Biomaterials for tissue engineering applications pdf

Collagen plays a dominant role in maintaining the biologic and structural integrity of ECM and is highly dynamic, undergoing constant remodeling for proper physiologic functions 1. Hence, the ideal goal of tissue regeneration is to restore both the structural integrity and the vivid remodeling process of native ECM, especially restoring the delicate collagen networks under which normal physiologic regeneration occurs.

Collagen molecules have a triple-helical structure and the presence of 4-hydroxyproline resulting from a posttranslational modification of peptide-bound prolyl residues provides a distinctive marker of these molecules 2. To date, 28 collagen types have been identified; I, II, III, and V are the main types that make up the essential part of collagen in bone, cartilage, tendon, skin, and muscle.

They also exist in fibrillar forms with elaborate 3D arrays in ECM 3 — 5. Bone tissues are mainly constructed from type I with a small quantity of type V collagen forming a framework that anchors nanosized hydroxyapatite HA crystals. Both organic and mineral phases compose a closely interwoven and highly complex but ordered composite which is further organized into layers, or lamellae, that are a few microns thick.

The lamellae are arranged into higher order structures. The organic framework is mainly constructed by type I collagen fibrils in elaborate 3D arrays as concentric weaves 4 , and the anchored bone crystals are 50 nm long, 28 nm wide, and 2 nm thick 6. Bioengineered bone was formed by collagen synthesis, fibril formation, assembling and crosslinking with mineralization by Ca-P crystals to form a continuous network and further remodeling.

Collagen fibers, while providing the framework for HA deposits, are the principal source of tensile strength of bone tissues. The mechanical properties of bone can be adjusted by varying fibril orientation which may result from the degree of collagen crosslinking and the mineral-organic composite remodeling process 7 , 8. There are three main types of cartilage: hyaline articular cartilage, fibrocartilage, and elastic cartilage. This review will only discuss hyaline articular cartilage which is found on the surface of long bones in articulating joints and is extremely important in the lubrication and load distribution in the joint.

Usually, articular cartilage is subdivided into three zones: superficial, middle, and deep zones 9. According to the work by Muir et al. Assuming that different aligning patterns and distribution of collagen fibers account for this difference, cartilage engineering relies on successful engineering of collagen frameworks in different zones for further deposition of other proteins, especially proteoglycans or GAGs.

Tendon transmits the force of muscular contraction to bone 9 , 12 and most tendon types experience tension, such as Achilles or patellar tendon, although a few types also experience a certain degree of compression, such as the supraspinatus tendon Collagen fibrils in tendon vary in diameter from 30 nm to nm, depending on the stage of development Usually a nonuniform or bimodal distribution of fibril diameters can be observed, which allows for tight lateral packing These fibrils are generally oriented longitudinally to support the great tensile stress exerted on the tissue although some may be transversely arranged 9 , Collagen fibrils are essential in tendon for maintaining mechanical properties and for transmitting force during movements 5.

Skin is the body's largest organ composed primarily of epidermis, dermis, and hypodermis layers with a complex nerve and blood supply. Dermis constitutes the bulk of the skin and is composed of collagen with some elastin and GAGs cushioning the body against mechanical injury by conferring elasticity and plasticity to the skin Collagen types III and V have also been detected in the dermis of skin 5 and are primarily responsible for tensile strength.

The dermal layer network also maintains stability of the epidermis The cornea consists of three distinct cellular layers: outer epithelium, central stroma, and inner endothelium. The stroma is the major component of transparent corneal tissue and is mainly comprised of type I and V collagens existing in parallel-organized fibrils with a uniform diameter of around 30 nm 20 that are further arranged in lamellae. Intensive intra- and intermolecular covalent crosslinking takes place universally to stabilize the collagen fibrils against proteolytic degradation and the fibrils confer the desired tensile strength to the stroma.

The 3D architecture of the thin collagen fibrils may contribute to cornea transparency The complex structure of fibrillar type I collagen presents different morphologies in different tissues performing different functions. When associated with HA crystals in bone, it provides rigid and shock-resistant tissues with high Young modulus. It can behave like an elastomer with low rigidity and high deformation to rupture in tendon.

In cornea, it shows optical properties such as transparency. Capturing the intrinsic building blocks and more complex, higher order supramolecular structures of collagen has been the current focus of the biomaterial society. In view of the variety of roles played by collagen in different tissues, research has focused on developing novel biomaterials to mimic the intricate fibrillar architecture which would be able to function as cell scaffolding replacing native collagen-based ECM.

One promising candidate is peptide-amphiphile PA nanofibers which are produced by molecular self-assembly as developed by Stupp and coworkers 22 , The self-assembling unit is a PA, a hydrophilic residue peptide with an N-terminal alkyl tail of 16 carbon atoms.

PAs are synthesized by standard solid phase chemistry that ends with the alkylation of the NH 2 terminus of the peptide and they further self-associate to form cylindrical micelles or nanofibers because of the amphiphile's overall conical shape. The alkyl tails are packed in the centers of the fibers, while the peptide segments are exposed to the aqueous environment so the chemistry of the peptide region is repetitively displayed on their surface.

The incorporation of four consecutive cysteine amino acids helps to stabilize the fibrous supramolecular structure and the introduction of a phosphoserine residue allows the fiber to display a highly phosphorylated surface to facilitate the formation of calcium phosphate minerals.

The sequence of Arg-Gly-Asp in the peptide is expected to promote the adhesion and growth of cells on fiber surfaces. Mineralization of the fibers was demonstrated and showed that plate-shaped polycrystalline minerals deposited on the PA fibers were consistent with HA found in native bone tissue. Thus, these mineralized nanofibers resemble the lowest level of hierarchical organization of bone 22 , Furthermore, the same group extended their work in such supramolecular structures by synthesizing self-assembling oligomeric-amphiphiles that allow incorporation of specific biomolecular signals, which is expected to direct and regulate the cellular behaviors Zhang developed a class of nanofibrillar gels crosslinked by self-assembly of self-complementary amphiphilic peptides in physiologic medium These simple, defined, and tailor-made self-assembling peptides have provided de novo -designed scaffolds for 3D cell culture and they biomechanically organize cells in a 3D fashion.

Continued efforts to control nanofibrillar morphologies include the elucidation of rational design principles to better control the production of kinked, wavy, or branched fibers The development of artificial collagen-like materials is another important area of research. By using native chemical ligation techniques that synthetic peptide units spontaneously polymerize in aqueous solutions, Paramonov et al.

The triple-helical polymers can also form a nanofiber-like structure with a length in micrometers. Since the polymerization proceeds under mild conditions, it is possible to incorporate amino acid residues with various side-chain functional groups.

The presence of Cys residues on the polymers also enables further crosslinking and modification by functional moieties. Donor-site availability, deformation and infection also need to be considered. Figure 1 Optimum model of mechanical properties of a scaffold as new The limited donor sites from which cartilage may be tissue develops acquired poses a significant problem for the treatment of significant cartilage defects such as those occuring in load- bearing joints and facial structures.

CC or stem cells can be expanded in vitro into a and waste products. Furthermore, the scaffold should larger population of cells and its surrounding matrix [15]. The restricted replicative capacity of CC is due to cartilaginous tissue.

Initially culture [18]. Homicz and co-workers found that primary the scaffold withstands most of the stress. Stem cells possess a multipotent, regenerating cells are gradually loaded with physiological highly replicative capacity, so fewer cells are required at stress, further simulating tissue regeneration.

Eventually the harvest. They can be acquired from a selection of sites such scaffold completely degrades and the regenerated tissue as blood, adipose tissue, bone marrow and even muscle bears the stress. The integral aim of tissue engineering is to create Early studies on chondrocyte proliferation used algi- new tissue from specific cells by providing an artificial nate, agarose and micromass pellet culture.

Over the past environment which mimics that of the native tissue [21]. This few decades, a multitude of natural and synthetic polymers is achieved with the aid of a scaffold and the incorporation of have been used as a delivery vehicle for these cells Table 1 exogenous factors required for cell development. Advances in biomaterial engineering have afforded the expansion of fabrication techniques and the development of novel porous Natural materials biodegradable scaffolds. Natural materials used to produce a bioactive scaffold The present Review concentrates on the scaffold ele- for tissue-engineered cartilage include agarose, alginate, ment of tissue-engineering cartilage and considers the cur- hyaluronic acid, gelatin, fibrin glue, collagen derivatives and rent advancements in polymer development and fabrication acellular cartilage matrix.

Potential problems, however, lie in their inferior mech- anical strength, disease transfer and antigenicity. They are also prone to rapid and variable host-related degradation. Materials used in development of cartilage scaffold Alginate and agarose Alginate and agarose are polysacchar- ide hydrogels derived from marine plants.

They are therefore A tissue-engineering scaffold should allow cell adherence and not easily degraded by mammalian metabolism. CC cannot migration, with subsequent differentiation and proliferation. Such scaffolds have a useful role in tissue formation, whereas the mechanical properties should be engineering, mainly for in vitro investigation, because they similar to that of native tissue. A porous scaffold will provide are relatively inert.

As discussed later, hybrid scaffolds of the 3D environment required for phenotypic stability of alginate with synthetic polymers offer more promise for use CC, and will also allow for mass transfer of nutrients in clinical applications [22]. Raghunath and others Alginate is derived from brown marine algae. The cells may be released from the alginate by mild [40]. Collagen can be fabricated as a gel, sponge or foam chelation.

This scaffold system has been well documented and is subject to enzymatic degradation. It has been shown for chondrocyte and chondro-induced stem-cell growth to maintain the chondrocytic phenotype in the gel form for in vitro [23—25]. Researchers at the Department of Ortho- up to 6 weeks [41]. Collagen I an appropriate environment for maintaining phenotype and sponges seeded with foetal bovine epiphyseal CC showed expression of cartilaginous markers appropriate to the comparable chondrogenesis in the native and cross-linked hyaline form.

In recent years, alginate has been used to form, with notably lower matrix accumulation of GAGs than propogate chondrocyte growth and cartilaginous matrix that which is found in normal cartilage [44]. When combined production prior to its recovery for seeding on to another with chondroitin sulfate, collagen II appears to promote scaffold [28,29]. Alginate gel sheets have recently been more diffuse chondrogenesis and matrix production than evaluated as a potential scaffold for cell proliferation [30].

Current research into the fabrication of a porous alginate gel may be useful for neocartilage formation, but it is unlikely Gelatin Gelatin is a porous compound derived from the that this substrate will ever have a role in vivo because of its hydrolysis of collagen. It has known biocompatibility in mechanical fragility.

Its use Agarose is a linear polysacharride consisting of basic as a scaffold for cartilage engineering has been relatively repeating units of agarobiose. It is derived from Asian unexplored [23,45—47]. As with come by the development of a low-melting-point agarose. This may interfere with integration of the im- optimize nutrient and waste exchange, as the stiffer gels plant in vivo.

A recent study comparing CC seeded biocompatibility. This result Good biocompatibility was demonstrated by minimal demonstrates the importance of determining the ideal scaf- inflammation and giant-cell infiltration in the explant from fold to support the production of constituent proportions which the gelatin sponge had completely disintegrated.

At 14 days on cartilage formation [34—36]. Rahfoth and colleagues post-implantation, there was considerable scatter of the in Erfurt, Erlangen and Jena in Germany [37] performed labelled MPCs into the deeper regions of the defect, but they allograft transplants of CC in agarose gel into osteochondral remained in contact with the gelatin. There also appeared defects in the knees of rabbits. They reported no graft- to be marked recruitment of surrounding host cells into versus-host rejections or foreign-body immune responses the defect.

A study using sternal chick-embryo indeed integrated into the host tissue. Although these this biomaterial to have therapeutic potential. It would results are encouraging, the choice of cell source is ques- also be valuable to consider the differences in mechanical tionable, as the CC of embryonic avian species are far loading and wear found in a human joint, as these may differ removed from human CC. HYAFF is more demonstrable considerably from the situation in goats.

Biocompatibility, bioresorbabilty and its ability to scaffold because it can be made from autologous blood integrate with normal hyaline articular cartilage are all and already has a good record of biocompatibility as a promising characteristics reported from in vivo studies on wound adhesive.

It can be produced in an injectable form athymic mice and rabbits [55—57,59]. CC and stem cells possess Chitin Chitin and its partially deacetylated derivative integrins that can bind directly to this endogenous protein. Chitin is an infinitely available polymer interactions on cell behaviour, but it has been postulated found in the exoskeletons of arthropods.

This polysac- that fibrin may provide a chemotactic and mitogenic stimuli charide-based analogue of GAG is degraded in vivo by for MSCs [48]. It forms water-insoluble deficiency disease mice resulted in the production of hyaline ionic complexes to which a large repertoire of moieties neocartilage [49,50]. An excellent illustration of the biocom- may be added.

Sechriest and colleagues found that chitosan patibility of fibrin was demonstrated by a recent auto- combined with chondroitin sulfate A maintained the chon- logous study where auricular cartilage of immunocompetent drocytic phenotype and supported proteoglycan production rabbits were suspended in fibrin and implanted for [60]. Chitin fibre meshes have been used to transport rabbit 3 months in the dorsum of its donor [51].

The angulation and length discrepancies and basic fibroblastic growth factor in the scaffold resulted in were significantly corrected at 16 weeks with chitin scaffolds minimal cartilage production. Further studies into growth- alone, and normalized with progenitor-cell-seeded chitin factor concentrations may be required to fully explain these scaffolds. The main problem with fibrin constructs is shrinkage Synthetic materials in vivo.

The aim of this special issue is to bring forth the synergy between material design strategies and biological evaluations through new and significant contributions from active researchers in the field. Topics of interest include, but are not limited to:. Article of the Year Award: Outstanding research contributions of , as selected by our Chief Editors. Read the winning articles. Journal overview. Special Issues. Polymeric Biomaterials for Tissue Engineering Applications.

Publishing date. The presence of multivalency and molecular mobility in PRX supramolecular assemblies is also one of the key mechanisms which leads to improved efficiency of cell-ligand interactions [ 23 , 24 , 25 ]. The choice of determining the suitable axle polymer and the type of CD usually depends on the cavity diameter of the CD. PRXs can be processed into scaffolds with controllable pore sizes. This helps to widen the utility for tissue engineering applications. PEG hydrogel scaffolds were produced by crosslinking hydrolyzable PRXs and also with cholesterol functionalized PRXs which were explored for cartilage regenerative applications.

Hydrolyzable PRX-scaffolds containing hydroxyapatite particles were also synthesized and were suggested to exhibit promising outcomes for bone tissue engineering. PRXs can be utilized as crosslinkers for the poly N -isopropylacrylamide PIPAAm polymeric networks, thereby yielding a thermoresponsive and extremely stretchable hydrogel, which could be used for muscle tissue engineering [ 22 , 30 ].

Similar high elastic hydrogels can be prepared by crosslinking acrylate monomers with slide-ring PRXs by exploiting the pulley effect [ 31 ]. Self-healing hydrogels can also be synthesized by creating a reversible bond between the ring molecules of PRXs and poly acrylamide [ 32 ]. Various stimuli-responsive PRX hydrogels, such as pH-sensitive, thermo-sensitive, photo-sensitive, etc.

These are usually synthesized by adjusting the threading CDs and the length of the polymeric axle, making chemical substitutions in the PRXs or by introducing stimuli-sensitive groups, which either breaks down or changes the conformation upon the necessary stimuli [ 33 , 34 ]. For example, self-assembled PRX particles with disulfide bonds that could degrade and release the encapsulated drugs in the reducing environment have been developed [ 35 ].

With such numerous options to modify, improve, and incorporate unique features, PRXs can be considered to be one of the very versatile biomaterials allowing us to cater a wide array of clinical requirements. Controlling the commitment of stem cells has a lot of implications in the field of tissue engineering and regenerative medicine.

Growth factors, chemical supplements, siRNAs, and a combination of these have been studied extensively for the purpose of directing stem cell commitment [ 41 , 42 , 43 ]. It has been well established that the fate of stem cell differentiation also depends on the mechanical cues from the microenvironment [ 44 ].

Taking insights from this fact, PRXs with different molecular mobilities have been synthesized and their potential to alter the fate of stem cells were explored. Seo et al. It was found that molecular mobility influences the cellular response through focal adhesion kinases, actin organization, and RhoA-ROCK mechanosignaling pathway.

Thereby, the possibility of using thin coatings of PRXs with different molecular mobility on the materials-surface to alter the stem cell commitment was established. Culturing iPSCs on PRXs with high mobility lead to the higher expression of Rac1 and N-cadherin expression indicating the strong cell-cell interactions when compared to cells cultured on gelatin-coated surfaces.

The number of beating colonies indicating the successful cardiomyogenic commitment was higher on the high mobile PRX surface. Recently, it was reported that molecular mobility of PRXs could alter the stemness and differentiation capability of bone marrow-derived MSCs even during a short culture span [ 47 ].

The MSCs cultured over high mobile S-PRX surfaces exhibited poor actin organization, retention of a key mechanosignaling element of YAP in the cytoplasm, and higher levels of stemness marker gene expression such as Nanog and Oct4. Further, when these MSCs were collected and replated onto regular tissue culture polystyrene plates, only the cells grown on high mobile S-PRX surfaces with bFGF were able to differentiate into osteoblasts.

Biomaterials for bone tissue engineering and enhancing osteogenic differentiation are of great importance as bone is the most widely transplanted tissue next to blood transfusions [ 48 ]. Although autografts remain as the gold standard for bone tissue repair, bone graft substitutes and alternative solutions to enhance the regenerative capacity of bone tissues are persistently being sought upon, owing to the practical limitations of autografts [ 49 ].

The molecular mobility and the ability of PRXs to be modified with various functional groups and complex with growth factors have been explored for the purpose of enhancing bone regeneration. Sulfonation of CDs in the S-PRX helps to enhance the activity of BMP2, mimicking the function of heparin, proving to be advantageous than heparin to be used in a clinical scenario since S-PRXs did not exhibit the anticoagulant activity.

Further, these complexes were shown to significantly improve the bone regeneration in vivo, when it was encapsulated within collagen sponges and implanted into the mouse calvarial defects [ 51 ].

This possibly allows PRX to be stored for a longer duration and immobilization of BMP2 can be done a few hours before the clinical application. Cartilage regenerative therapies mainly focus on finding suitable biomaterial scaffolds to support the growth of chondrocytes and aid the secretion of extracellular matrix proteins in a spatiotemporally controlled pattern for matching the native healthy tissue as close as possible [ 53 ]. Lee et al.

The scaffold showed promising results for the proliferation of the chondrocytes in vitro. Further, the microporous hydrogel scaffold was able to trap a higher number of chondrocytes due to the primary amine functionalization in the PRXs [ 55 ]. The erosion rate of the PRX-based hydrogel scaffold was also controllable, thereby providing ample time for the chondrocytes to produce of glycosaminoglycan GAG , indicative of promising use in cartilage regeneration.

A similar hydrogel scaffold was synthesized by introducing cholesterol moieties in the PRXs, which has the advantage to mimic the lipid layers in the cell membrane, thereby improving the biocompatibility [ 56 ].

Cellular adhesion is one of the most important requirements for biomaterial interfaces in the tissue engineering approach. It has been found that the molecular mobility of PRXs could be used to alter the cellular adhesion patterns. This could be because the low mobile PRX could mimic the mechanics of the native matrices of adherent cells leading to higher cell-biomaterial interactions.

It could also be reasoned that due to excessive sliding of CDs in the high mobile PRXs, the cells may not achieve enough tractions for effective spreading seen on soft hydrogels, thereby inhibiting the other functions such as cell migration or proliferation [ 57 ]. Apart from utilizing the low molecular mobility of PRXs to enhance cellular adhesion, RGD peptides have also been conjugated to the dynamic CDs in the PRXs to synthesize cell adhesive hydrogel scaffold [ 58 ].

Formation of microvascular networks in implanted biomaterials is one of the key events for a successful tissue regeneration, as these microvessels are necessary for the efficient transport of cells, nutrients, and removal of the wastes [ 59 ]. It was observed that the molecular mobility of PRXs could be used to modulate the formation of microvascular networks [ 60 ].

The effect of molecular mobility of S-PRXs on hepatocyte functions has also been explored by using HepG2 cells [ 61 ]. Consistently, the albumin secretion, one of the key hepatic functions, was significantly higher in the cells cultured on high mobile S-PRX with immobilized HB-EGF, indicating that these surfaces could be useful in hepatocyte cultures and liver tissue engineering applications.

The molecular mobility of polyrotaxane surfaces also affected collagen fibrillogenesis. Although the extent of molecular mobility was independent of the absorbing amount of collagen, higher mobility of polyrotaxane surfaces preferentially induced rearrangement of adsorbed collagen and caused collagen fibrillogenesis. When PRX-coated substrates were implanted subcutaneously in rats, it was observed that the recruitment of macrophages at the implant site was suppressed by polyrotaxane surfaces with higher mobility.

These results suggest that the molecular mobility of polyrotaxane helps suppress inflammation and promote regeneration in vivo [ 62 ]. As briefly reviewed above, we could see that PRXs are a novel class of supramolecular assemblies which exhibit mechanical movements at a molecular scale.

It has also been shown that these mechanical signals from the molecular mobility of PRXs could be harnessed and utilized for modulating various mechanosignaling pathways and thereby driving specific gene expressions Table 1. One of the peculiar advantages of PRXs is that the effects of molecular mobility holds true at various scales, such as hydrogels, scaffolds, and even in thin coatings.

This allows us to utilize PRXs in various forms for different clinical scenarios. Furthermore, with the other advantages of being able to synthesize from biocompatible and biodegradable precursors and the ability to add various chemical groups, peptides, growth factors, and vitamins to impart cell-specific functions in the field of tissue engineering in the future, it should also be noted that different types of cell or tissues could exhibit different responses even to the same kind of PRXs.

Therefore, it is necessary to design and choose the appropriate type of PRXs for a specific application. However, it could also be foreseen that researchers might develop multilayered or multiscale constructs in which different segments of the construct could exhibit different molecular mobility, thereby closely mimicking the complex layers and mechanics of the real tissues. However, we could see that most of the studies have shown the intended effects of PRXs only in the in vitro stage, except for a handful of studies in vivo.

Exploring the in vivo effects could possibly bring the interesting potentials of PRXs to the clinical benches. Thus, PRX-based biomaterials hold promising directions in the field of mechanobiology mediated tissue engineering. Stratakis E.