Scaffolds in tissue engineering pdf

Log in with Facebook Log in with Google. Remember me on this computer. Enter the email address you signed up with and we'll email you a reset link. Need an account? Click here to sign up. Download Free PDF. Scaffold-based tissue engineering: rationale for computer-aided design and solid free-form fabrication systems Trends in Biotechnology, Engineering Computer. A short summary of this paper. Scaffold-based tissue engineering: rationale for computer-aided design and solid free-form fabrication systems.

Hutmacher1, Michael Sittinger2 and Makarand V. Furthermore, scaffold architec- the development of 3D scaffolds that guide cells to form ture should enhance initial cell attachment and subse- functional tissue. Recently, mouldless manufacturing quent migration into the matrix; it must also enhance techniques, known as solid free-form fabrication SFF , the mass transfer of metabolites and provide sufficient or rapid prototyping, have been successfully used to space for remodelling of the organized tissue matrix and fabricate complex scaffolds.

Similarly, to achieve simul- development of a vasculature. To achieve this goal, the taneous addition of cells during the scaffold fabrication, scaffold degradation profile should be designed so that it novel robotic assembly and automated 3D cell encapsu- supports the construct until neotissue cells plus organized lation techniques are being developed. As a result of extracellular matrix without vascularization is formed these technologies, tissue-engineered constructs can [1]. Factors structure.

Here, we review the application, advance- affecting the rate of remodelling include the type of tissue ment and future directions of SFF techniques in the and the anatomy and physiology of the host tissue [2]. The design and creation of scaffolds for use in clinically external size and shape of the construct must also be driven tissue engineering.

It is also imperative Currently, scaffold-based tissue engineering strategies are that scaffolds are manufactured in a reproducible, con- expanding to encompass cells, bioactive molecules and trolled, and cost-effective fashion with the flexibility to structural matrices. To case scenario — regeneration of damaged or diseased date, two methods of incorporating cells into scaffolds are tissues.

In terms of organ Factors governing scaffold design are complex and printing, the second process is of considerable interest. Moreover, because scaffolds are often composite structures, other confound- Basic considerations for scaffold fabrication by SFF ing variables include the composition of biological com- Advanced mouldless manufacturing techniques, com- ponents and variation of these and other factors with time. Scaffold design should shaped scaffolds.

Unlike conventional machining, which therefore begin with at least a set of minimum biochemical involves constant removal of materials, SFF builds parts and physical requirements [1]. The scaffold must provide by selectively adding materials, layer by layer, as specified sufficient mechanical strength and stiffness to substitute by a computer program.

Each layer represents the shape of initially for wound contraction forces, and later for the the cross-section of the model at a specific level.

Hutmacher biedwh nus. In www. All rights reserved. Over the past two decades,. Stereolithography SLA is cations. Very recently, biomaterial scientists used these often considered the pioneering RP technique with the technologies to fabricate scaffolds for tissue engineering. This technique is based on the matrix architecture size, shape, interconnectivity, branch- use of a UV laser that is vector scanned over the top of a ing, geometry and orientation yielding biomimetic struc- bath of a photopolymerisable liquid polymer material.

As tures varying in design and material composition, thereby polymerisation is initiated, the laser beam creates a first enhancing control over mechanical properties, biological solid plastic layer, at, and just below the surface of the effects and degradation kinetics of the scaffolds. RP bath. This laser polymerisation process is repeated to techniques can be easily automated and integrated with generate subsequent layers by tracing the laser beam imaging techniques to produce scaffolds that are custom- along the design boundaries and filling in the 2D cross- ised in size and shape allowing tissue-engineered grafts to section of the model, layer-by-layer.

Once the model is be tailored for specific applications or even for individual complete, the platform is raised out of the vat and the patients Figure 1.

The model is cured in a UV oven A recent milestone in scaffold fabrication by SFF was and finished by smoothing the surface irregularities. By allowing inclusion of bioactive components, such elevator layer resolution up to 1. In the following sections we discuss the state limited to the creation of anatomical models for surgical of the art in SFF techniques applied to scaffold fabrication. Because of curing and We have classified the SFF systems based on their shrinkage after post-processing a shortfall of the SLA processing technology.

Future directions in the design model is compromised resolution [13]. Also, due to absorp- and manufacturing of novel and patient-specific matrices tion and scattering of the laser beam, a pronounced a b c d e f g defect TRENDS in Biotechnology Figure 1. Tissue engineering of patient-specific bone grafts. This innovative treatment protocol uses medical imaging, computational modelling and bioresorbable scaffold fabricated with rapid prototyping RP technique.

CT scan data of the patients bone defect a are used to generate a computer-based 3D model b. RP technology produces excellent templates for the treatment of intricate bone defects a and f. Custom-made scaffold and cell constructs g, see arrows exactly follow the complex shaped 3D contour of the skull. Generally, the manufactured part is parts using suspensions of alumina, silicon nitride and weak at the time of removal and needs post-processing for silica particles in a UV photocurable monomer by SLA [22].

It is noteworthy that there is a Using a similar technique, a suspension of hydroxyapatite limited choice of photopolymerisable biomaterials that HA in a photocurable monomer was formulated to have the required biodegradability and biocompatability, produce scaffolds for orbital floor prosthesis [23].

It was mechanical stability and other prerequisite properties for concluded that for bone graft applications HA scaffolds scaffold applications. Examples of photopolymerisable provide a superior cosmetic appearance compared with macromers include derivatives of polyethylene glycol conventional techniques [23].

Porter et al. Polypropylene fumarate, implants using SLA [24]. Sintering CPP at C for 1 h anhydride and polyethylene oxide PEO precursor sys- produced a crystalline material average porosity of Tan et al.

Cooke et al. Likewise, Matsuda and Mizutani have developed a biodegradable, System based on print technology — 3D printing. An advantage of 3DP is that it can be performed [15]. Based on such encouraging results micro stereolitho- in an ambient environment. The 3D printer constructs the graphy mSL in particular offers great potential for the 3D model by first spreading a layer of fresh powder over a production of 3D polymeric structures with micrometer building platform.

Recently, mSL was also tested for fabricating the binder solution onto the powder bed. After the 2D layer high aspect ratio and complex 3D microstructures [16,17]. The printing cycle continues and the layers merge together Selective laser sintering. The selective laser sintering when fresh binder is deposited until the whole part is SLS technique uses a CO2 laser beam to sinter thin layers completed.

After the binder has dried in the powder bed, of powdered polymeric materials, forming solid 3D objects. However, if the part is designed to be porous, one scanned over the powder surface following the cross- drawback of a powder-supported and powder-filled struc- sectional profiles carried by the slice data.

The interaction of the laser beam with the powder raises the powder ture is the difficulty in removing internal unbound powder. Subsequent layers are position controller that defines print head movement. In built directly on top of previously sintered layers with new addition, the particle size of the powder governs the layers of powder being deposited by a roller [18]. As a result, various efforts such as agitation, perfusion and enlarged pore size are needed to enhance cell seeding efficiency and these methods are not free of problems, to list a few, low cell viability and high cost.

Acellular ECM processed from allogenic or xenogenic tissues are the most nature-simulating scaffolds, which have been used in tissue engineering of many tissues including heart valves [ 62 ], vessels [ 15 ], nerves [ 44 ], tendon and ligament [ 56 ]. This scaffolding approach removes the allogenic or xenogenic cellular antigens from the tissues as they are the sources for immunogenicity upon implantation but preserves the ECM components, which are conserved among species and therefore well tolerated immunologically.

Specialized decellularization techniques are developed to remove cellular components and this is usually achieved by a combination of physical, chemical and enzymatic methods [ 9 , 40 ]. The cytoplasmic and nuclear cellular components will then be solubilized and removed by a handful choice of detergents. The process parameters need to be optimized to aid complete decellularization with minimal disturbance on biochemical composition and mechanical properties of the ECM. Decellularized ECM can be used for homologous functions when the decellularized ECM is used to replace an analogous structural tissue that has been damaged.

An example is to decellularize vessels as allogenic vascular grafts [ 15 , ]. Decellularized ECM can also be used for non-homologous functions when it is used for a purpose different from which it fulfills in its native state, or in a location of the body where such structural function does not normally occur.

Examples are to use small intestinal submucosa SIS for vascular graft, tendon, dura mater, skin and other tissues [ 7 , 8 , 38 , 49 , ] and to use amnion membrane for peripheral nerve regeneration [ 77 ].

The major advantage of this scaffolding approach is the most close-to-nature mechanical and biological properties of the decellularized ECM. Moreover, apart from the excellent biocompatibility of the natural ECM, growth factors preserved in the decellularized matrix may further facilitate cell growth and remodeling [ 9 ]. Nevertheless, cell seeding in decellularized ECM may also lead to inhomogeneous distribution while incomplete removal of cellular components may elicit immune reactions upon implantation [ ].

Cell sheet engineering represents an approach where cells secrete their own ECM upon confluence and are harvested without the use of enzymatic methods. This is achieved by culturing cells on thermo-responsive polymer, such as poly N -isopropylacrylamide coated culture dish until confluence.

The confluent cell sheet is then detached by thermally regulating the hydrophobicity of the polymer coatings without enzymatic treatment.

Such approach can be repeated to laminate multiple single cell layers to form thicker matrix. This technology is pioneered by Japanese group [ 84 , 85 , ] and has been systemically applied to a number of applications such as cornea [ 82 ] in clinical trials and myocardium [ ] in preclinical trials.

The same approach has been modified and improved to produce patterned substrates for thicker tissue fabrication and for injectable applications [ 57 ]. Thermosensitive chitosan has also been investigated to create an aligned cell sheet [ 32 ]. Cell sheet engineering approach is excellent for epithelium, endothelium and cell-dense tissues [ , ] as the formation of cell sheets require cells to grow into confluence at high density such that cells can form tight junctions with each other and secrete ECM of their own.

High cell density and close association is a characteristic of epithelium and endothelium. Therefore, corneal epithelium and endothelium, vessel endothelium and tracheal epithelium are good candidates of this approach. Another advantage of the cell sheet engineering approach is that the laminated layers resulted in rapid neovascularization unlike transplantation of thick constructs with cell-seeded scaffolds. A recent report reviewed other advantages of the cell sheet engineering approach such as easy harvesting procedure, possibility of sutureless transplantation, and even distribution of cells in the sheet without mass transfer problem [ 31 ].

This is less clinically feasible as multiple surgeries in patients are unlikely and therefore preformed patches are needed. The other disadvantage or perhaps limitation of cell sheet engineering approach is that constructing ECM rich tissues and hypocellular tissue such as bones, cartilage and intervertebral disc are unlikely as the amount of ECM secreted upon cell confluence is severely limited. In another word, tissues with rich ECM for load bearing purposes are unlikely to be fabricated by a cell sheet engineering approach [ ].

Encapsulation is a process entrapping living cells within the confines of a semi-permeable membrane or within a homogenous solid mass [ 66 , 86 , 87 , ]. The biomaterials used for encapsulation are usually hydrogels, which are formed by covalent or ionic crosslinking of water-soluble polymers. Many types of biomaterials including natural and synthetic hydrogels can be used for encapsulation provided that the conditions inducing the hydrogel formation or the polymerization are compatible with living cells.

Encapsulation has been developed over several decades and the predominating use is for immunoisolation during allogenic or xenogenic cell transplantation [ 87 , ].

Naturally occurring polysaccharides derived from algae, sodium alginate is the most commonly used material while other natural materials such as agarose [ 10 ] and chitosan [ ] and synthetic materials such as poly ethylene glycol PEG [ 83 ] and polyvinylalcohol PVA [ 58 ] are also used.

The most well-known application is xenogenic pancreatic cell transplantation for diabetes [ 41 , 66 ] while applications for other disorders such as CNS insufficiency [ 86 ] and liver failure [ 24 ] are also reported. For immunoisolation to work, biomaterials encapsulating the cells need to be crosslinked or processed to become impenetrable to cells, impermeable to large molecules such as antibodies and cellular antigens but permeable to nutrients such as oxygen and glucose, metabolites such as carbon dioxide and lactic acid, and secreted therapeutic biomolecules from the encapsulated cells such as insulin from pancreatic beta cells.

In case of a semi-permeable membrane, the encapsulated cells should have the ability to maintain viability and functionality in aggregates even though there is only solid anchorage support limited to the luminal surface. Pancreatic cells, hepatocytes and haematopoietic cells are of this type. In case of a homogenous solid mass, where the entrapped cells are closely interacting with the biomaterials, the biomaterials should have good biocompatibility enabling cellular attachment and growth.

Nevertheless, the commonly used encapsulating materials such as alginate and agarose have limited ability to support cell attachment growth and differentiation, resulting in low cell viability and growth [ 42 , 83 , ]. In many cases, a biomaterial with better biocompatibility such as collagen must be supplemented for improvement in cell viability [ 10 , 42 ].

Collagen is a natural biocompatible and biodegradable material [ ] and can be reconstituted into fibrous structures simulating the native ECM in tissues. Recently, a microencapsulation system immobilizing living cells within reconstituted collagen fiber meshwork has been established [ 19 ] and the collagen meshwork is able to provide a bio-mimetic scaffold supporting cell growth, migration [ 21 ], therapeutic protein secretion [ ] and stem cell differentiation [ 52 ].

Moreover, chemical approaches have been used to design self-assembled peptides [ 39 , ] and these biomimetic peptides can also be used to entrap cells [ ]. One important feature of encapsulation is that the biomaterials used are able to self-assemble from liquid monomers to solid polymer meshwork upon initiation, which is usually pH, temperature, ionic strength and light controlled.

To list a few examples, alginate solidifies when its monomer solution is exposed to divalent ion solutions such as calcium chloride, where the calcium ion crosslinks the alginate; collagen monomers polymerize when they are switched from an acidic pH and a low temperature to a neutral pH and a body temperature; ethylene glycol modified with acrylic moiety starts to polymerize when it is exposed to UV light in the presence of a photo-initiator [ 81 ].

This unique feature combines the scaffold fabrication and the cell seeding into one-step procedure as cells can be mixed with the liquid biomaterials before initiation of polymerization. The advantages of this approach include simple one-step procedure, homogenous cell distribution in the hydrogel and excellent cell viability.

This represents a minimally invasive approach of tissue engineering and is advantageous when the defect is irregularly shaped. Because of the injectable feature of the scaffold, this approach is sometimes referred as injectable scaffolding or in situ tissue engineering.

Nevertheless, hydrogel materials in this approach used usually have poor mechanical properties. As a result, this scaffolding approach is seldom used for tissues with load bearing functions. Selecting the scaffolding approach for tissue engineering is tissue- and application-specific.

Intervertebral disc is chosen as an illustrating example in this review. Extensive efforts have been made to search for biological therapeutics for disc degeneration of different severity [ 48 , ]. Readers are directed to excellent reviews on structural functional relationship and pathophysiological aspects of IVD [ 1 , 12 , ] and excellent reviews on potential biological therapies including growth factor, cell and tissue engineering approaches [ 3 , 6 , 37 , 43 , 88 ].

In this review, existing scaffolding approaches for disc regeneration and their insufficiencies, the unique considerations of intervertebral disc and the future directions of scaffolding in IVD tissue engineering will be reviewed Table 3.

Existing scaffolding approaches, insufficiencies and future directions for IVD tissue engineering. In early disc degeneration, disc cells in particular the nucleus pulposus NP cells become less capable to synthesize the proper ECM.

The gelatinous NP becomes more fibrous with reduced water content. The primary to be repaired should be the capability of the NP cells to secrete the right matrix components in particular proteoglycans, which are responsible for the water absorbing function of the nucleus matrix.

At this early stage, matrix loss is still limited or partial and there is no need to replace the bulk matrix. As a result, minimally invasive treatments such as injecting growth factors to stimulate the NP cells to synthesize proteoglycans [ 74 ] or injecting cells able to synthesize appropriate ECM such as bone marrow mesenchymal stem cells [ 48 , 67 , ] are the most commonly proposed.

Nevertheless, one important inadequacy of this approach is that the high disc pressure and the aqueous nature of the cell suspension usually result in depletion of the local availability of cells [ 5 , 53 ]. As a result, use of injectable carriers is necessary to effectively deliver cells into the degenerative NP space. Therefore, self-assembled hydrogels able to suspend and deliver cells in solution via injection but which can solidify after injection presents an appropriate scaffolding approach.

An example is to deliver mesenchymal stem cells in hyaluronan gel [ 30 ]. This has been observed in our own group using other hydrogels unpublished work. Therefore, better injectable approaches such as carriers with higher viscosity and stiffness, and in situ welding techniques such as laser welding, photochemical welding and use of biological glues at the injection site should be explored in order to prevent leakage and extrusion and thus improve the effectiveness of the injectable therapy.

As the disc degeneration progresses, more ECM and structural changes such as proteoglycan and collagen degradation involving NP and sometimes annulus fibrosus AF ensue, leading to disc height reduction.

At this stage, replacement of cells with capacity to synthesize ECM and compensation of the substantial loss of matrix components are appropriate therapeutic approaches. This can be achieved by implanting cell-seeded pre-made scaffolds approach 1 or decellularized ECM approach 2 or injecting cells encapsulated in self-assembled hydrogels approach 4 into the intradiscal space.

In reviewing the existing scaffolding approaches for IVD tissue engineering, approach 1 using pre-made scaffold is most common. A wide range of biomaterials dominated by natural biomaterials has been used for nucleus replacement.

Reviews for nucleus replacements have been reported elsewhere [ 29 , ]. In most if not all cases, freeze-drying has been employed as the fabrication method to create porous structures. In most nucleus and annulus replacement strategies, survival and growth of seeded cells, and enhanced synthesis of collagen II and proteoglycans by these cells are reported [ 2 , 23 , 70 , 76 , 94 , 97 , , ].

Nevertheless, almost all studies using the pre-made scaffold approach are in vitro while only a few are in vivo [ 76 ].

Subcutaneous implantation was used and the implant was never exposed to physiological loading. Recently, there is one ex vivo study reporting collagen nucleus replacement in bovine discs under mechanical loading [ ]. Although restoration of disc height is possible, extrusion of the whole implant after a few loading cycles have been reported.

As a result, maintaining annulus integrity by sealing or welding methods after nucleotomy or injection or insertion of nucleus replacements is crucial. As for annulus replacement, the intrinsic mechanical properties of the annulus replacements have to be comparable to that of the native disc in order to be able to resist physiological loading. This is extremely challenging because the native annulus fibrosis serves a highly demanding and complex mechanical function.

In existing literature, only a few annulus replacement strategies mentioned enhanced mechanical properties using composites materials [ ] and electrospun fibers with alignment [ 80 ]. As a result, further efforts in enhancing the mechanical integrity of scaffolds for annulus replacement should be encouraged.

The efforts using the second scaffolding approach, decellularized ECM for nucleus replacement is minimal. Only one report uses porcine SIS as the scaffold for human disc cells [ 68 ]. Cell survival and enhanced matrix deposition have been reported in vitro [ 68 ]. In a pilot animal study using SIS in nucleotomized baboons [ 69 ], MRI suggested a higher water content in the treatment groups compared to the nucleotomy group and some tissue remodeling in the disc space in the group with bone marrow-soaked SIS.

Moreover, SIS has good tissue biocompatibility in the disc space and seemed to have biodegraded over the six-month period. Nevertheless, a sizable study with more animals is needed before a conclusive statement on the efficacy of using SIS as nucleus replacement can be made. There is no study in decellularizing allogenic or xenogenic nucleus or annulus for replacement at all.

Nevertheless, there is one study evaluating the mechanical properties of decellularized temporomandibular joint disc [ 73 ]. Although the immune privileged status of IVD, which is avascular, allows the application of allografts without decellularization, research efforts in decellularizing nucleus and annulus grafts from xenogenic sources for scaffolding should be encouraged because allografts are not always available.

Preservation of proteoglycans will, however, be challenging. Mesenchymal stem cells and disc cells have been encapsulated in thermosensitive hydroxybutyl chitosan gel and cell proliferation and matrix production has been demonstrated [ 33 ].

Atellocollagen type II, hyaluronan and aggrecan gels support NP viability [ 45 ]. Another study using atellocollagen shows that atellocollagen is better than alginate in supporting NP cells with better water and proteoglycan retention [ 98 ].

Nevertheless, most studies on injectable modality using approach 4 are in vitro. Encouraging in vivo studies have been relatively scarce. In a rabbit model, MSCs were loaded in solution atellocollagen [ 99 ] and injected to degenerative discs. Disc degeneration is effectively arrested by the treatment, and evidence of atellocollagen supporting cell growth, differentiation and matrix production is demonstrated.

In a pig model, MSCs were loaded with hyaluronan derivatives [ 93 ] and injected into the nucleotomized discs. Close similarity in disc biconvex structure and viable chondrocytes like cells were reported [ 93 ].

Further enhancement in the swelling and mechanical stability of the scaffolds for nucleus or annulus replacement such as crosslinking without compromising cell viability [ 45 ] should be encouraged as these physical properties favor the restoration and maintenance of disc height.

Evaluation of the mechanical properties of injectable polymers such as hyaluronic acid gel [ 28 ] and polyethylene glycol [ ] against different engineering parameters should also be encouraged as replacing the mechanical function is equally important as replacing the cellular function in matrix secretion in this stage of degeneration.

At the advanced stage of disc degeneration, structural collapse and eventual loss of disc function in resisting loading set in motion. At this stage, probably only an engineered IVD tissue idealized in the discussion above can do the job. Decellularized IVD allograft is theoretically the best scaffold for late stage replacement. To date, there is no decellularization study in intervertebral disc, but there are several in vivo studies transplanting undecellularized allogenic discs.

A few groups have also transplanted cryopreserved allogenic disc grafts in dogs [ 60 , 75 ]. Nevertheless, dramatic decrease in cellular activity has been reported, indicating that preservation of cellular synthetic and remodeling activities is important for long term viability of the grafts. Furthermore, degenerative changes have been reported at 1 year post-implantation, also suggesting the need to preserve better cellular remodeling capacity. Larger animal model in monkeys [ 71 , 72 ] showed that fresh frozen allografts can maintain the mechanical properties and some degree of cell metabolism, but severe degeneration has been observed in 2 years time, associating with decreased biochemical contents of the disc.

This study also suggests the need to preserve or supplement viable cells with remodeling capacity that is able to maintain long term merits of allografting. Very recently, this group conducted a first allograft disc segment transplantation study in human [ 96 ]. The transplanted segments were able to preserve motion and hydration for at least 5 years with partially recovered disc height and improved neurological symptoms, and with no immunoreaction.

Nevertheless, mild degenerative changes have been found after years of follow-up, suggesting again that repopulation or supplementation of live cells with matrix remodeling capacity is necessary for long term functionality of allografts. As a result, a promising direction is to develop better graft preservation technologies for maintaining the disc cell remodeling capacity.

Alternatively, supplementing stem cells or other cells with the ability to respond to physiological challenges in terms of synthesis and secretion of appropriate ECM into the allograft disc space is also promising.

In this regard, searching for injectable carriers to provide an appropriate microenvironment to the encapsulated cells for proper remodeling response toward mechanical stimulation deserves more attention.

Apart from using the allograft, building an implantable IVD with multiple tissue components with mechanical properties comparable with the native disc presents another possibility. Activities with this line of research have been scarce, and only a few groups are attempting to build multiple tissue components and to target the interface problems [ 4 , 46 ].

A few unique considerations should be highlighted as they are the most challenging tasks for tissue engineering. Firstly, IVD is a complex tissue with multiple tissue components. It is unlikely to utilize single biomaterial and single scaffolding approach. Maintaining the differential hydration and mechanical properties of the nucleus and the annulus is important to the maintenance of the disc height and its load-resisting properties.

Engineering designs to better retain the water absorbing capacity in the biomaterials used for nucleus replacement and engineering designs to better maintain the enclosure of the swelling nucleus within the confines of the annulus replacement warrant further attention. Secondly, IVD involves multiple tissue interfaces, which are essential for load transfer and distribution between hard and soft tissues [ 17 , ] and are crucial for maintaining proper disc functions.

Chondrocytes seeded between a pre-made bone scaffold and NP cells resulted in the formation of a cartilage-like layer between bone construct and nucleus cells [ 46 ] while chondrocytes supplemented with osteogenic differentiation signal led to the formation of a calcified zone between bone and cartilage interface [ 4 ]. Both studies demonstrated benefits in mechanical performance, in particular improving the interfacial shear stress.

This suggests that engineering the interfaces among different tissue components should deserve special attention and more enthusiastic research efforts. Disc tissue engineering is by default multidisciplinary, and the scaffolding approach is just one of the disciplines involved. Its success obviously relies on the corroborative efforts from other aspects. For examples, delineation of the etiology of disc degeneration, understanding of the disc nutrition mechanism, cell sourcing, identification of disc specific markers, growth-stimulating and differentiation-stimulating signals, etc.

Scaffolds in engineered tissues are to mimic the ECM in native tissues, at least partially. Unsurprisingly, their functions should mimic the ECM of the target tissue. Over the past few decades, four major scaffolding approaches, namely implanting cell-seeded pre-made porous scaffolds, implanting cell-seeded decellularized allograft or xenograft ECM, implanting laminated cell sheets with secreted ECM and injecting cell encapsulated self-assembled hydrogels, have been developed.

Each approach has its own pros and cons and preferred tissue engineering applications. In planning for tissue engineering for a complex tissue such as IVD, these scaffolding approaches serve as important guidelines and can be used in combinations. Moreover, tissue-specific considerations in relation to the extent of injury, the unique structural functional relationship, multiple tissue composition and interfaces in IVD deserve special attention.

Conflict of interest statement None of the authors has any potential conflict of interest. National Center for Biotechnology Information , U.

Journal List Eur Spine J v. Eur Spine J. Published online Nov Chan 1 and K. Leong 2. Author information Article notes Copyright and License information Disclaimer. Corresponding author. This article has been cited by other articles in PMC. Abstract Scaffolds represent important components for tissue engineering. Introduction Since its emergence in the mids, tissue engineering has continued to evolve as an exciting and multidisciplinary field aiming to develop biological substitutes to restore, replace or regenerate defective tissues [ 59 , 64 ].

Analogous functions of scaffolds and extracellular matrix Apart from blood cells, most, if not all other, normal cells in human tissues are anchorage-dependent residing in a solid matrix called extracellular matrix ECM. Table 1 Functions of extracellular matrix ECM in native tissues and of scaffolds in engineered tissues.

Functions of ECM in native tissues Analogous functions of scaffolds in engineered tissues Architectural, biological, and mechanical features of scaffolds 1. Provides structural support for cells to reside Provides structural support for exogenously applied cells to attach, grow, migrate and differentiate in vitro and in vivo Biomaterials with binding sites for cells; porous structure with interconnectivity for cell migration and for nutrients diffusion; temporary resistance to biodegradation upon implantation 2.

Contributes to the mechanical properties of tissues Provides the shape and mechanical stability to the tissue defect and gives the rigidity and stiffness to the engineered tissues Biomaterials with sufficient mechanical properties filling up the void space of the defect and simulating that of the native tissue 3. Provides bioactive cues for cells to respond to their microenvironment Interacts with cells actively to facilitate activities such as proliferation and differentiation Biological cues such as cell-adhesive binding sites; physical cues such as surface topography 4.

Acts as the reservoirs of growth factors and potentiates their actions Serves as delivery vehicle and reservoir for exogenously applied growth-stimulating factors Microstructures and other matrix factors retaining bioactive agents in scaffold 5.

Provides a flexible physical environment to allow remodeling in response to tissue dynamic processes such as wound healing Provides a void volume for vascularization and new tissue formation during remodeling Porous microstructures for nutrients and metabolites diffusion; matrix design with controllable degradation mechanisms and rates; biomaterials and their degraded products with acceptable tissue compatibility.

Open in a separate window. Scaffolding approaches in tissue engineering Over the last two decades, four major scaffolding approaches for tissue engineering have evolved Fig. Schematic diagram showing different scaffolding approaches in tissue engineering. Table 2 Characteristics of different scaffolding approaches in tissue engineering. Thus supplementation of the culture medium 49 of MSCs with ascorbic acid, dexamethasone dexa , beta-glycerophosphate b-GlyP 50 and 1,dihydroxyvitamin D3 can induce the appearance of several osteoblast features 51 [12—15].

The pH of collagen gel was 63 adjusted at 7. The gels were then freeze-dried using the Christ 68 Model Delta 2—24 KD lyophilizer, Germany, to obtain 3D collagen scaffolds spongious 69 forms.

All experiments were done after 1 week of culture. Specimens were frozen in liquid nitrogen 86 and sectioned with a Leica CM cryotome; the thickness of the sections was 4—6 mm. Specimens 96 were dehydrated in graded ethanol series, critical point dried, and gold-sputtered prior 97 to observation series, critical point dried, and gold-sputtered prior to observation. The MTT is a reliable assay method for measuring cell viability in different scaffolds [17].

The assay is dependent on the cleavage of the yellow tetrazolium salt to the purple formazan crystals by metabolic active cells [18].

The cells cultured on collagen scaffolds disks 5 mm in diameter and 2 mm thick were incubated with 0. As negative control we used collagen with gentian bleu scaffolds. The results were expressed as viability percentage.

The concentration and quality of RNA was determined by spectrophotometry. To evaluate cellular colonization on collagen-Dex and coll-D3 scaffolds we performed Hoechst staining. We observed that all scaffolds tested allow the cells not only to attach but also to migrate inside and colonize the scaffold to form 3D structure Figs. Also, numerous cytoplasmic protrusions towards the substrate, and also between cells, were clearly visible Figs.

These findings suggest a significant increase in cell-cell interactions and a good affinity of MG63 and hFOB 1. MTT test results were given in Figs. However, a small decrease in viability of hFOB 1. This result was confirmed by SEM analysis. Micrographs of human osteoblast-like MG 63 cells on collagen scaffolds after seven days in culture: left panels—phase contrast microscopy; right panels—blue fluorescent staining with Hoechst identifying osteosarcoma cell nuclei observed by epifluorescence microscope.

BSP I gene expression was not observed because osteosarcoma cells do not express this marker. Exposure of cells to collagen-Dex and collagen-D3 scaffolds increased the gene expression of osteocalcin Fig. Compare with cells grown on controls borosilicate glass and collagen scaffolds BSP II gene expression increased in cells on collagen-Dex and collagen-D3 scaffolds Fig. Micrographs of human osteoprogenitor h FOB 1. Data representative of two independent experiments are shown in this paper.

The gene expression of bone sialoglycoprotein I decreased in cells cultured on collagen-Dex and collagen-D3 collagen scaffolds. By contrast, the BSP II gene was weakly expressed in cells grown on borosilicate glass and collagen control scaffolds and increased in cells culture on collagen-Dex and collagen-D3 scaffolds Fig.

MG63 cells grown for seven days on collagen scaffolds: a collagen scaffold, b collagen dexamethasone scaffold, c collagen D3 scaffold, d MG63 cells on collagen scaffold, e MG63 cells on collagen dexamethasone scaffold, f MG63 cells on collagen D3 scaffold.

Figure 4.