug. Many formulation techniques have been used to build the barrier into the peroral dosage form, e.g., the coating of a core containing the active ingredient or the embedding of the active ingredient in a polymer matrix. Alginate has mainly been applied in systems based on diffusion (70).


Chitosan is structurally similar to glycosaminoglycans (GAGs) and is degradable by enzymes in humans. It is a linear polysaccharide of (1-2-3-4) linked D-glucosamine and N-acetyl-d-glucosamine residues derived from chitin. Once dissolved, chitosan can be gelled by increasing the pH or extruding the solution into a non-solvent, by glutaraldehyde crosslinking, UV hyaluronic acid. After cellulose, chitin is probably the most common polymer found in invertebrates as crustacean shell or insect cuticles, but also in some mushroom envelopes, the green algae cell wall, and yeasts (71). In contrast to most other biopolymers, chitosan has a positive electrical charge below the pKa value of the amino group which is around pH 6.0, this makes it electrostatically attach to most living tissues that contain negatively charged surface matrices (72). The presence of amino groups in the chitosan structure gives to this polymer many peculiar properties; chitosan with protonated amino groups becomes a polycation that can form ionic complexes with natural or synthetic anionic species such as lipids, proteins, DNA, and some negatively charged synthetic polymer (71).

Chitosan can be easily processed into hydrogels, membranes, nanofibers, beads, micro/nanoparticles, scaffolds, foams, and sponges for biomedical applications (drug delivery), wound healing, and tissue engineering. (72). The antifungal and bactericidal properties of chitosan, its permeability to oxygen are very important characteristics associated with the treatment of wound and burns. Chitosan is able to form a gel without the need of any additive; the process is based on the neutralization of chitosan amino group and thus the repulsion between chitosan chains. The formation of hydrogel occurs via hydrogen bonds, hydrophobic interactions, and chitosan crystallites (72).

Chitosan in tissue engineering

Chitosan is a promising polymer for tissue engineering due to its favorable properties: it is non-toxic, non-allergenic, mucoadhesive, biocompatible, and biodegradable, and also accelerates cell proliferation. These properties make chitosan an outstanding candidate for biomedical applications, notably for dental and bone implants, cartilage, and artificial skin (71).

A number of techniques, such as phase separation, self-assembly, and electrospinning (15) have been developed to fabricate nanofibrous scaffolds with unique properties. Thermal-induced phase separation is based on the decrease in solubility associated with temperature increase. Among these techniques, electrospinning technology has become popular in recent years because it is a simple, rapid, efficient, and inexpensive method for producing nanofibers by applying a high voltage to electrically charged liquid (15). To improve the wound-healing properties, chitin and chitosan-based membranes have been developed with different types of polymers such as alginate, hyaluronic acid, polyethylene glycol diacrylate, poly(vinyl alcohol), γ-poly (glutamic acid), and 2-hydroxyethyl methacrylate. The type of chitosan has significant effects on scaffold properties; the molecular weight of chitosan has been shown to influence swelling and biodegradation properties as well as cell proliferation (73). A decrease in molecular weight resulted in lower water uptake and favored dissolution (74).

Covalently crosslinked chitosan hydrogels are interesting for applications where a well-shaped system is required, for the formation of implants or bandages, or for the preparation of gel particles for oral administration exhibiting pH-dependent drug release in acidic conditions (44, 75).

Chitosan microspheres for drug delivery

The microencapsulation technique has also been applied to making controlled- release microspheres of natural biopolymers. It is important to control the factors that influence the diameter and drug-loading rate of microspheres. These include the mechanical stirring time, the volume of added glutaraldehyde solution, the ultrasonic emulsification time, the water-oil volume ratio, the concentration of chitosan and the amount of drugs. Whatever the type of structure, networks containing covalently cross-linked chitosan are considered to be porous. Due to these pores, chitosan hydrogels can be used as drug delivery systems from which drugs are released by diffusion (44).

Chitosan nanofibers, sponges, films, and scaffolds for wound healing

Burn healing is one of the most significant problems of modern surgery due to the high percentage of burns compared to other traumas. Chitosan nanofiber dressings provide effective absorption of exudate, ventilation of the wound, protection from infection, and stimulation of the process of skin tissue regeneration. To create a moist environment for rapid wound healing, a chitosan-PVA-alginate film with sustained antibacterial capacity had been developed by the casting/solvent evaporation method (76). A porous collagen-chitosan skin was developed as a scaffold for the reconstruction of skin in vitro: this artificial skin promoted the remodeling of an ECM similar to normal dermis (71).

Chitosan sponges for wound healing and bone tissue engineering

While mainly used as wound healing materials to aid tissue regeneration, chitosan sponges are also applied in bone tissue engineering. Sponges are soft and flexible materials with a well interconnected micropore structure. Due to their unique structural features, they have good fluid absorption capability, cell interaction and hydrophilicity but they are mechanically too weak to maintain the desired shape until newly formed tissue matures. Sponges are mainly obtained by freeze drying or lyophilization, a process that consists in freezing a solution of chitosan followed by sublimation of the solvent under reduced pressure (71). The method is based on the formation of ice crystals that induce porosity through ice sublimation and desorption. It is possible to control the porosity level of the foams by varying the freezing time and the annealing stage. The main difficulty associated with this process is to ensure structural stability and adequate mechanical properties of the porous constructs after subsequent hydration (77).

Hyaluronic acid (HA)

HA is a glycosaminoglycan present in all vertebrates. It is a major constituent of the ECM, in the vitreous humur of the human eye, in synovial joint fluid, and in the matrix produced by the cumulus cells around the oocyte prior to ovulation. HA has unique viscoelastic properties as well as good biocompatibility and biodegradability that make it a good candidate for tissue engineering applications (78). One of the key advantages of using HA gels is that their degradation can be mediated by hyluronidase, an enzyme secreted by the mammalian cell type (78).

The most common modification of hyaluronan is crosslinking to form a hydrogel; crosslinking has been accomplished under acidic, neutral, and alkaline conditions. HA structural and biological properties mediate several processes involved in wound repair, such as cellular signaling, angiogenesis, morphogenesis and matrix organization. This, coupled with the ability of HA-based scaffolds to provide an open, hydrated structure for the passage of nutrients, make them ideal candidates for tissue regeneration and repair. Its properties, both physical and biochemical, in solution or hydrogel form, are extremely attractive for various technologies involved with body repair.

Hyaluronic acid in tissue engineering

Hyaluronic acid offers great practical potential as a scaffolding material, due to its ability to maintain a hydrated environment conducive for cell infiltration. HA-based hydrogels are ideal as wound grafts to treat chronic wounds or wounds in patients with impaired healing such as diabetic patients (10).

Researchers have developed scaffolds based on hyaluronic acid in the form of hydrogels, sponges, and meshes (77). Electrostatic repulsion between the negative charges is countered not only by hydrophobic interactions but also by H-bonding between acetamido and carboxylate groups (non-covalent reaction). The hyaluronan esters can be extruded to produce membranes and fibers, lyophilized to obtain sponges, or processed by spray-drying, extraction, and evaporation to produce microspheres. These materials have been used for growth of cultured human fibroblasts and for culture of chondrocytes and bone marrow-derived mesenchymal cells for repair of cartilage and bone defects. HA-based hydrogels are able to alter the microenvironment of the injured heart to demote scar tissue formation and induce neovascularization, resulting in an improvement of the cardiac function (79).

HA for bone and cartilage tissue repair and regeneration

The presence of HA is sufficient to increase both chondrocyte proliferation and protein secretion. In general, initial proliferation is important to ensure an adequate cell population, but ultimate cartilage formation depends on synthesis of GAGs and type II collagen. The encapsulation of cells such as auricular chondrocytes for cartilage regeneration has been investigated with photopolymerizable HA-based scaffolds and freeze-dried HA/collagen-based scaffolds (80).

Other rapid prototyping techniques such as fused deposition modeling (FDM) and stereolithography are also being explored for scaffold fabrication (80). A hyaluronic acid dextran hydroxyethyl methacrylate (HA-dex HEMA) material has been produced by bioprinting scaffolds using a layer-by-layer deposition technique, where scaffolds are formed by deposition of cell-laden material (81). Three-dimensional printing produces components by ink-jet printing a binder into sequential powder layers (81). This is repeated until the entire part, such as a porous scaffold, is fabricated. Leukers et al used a 3D printing technique to fabricate scaffolds based on HA with a complex internal structure (82). A crucial advantage of 3D printing is the wide range of materials that can be used, from synthetic and natural polymers to ceramics, as well as the availability of the material in powder form (polymeric, ceramic and composites) used in bone scaffold engineering (10).


Collagen is the primary protein component of animal connective tissues. It is composed of different polypeptides, which contain mostly glycine, proline, hydroxyproline, and lysine. The flexibility of the collagen chain depends on the glycine content (83). Collagen is enzymatically degradable and has unique biological properties; it has been extensively investigated for biomedical applications (84) due to its special characteristics, such as biodegradability and weak antigenecity (85).

Collagen-based materials are considered to be a favorable biomaterial for both cartilage and bone scaffolds due to the fact that collagen is the major matrix component in ECM; collagen type II is the major component in articular cartilage; and collagen type I is the major component in bone (15). Chemical cross-linking of collagen using glutaraldehyde or diphenylphosphoryl azide can improve the physical properties. However, these gels are still short of physical strength, potentially immunogenic, and can be expensive.

Collagen has been used as a tissue culture scaffold or artificial skin due to the ready attachment of many different cell types and its cell-based degradation. The attachment of cells to collagen can be altered by chemical modification, including the incorporation of fibronectin, chondroitin sulfate, or low levels of hyaluronic acid into the collagen matrix. Thus, collagen gels have been utilized for reconstruction of the liver, skin, blood vessels, and the small intestine (52). There are some intrinsic relationships between collagen and many diseases such as rheumatoid arthritis and systemic sclerosis (84). By denaturation and/or physical-chemical degradation of collagen, a high molecular weight polypeptide is produced, called gelatin (84).


Gelatin is a derivative of collagen, formed by breaking the natural triple-helix structure of collagen into single-strand molecules. Gelatin easily forms gels by changing the temperature of its solution. It has been used in many tissue engineering applications due to its biocompatibility and ease of gelation. Gelatin gels have also been utilized for delivery of growth factors to promote vascularization of engineered new tissue. However, the weakness of the gels has been a problem, and a number of chemical modification methods have been investigated to improve the mechanical properties of gelatin gels (84). Gelatin has been used for coatings and microencapsulating various drugs for biomedical applications; it has also been employed for preparing biodegradable hydrogels.

A novel osteochondral scaffold based on a ceramic-gelatin assembly for articular cartilage repair and a porous ceramic for bone repair was developed. This scaffold design was motivated by the problem of achieving improved joining strength at the interface between cartilage and bone (15).

Synthetic biodegradable polymers

Synthetic polymers represent the largest group of biodegradable polymers, and they can be produced under controlled conditions; they exhibit predictable and reproducible mechanical and physical properties such as tensile strength and an elastic modulus and degradation rate; they are often cheaper than biologic scaffolds (6) and should ensure an optimal interaction with endothelial cells to promote angiogenesis. The more common polymers used in fibrous meshes and foams include the linear polyesthers, poly-lactic acid (PLA), poly-glycolic acid (PGA) poly lactic-co-glycolide (PLGA) copolymers (86), and polycaprolactone (PCL) (15) (Tab. II). In general, synthetic polymers have limitations in bioactivity because of their hydrophobic surface (15).


Polymers Biomedical Applications Ref. No
PLA = Poly(lactic acid); PGA = poly(glycolic acid); PLGA = poly(lactic-coglycolide); HAP = hydroxyapatite; HA = hyaluronic acid; PCL = Poly(ε-caprolactone).
PLA, PGA, PGLA copolymers Bone internal fixation device, resorbable sutures, bone tissue engineering, artificial vessels 3, 9, 16, 73, 86-95, 96-98, 100-104, 114, 119
PLA/HAP bone tissue engineering 15, 82, 97, 98
Bioss/PLA bone defects in periodontal and maxillofacial surgery 104-107
PCL bone and cartilage repair, surgical sutures bone regeneration 15, 85, 99, 108
PLGA/HA delivery of growth factor, vascularization of engineering new tissue 33

PLA (poly-lactic acid), PGA (poly-glycolic acid), PLGA (poly-lactide-co-glycolide)

These are the most often utilized biodegradable synthetic polymers for 3D scaffolds (Fig. 3C) in tissue selection of the material and consequently of the nanostructure and the architecture of scaffolds. Typical scaffold designs have included meshes, fibers, sponges, and foams; these designs are chosen because they promote uniform cell distribution, diffusion of nutrients, and the growth of organized cell communities (6). The use of these polymers began in pharmaceutics and was consolidated in that sector, however, they have been widely studied in the last few years, mainly due to their good biocompatibility, their chemical versatility, and good biological performance. They also do not entail the danger of immunogenic reactions or possibility of disease transmission.

They biodegrade by random chain scission, generating monomers of lactic and glycolic acid that are eliminated through the metabolic pathways. The intrinsic properties of the raw materials play a strategic role in their production, structure, and morphology and, consequently, in the functional performance of the polymer scaffold (86). To act as an artificial ECM (Fig. 4A), the structure and surface morphology of the scaffolds have to meet general requirements specific for the targeted tissue: (I) interconnected pores to ensure cell growth and nutrients and metabolic waste transport flow; (II) three-dimensional architecture; (III) suitable mechanical properties; (IV) suitable surface chemistry; (V) controllable biodegradation and bioresorbability (87) (Figs. 4A and B) The scaffold shape should also facilitate cell seeding and attachment and promote cell proliferation and differentiation (9) (Fig. 3C). Moreover, the bioresorbable scaffold should present mechanical properties (strength and stiffness) equivalent to those of the host tissue until the bioresorbable scaffold matrix is substituted by the new tissue (88).

Scanning electron microscopy (SEM). A) PLA/Chitosan scaffold; it’s possible to observe cells embedded in the ECM, bar = 20 μm. B) PLA scaffold with cells detectable on the surface; bar = 100 μm.

Pore size refers to the distance between solid sections of the porous matrix; it is typically reported as the diameter of circular pores or the major axis for noncircular pores. Pore size affects cell binding, migration depth of cellular in-growth, cell morphology, and phenotypic expression (89). Scaffolds with a mean pore size ranging from 20 μm to 1500 μm have been used in bone tissue engineering applications. By facilitating capillary formation, pores greater than 300 μm lead to direct osteogenesis, while pores smaller than 300 μm can encourage osteochondral ossification (90, 91). The elasticity of the microporous scaffold increases as the number of pores within the scaffold increases (92). The optimized size of the pore interconnection of cartilaginous scaffolds should be less than that of subchondral bone scaffolds because articular cartilage is normally fed by articular fluid, whereas bone is fed by the nutrients from blood circulation (73).

Poly(lactic acid) (PLA) and poly(glycolic acid) (PGA), as well as poly(lactic-coglycolide) (PLGA) copolymers. PLA exists in three forms: L-PLA (PLLA), D-PLA (PDLA), and a racemic mixture of D, L-PLA (PDLLA) (93). The chemical properties of these polymers allow hydrolytic degradation through de-esterification. The body already contains highly regulated mechanisms for completely removing monomeric components of lactic and glycolic acids. PGA is converted to metabolites or eliminated by other mechanisms, and PLA can be cleared through the tricarboxylic acid cycle. Due to these properties PLA and PGA have been used in products and devices, such as degradable sutures which have been approved by the US Food and Drug Administration (94). PLA and PGA can be processed easily and their degradation rates and physical and mechanical properties are adjustable over a wide range by using various molecular weights and copolymers. However, these polymers undergo a bulk erosion process that can cause scaffolds to fail prematurely; abrupt release of these acidic degradation products can cause a strong inflammatory response (95, 96). Different factors affect the degradation kinetics, such as: chemical composition and configurational structure, processing history, molar mass (Mw), polydispersity (Mw/ Mn), environmental conditions, stress and strain, crystallinity, device size, morphology (e.g., porosity) and chain orientation, distribution of chemically reactive compounds within the matrix (97, 98) and overall hydrophilicity.

Polyester PLGA is a copolymer of PLA and PGA: it has a wide range of degradation rates, the degradation kinetics being governed by both hydrophobic/hydrophilic balance and crystallinity (97). Poly(e-caprolactone) (PCL), on the other hand, can take several years to degrade in vivo (99). Polymeric scaffolds that undergo bulk degradation tend to break down the internal structure of the scaffold, thus reducing the molecular mass. A polymeric scaffold that primarily undergoes surface degradation can be described as similar to the dissolution of soap. Due to their good mechanical properties, PGA and PLLA have been used as bone internal fixation devices. They also have excellent fiber forming properties and thus PGA was used to prepare resorbable sutures and PLLA to replace ligament and non-degradable fibers (100).

Synthetic biodegradable polymers for bone tissue engineering and for artificial vessel

The combination of degradable polymers and inorganic bioactive particles yields the highest achievable mechanical and biological performances in hard tissue today (6, 100). The replacement of diseased bone tissues has taken a variety of pathways: metals, ceramics, polymers and bone itself, none of which has proven ideal for tissue engineering (101). In cell culture studies with osteoblast-like cells, the surface characteristic of the material plays an important role because osteoblastic cells require a supportive matrix in order to survive (102). With regard to architecture, macroscopic 3D shapes are typically defined by traditional processes such as extrusion, melt molding, and solvent casting (2-3-4-5-6). Foaming during melt extrusion or injection molding is based on the use of physical or chemical blowing agents that are responsible for inducing porosity. Research on these processing routes has mainly been focused on the use of chemical blowing agents. These scaffolds have been reported to be highly biocompatible and to exhibit adequate properties, both in terms of porosity and pore geometry, for supporting cell growth and apparent bone formation (77). Material microstructure, in contrast, is often controlled by process parameters such as the choice of solvent in phase separation, doping with particulate leachants, gas foaming, woven fibers, and controlled ice crystal formation and subsequent freeze- drying to create pores (77). However, these scaffolds lack the well-defined organization that is found in most tissues in vivo.

It is difficult to deliver cells into porous scaffolds made of biodegradable polymers such as PLA and PLGA because of their hydrophobic surfaces. Thus, the combination of bioactive inorganic materials such as HA, TCP, and bioactive glasses with a polymer matrix-forming composite material should overcome this problem (15, 98).

More intricate scaffolds that contain small pores and features can also be fabricated using lamination techniques casting thin films of poly(DL-lactic-co-gycolic) acid (PLGA) onto microfabricated silicon wafers to create biodegradable membranes containing small trenches that are the inverse of the silicon masters. By laminating the patterned PLGA membranes to each other, channels were formed between the layers to create a scaffold for vascular tissue engineering

Several biocompatible and biodegradable polymers have been used as scaffolds for the construction of artificial vessels (103). Polyglycolic acid (PGA) is most commonly used because its high porosity permits the diffusion of nutrients upon implantation and subsequent neovascularization, and it is easily handled and fabricated into different shapes. Unfortunately, PGA meshes are rapidly bioabsorbed (within 6 to 8 weeks) and unable to withstand systemic pressures. To improve the physical properties of PGA, several copolymers have been produced by combining PGA with polyhydroxyalkanoate (PHA), poly(lactic acid-colysine, poly-4-hydroxybutyrate (P4HB), poly-Llactic acid (PLLA), or polyethylene glycol (PEG). These combinations not only define the mechanical characteristics of the scaffold but also regulate the phenotype of the cells grown on them through cellular interactions with the scaffold material (97).

Bio-Oss / PLA

Bio-Oss is a natural bone substitute made of the mineralized portion of bovine bone. The material is used mainly to fill bone defects in periodontal and maxillofacial surgery and to enable reossification. Bio-Oss is obtained via a proprietary extraction procedure involving denaturation and elimination of the organic matrix of bone. Through its trabecular architecture, pore size (300-500 μm), and high porosity (70-75%) Bio-Oss promotes the invasion of blood vessels and bone cells, thereby inducing revitalization and ossification of the defect (104, 105). Bio-Oss has been used in clinical applications for more than 15 years and has been scientifically investigated. Results show that this material, thanks to the large hydrophilic inner surface area similar to human bone, presents superior handling characteristics. Nevertheless, the cell proliferation rate is not as high as those of other biomaterials in use (106). PLA-coated scaffolds improved cell growth, polymer degradation behavior, extraction and measurement of the ECM proteins of the cultured scaffolds (such as decorin, fibronectin osteocalcin, osteonectin, osteopontin and type-I and type-III collagen), and immunolocalization of bone proteins. Morphological analysis of the scaffolds confirmed the bioactive properties of Bio-Oss/PLA4M, suggesting that it could be a valuable grafting material (107).

Poly (ε-caprolactone) (PCL)

PCL is a versatile synthetic polymer with a low melting point (60°C) which allows easy processing. Mechanical properties and non-enzymatic degradation (by hydrolysis) of PCL can be altered by regulating its crystal structure. However, PCL has a limited bioregulatory activity, hydrophobicity, neutral charge contribution, and susceptibility to bacteria-mediated degradation.

ε-caprolactone and copolymers have been used intensely and are commonly studied materials for biomedical applications in bone and cartilage repair, as surgical suture as well as for drug delivery systems, especially those with longer working lifetimes (85, 108). Electrospun nanocomposite PCL mats were explored for potential bone regeneration applications.

Different approaches have been used to prepare hybrid polymeric scaffolds: blending PCL/chitosan solutions and HA nanoparticles under the determined conditions and then electrospinning them (PCL/chitosan/HA blend scaffolds).

Bioactive ceramics

Another important category of materials consists of bioactive ceramics that have been developed with the aim of increasing bioactive glasses or glass ceramics (10). Hydroxyapatite (HAP), tricalcium phosphate (TCP), and certain compositions of silicate and phosphate glasses (bioactive glasses) and glass ceramics react with physiological fluids and through cellular activity bond to hard and, in some cases, soft engineered tissue. Ceramics are known for their good compatibility, corrosion, resistance, and high compression resistance. The drawbacks of ceramics include brittleness, low fracture strength, low mechanical reliability, lack of resilience, and high density (6). Bioactive ceramic materials such as Bioglass (10), sintered hydroxyapatite (11), and glass-ceramics (12), are now clinically used as bone substitutes but these bioactive ceramics have low fracture toughness and a high Young modulus. These materials are not suitable in high mechanical loading parts, so the applications are limited for the replacement of bone parts of low loads and are used as bone fillers. Investigations into synthetic and natural inorganic ceramic materials such as HA as coated scaffold material have been employed mostly in bone tissue engineering (109, 110).


The characteristics of degradable polymers preferred for implantation have been divided into two main categories: biocompatibility and biofunctionality. Biocompatibility refers to the aspects concerning the absence of toxicity, immunogenicity, carcinogenicity, and thrombogenicity (90). Biofunctionality refers to the aspects of adequate properties (mechanical, physical, chemical, thermal, and biological), and whether it is easy to handle, sterilizable, storable, and reabsorbable. The biocompatibility of the biomaterial is very closely related to cell behavior and particularly to cell adhesion on the biomaterial surface (111-112-113), which can influence cell reaction through changes in the cytoskeleton. It is known that cell behavior and interaction with a biomaterial surface are dependent on properties such as topography, surface charges, and chemistry (114, 115).

As an alternative to hydrophobic scaffold materials, hydrogels offer an attractive opportunity to seed living cells and other biological species during the scaffold fabrication process (Fig. 5). If we consider cells embedded into hydrogels, they probably sense some sort of physical confinement that regulates growth, differentiation, and ECM accumulation. The confinement may be caused by the mechanical properties of the hydrogel itself as well as the pericellular accumulation of ECM macromolecules. The supply of nutrients, oxygen, and bioactive substances as well as the removal of waste products are also affected by the network properties and swelling characteristics of hydrogels (29, 31, 32). As with most biomaterials used for tissue-engineering applications, the initial mechanical properties are not retained over time. During the degradation of hydrogel, the average mesh size and swelling level increases, and the diffusion of macromolecules, e.g., ECM components, is facilitated. Concomitantly with the increase in mesh size, the mechanical properties of the degrading hydrogel decrease significantly (36). Mechanical properties of a hydrogel may be increased by combining the hydrogel with particles of a ceramic material, such as B-tricalcium phosphate, hydroxyapatite, or calcium carbonate (110). However, an increase in radical concentration during polymerization may lead to transport limitations for nutrients and oxygen to the encapsulated cells. It was also shown that lower initial viability was seen with a higher macromere concentration (101). The hydrogels may be suitable for non-adherent cells but not for osteoblasts, which need to adhere to develop and produce the extracellular bone matrix. PEG based hydrogels such as poly(ethylene glycol)-diacrylate (PEGDA) were analyzed for osteoblast cells viability and proliferation (101). The choice of safe, biocompatible covalent crosslinkers is quite limited, which is the main drawback of these systems. Interesting alternatives are emerging, such as genipin. (10).

Optical section of granulosa cells together with oocyte inside alginate barium. m = external alginate membrane; gc = granulosa cells; oo = o ocyte; zp = pellucid zone; c = core.

With regard to bone tissue repair, the advantage of porous materials is their ability to provide biological anchorage for the surrounding tissue via ingrowth of mineralized tissue into the pore space (116). A porous surface implant could improve early implant stability and resistance of mechanical removal (117). The high porosity (65-70%) and the broad pores (diameter of 350 to 550 μm) should be sufficient to enable an ample nutritional supply inside the scaffold. If the pores are too small, cell migration is limited, resulting in the formation of a cellular capsule around the edges of the scaffold and limiting ECM production and neovascularization of the inner areas of the scaffold. This, in turn, can limit the diffusion of nutrients, resulting in a necrotic region within the construct. Conversely, if the pores are too large, there is a decrease in surface area, limiting cell adhesion (88, 117). In a previous study, O’Brien et al (118) showed that the specific surface area decreases with increasing pore size; it is hypothesized that the effect of the specific surface area is due to the ligand density available for integrin-binding after initial seeding (89). Large pore size makes the scaffolds more fragile and decreases surface density, but larger pore size could improve infiltration of new blood vessels into the scaffold, thereby enhancing its vascularization and providing cell survival into the inner part of the scaffold (119). The critical determinant of blood vessel ingrowth is significantly faster in pores with a size greater than 250 mm than in those smaller than 250 mm (18).

Scaffolds with smaller pores have a greater surface area, which provides increased sites for initial cellular attachment post-seeding (83). Therefore maintaining a balance between the optimal pore size for cell migration and specific surface area for cell attachment is essential (91,102).

A study done by Rumpler et al (8) showed that if cells seeded on scaffolds were stained for actin, it become apparent that cells interacted through their cytoskeleton. Actin stress fibers between neighboring cells in the tissue border are aligned parallel to the tissue-fluid interface in such a way that they appear as a single ring-like structure (8). Such stress fibers may be associated with the interaction of neighboring cells via mechanical forces (120). The strong alignment of the stress fibers with the tissue interface suggests that mechanical forces may develop within the tissue; this observation is independent of the biomaterial utilized (7, 120). In contrast, the engineering of more complex tissues consisting of large 3D structures remains a critical challenge. Because the amount of oxygen required for cell survival is limited to a diffusion distance of approximately 150 mm to 200 mm from the supplying blood vessel, the long term survival and function of these 3D-constructed tissues depend on rapid development of new blood vessels, which provide nutrients and oxygen to the cells not only of the margin but also of the center of the tissue grafts. In fact, the growth of a new microvascular system remains one of the major limitations in the successful introduction of tissue engineering products to clinical practice (15).

Numerous degradable polymers such as polycaprolactone, polylacticacid (PLA), polyglycolic acid, chitosan, and their copolymers have been used to fabricate 3D scaffolds (2, 6, 8-11) with rapid prototyping (RP) (119, 121). The scalability of the 3D printing technique enables the manufacturing of large specimens in the meter range as well as small specimens of a few millimeters. RP structures show mechanical properties significantly higher than those of structures fabricated by other well-known techniques such as solvent-casting and particle leaching, thermal-induced phase separation and gas foaming, among others (77).


Natural materials may enhance biological interaction with host tissues. However, the clinical applications of such natural materials are still difficult to realize with current technologies because of their relative mechanical inferiority and instability compared with native cartilage (15). In contrast, synthetic materials, while lacking the intrinsic biocompatibility of natural materials, have important advantages, including some control of the chemical, mechanical, and structural properties of the scaffold and the possibility of satisfying the increasing clinical demand (15).

Due to the variation in material formulation, it is often difficult to come to conclusions on the effectiveness of a particular polymeric material for tissue regeneration. What is clear is that the mechanical conditions are of the upmost importance to ensure engineering success. Hard materials provide compressive and torsional strength, but they are often poor at promoting bone tissue formation. Soft composites made of natural polymers that more effectively promote cell expansion have a very low immunogenic potential, bioactive behavior, and capability of interacting with the host tissue. However, they lack mechanical strength to withstand the forces typically observed in natural bone, which is the prime requisite for bone tissue engineering and tissue formation (122).

Until now, most biomaterials currently available do not completely meet the demands because cellular functions are the result of highly diverse interactions between cells as well as between cells and biomaterials (122). The state of the art in biomaterial design has been continuously evolving over the past few decades, as the goal of biomedical engineering increases in complexity (6). Blending two polymers, natural and synthetic, or hybrid polymeric scaffolds combining natural and synthetic polymers, has gathered growing interest in mimicking the ECM of a natural tissue. These new hybrid structures can present a wide range of the physicochemical properties and processing techniques of synthetic polymers as well as the biocompatibility and biological interactions of natural polymers (123).


The authors wish to thank Prof. G. Magenes, Prof. M.L. Torre at the University of Pavia, Italy, and Prof. D. Vigo at the University of Milan, Italy for their kind suggestions.


Financial Support: This work was supported by the University of Pavia, Pavia, Italy.
Conflict of Interest: The authors state that there are no conflicts of interest.
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  • Dipartimento di Sanità Pubblica, Neuroscienze, Medicina Sperimentale e Forense, University of Pavia, Pavia - Italy

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