GITAM DENTAL COLLEGE & HOSPITAL
DEPARTMENT OF ORAL & MAXILLOFACIAL SURGERY
SEMINAR ON
Tissue engineering in oral and maxillofacial surgery Presented By: Dr. Sambhav K Vora III MDS
Contents
1.
Overview
2.
Cells as building blocks a. b.
3.
Extraction Types of cells
Scaffolds a.
Materials
b.
Synthesis
4.
Assembly methods
5.
Tissue culture a.
Bioreactors
6.
Organotypic And Histiotypic Models Of Engineered Tissues
7.
Fiber bonding
8.
References
Tissue engineering is considered as a field in its own right. It is the use of a combination of cells, engineering and materials methods, and suitable biochemical and physio-chemical factors to improve or replace biological functions. While most definitions of tissue engineering cover a broad range of applications, in practice the term is closely associated with applications that repair or replace portions of or whole tissues (i.e., bone, cartilage, blood vessels, bladder, skin etc.). Often, the tissues involved require certain mechanical and structural properties for proper functioning. The term has also been applied to efforts to perform specific biochemical functions using cells within an artificially-created support system (e.g. an artificial pancreas, or a bioartificial liver). The term regenerative medicine is often used synonymously with tissue engineering, although those involved in regenerative medicine place more emphasis on the use of stem cells to produce tissues. Definition – It is "an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function or a whole organ". Tissue engineering has also been defined as "understanding the principles of tissue growth, and applying this to produce functional replacement tissue for clinical use Biologic tissues consist of the cells, the extracellular matrix (made up of a complex of cell secretions immobilized in spaces continuous with cells), and the signaling systems, which are brought into play through differential activation of genes or cascades of genes whose secreted or transcriptional products are responsible for tissue building and differentiation.
The triad of tissue engineering, which is based on the three basic components of biologic tissues. The principal components of scaffolds (into which the extracellular matrix is organized in actual tissues) are collagen biopolymers, mainly in the form of fibers and fibrils. Other forms of polymer organization have also been used (gels, foams, and membranes) for engineering tissue substitutes. The various forms can be combined in the laboratory to create imitations of biopolymer organization in specific tissues. Scaffolds can be enriched with signaling molecules, which may be bound to them or infused into them. The focus of the triad is the prosthesis. The incorporation of cells in reconstituted prosthetic tissue devices often can provide the signals needed for tissue building, but the repertoire of feats of differentiation may be limited (see section on stem cells below). For example, although cultivated allogeneic keratinocytes and dermal fibroblasts, plus a collagen scaffold, can be assembled into a graftable organ that differentiates a fully formed epidermis, having a stratum corneum with barrier properties and a basal lamina, the secondary derivatives such as hair follicles and sebaceous and sweat glands do not develop. Improving the quality and functionality of tissueengineered skin will mean the introduction of new versions of skin that address the clinical needs in a way better than their precursors have addressed them. Oversimplified materials used for the scaffolding component (the extracellular matrix of the tissue being engineered) may be limiting. If the scaffold cannot provide the developmental signals for tissue building needed by the cells that are seeded into it in vitro, or mobilized by it in vivo, tissue building might fail, as it does when a Dacron sleeve is used in vivo to replace a segment of artery. Man-made biopolymers such as poly(L-lactic acid), poly(glycolic acid), polyglycolide, and poly(L-lactide) have built-in ranges of degradation times that may not be
in tune with the required rate of remodeling characteristic of regeneration, because the polymers are not susceptible to breakdown by metalloproteinases and tryptic enzymes, which function normally in the remodeling of collagenbased scaffolds. If they are out of tune with the remodeling activities of cells that occupy the transient scaffolds, including matrix biosynthesis, the process of matrix renewal may be compromised. A potentially valuable attribute of acellular materials installed in vivo as precursors of tissue replacements is their ability to mobilize appropriate cells from contiguous tissues, circulating body fluids, or stem cell sources, making it unnecessary to populate the prosthesis with cells before implantation. Because acellular implants of man-made biopolymers are information poor, cell mobilization and vascularization may fail, and so may the organization of the mobilized cells and their secretory output, needed to regenerate a replacement matrix and a functional tissue. STEM CELLS Scaffolds can be populated with adult-derived cells that are capable of undergoing subsequent differentiation after being cultivated in vitro. In this category are cells of the skin, cartilage, muscle, tendon, ligament, bone, adipose tissue, endothelium, and many others. Aside from skin, the foregoing cell types are harbored as stem cell populations in the marrow, in addition to those of the hematopoietic and immune systems, but the diversity of mesenchymal and possibly other cell types in the marrow still needs to be probed. Stimulating factors, the cytokines, which move some of the cells into the circulation, will be important for engineering acellular scaffolds. Other stem cells are available to tissue engineering, such as the satellite cells found in striated muscle and to some degree keratinocytes of the skin. Where host cells are available, an acellular scaffold, particularly one enhanced with signals and possessing the binding sites needed for cell attachment, can mobilize host cells that will populate the prosthesis. Already,
new sources of stem cells, particularly neuronal stem cells, have been discovered in the adult brain and are opening the door to the reconstitution of nerve tissue for tissue engineering. In addition to the striatum, which harbors extracellular growth factor-responsive stem cells, other central nervous system sources of stem cells in adult vertebrates include the hippocampus and the periventricular subependymal zone. Stem cells giving rise to neuronal and glial phenotypes, from the adult rat hippocampus, are isolated with the help of fibroblast growth factor 2 and are stimulated to differentiate with the help of retinoic acid. Stem cells from these sources are also present in the adult human brain. The discovery that embryonic stem cells can be recovered from human fetal tissue and propagated for long periods without losing their toti- or pluripotency is of huge importance for tissue engineering. How to direct their differentiation is a subject of high current interest. EXTRACELLULAR MATRIX STRUCTURE AND FUNCTION COMPOSITION AND ORGANIZATION One of the most critical elements of tissue engineering is the ability to mimic the ECM scaffolds that normally serve to organize cells into tissues. ECMs are composed of different collagen types, large glycoproteins (e.g., fibronectin, laminin, entactin, osteopontin), and proteoglycans that contain large glycosaminoglycan side chains (e.g., heparan sulfate, chondroitin sulfate, dermatan suflate, keratan sulfate, hyaluronic acid). Although all ECMs share these components, the organization, form, and mechanical properties of ECMs can vary widely in different tissues depending on the chemical composition and three-dimensional organization of the specific ECM components that are present. For example, interstitial collagens (e.g., types I and III) self-assemble into a three-dimensional lattice, which, in turn, binds fibronectin and proteoglycans. This type of native ECM hydrogel forms the backbone of loose
connective tissues, such as dermis. In contrast, basement membrane collagens (types IV and V) assemble into planar arrays; when these collagenous sheets interact with fibronectin, laminin, and heparan sulfate proteoglycan, a planar ECM results (i.e., the “basement membrane�). The ability of tendons to resist tension and of cartilage and bone to resist compression similarly result from local differences in the organization and composition of the ECM. Matricellular proteinsTHROMBOSPONDIN-1 AND THROMBOSPONDIN-2 Thrombospondin-1 is a 450,000-Da glycoprotein with seven modular domains. At least five different extracellular matrix-associated proteins are able to bind to thrombospondin-1: collagens I and V, fibronectin, laminin, fibrinogen, and SPARC . One cytokine known to interact with thrombospondin-1 is scatter factor/hepatocyte growth factor (HGF), a known angiogenesis-promoting factor. Thrombospondin-1 has also been shown to interact specifically with another cytokine, transforming growth factor (TGF- ). Thrombospondin-1 is also able to influence cell adhesion and cell shape. For example, it will diminish the number of focal adhesions of bovine aortic endothelial cells and thus will promote a migratory phenotype. Thrombospondin-1, therefore, has been proposed to modulate cell–matrix interaction to allow for cell migration when necessary. Thrombospondin-1 and -2 can act as negative regulators of cell growth. In particular, endothelial cells are susceptible to an inhibition of proliferation by both proteins and, as such, have been classified as inhibitors of angiogenesis TENASCIN-C Tenascin-C is a matricellular protein with a widespread pattern of developmental expression in comparison to a restricted pattern of expression in adult tissues
OSTEOPONTIN Osteopontin associates with the extracellular matrix, in that it binds to fibronectin and to collagens I, II, III, IV, and V. Osteopontin also affects cellular signaling pathways by virtue of its capacity to act as a ligand for multiple integrin receptors as well as CD44 (Denhardt and Noda, 1998; Weber et al., 1996). Thus osteopontin, like most of the matricellular proteins, is able to act as a bridge between the extracellular matrix and the cell surface. The promotion of cell survival is another property ascribed to osteopontin. Finally, osteopontin appears to be involved in inflammatory responses. Expression of osteopontin was found to increase during intradermal macrophage infiltration, and purified osteopontin injected into the rat dermis led to an increase in the number of macrophages at the site of administration. SPARC SPARC (also known as BM-40 and osteonectin) was first identified as a primary component of bone but has since been known to have a wider distribution. SPARC is found in the gut epithelium, which normally exhibits rapid turnover, and in healing wounds. Another significant effect of SPARC on cells in culture is its capacity to elicit changes in cell shape. Role of growth factors in bone healing:5 Growth factors that can help in bone healing are
Platelet derived growth factor
Transforming growth factor
Insulin deriver growth factor
Fibroblast derived growth factor Platelet derived growth factor-
It is released from platelet alpha granule, macrophages or monocytes,endothelial cells & as well as from osteoblast cells. The specific activities of PDGF includes mitogenesis(increase in the cell population of healing cells), angiogenesis(endothelial mitosis into functional capillaries), & macrophage activation(debridement of the wound site &second phase source of growth factors for continued repair & bone regeneration). Transforming growth factor: It is present in abundance in the bone matrix, with bone representing the major site for the storage of the TGF –beta in the body.the primary effect of TGF –beta is on the bone formation, particularly in the early phase of the osteoblast development.it stimulates matrix protein synthesis by human osteoblasts.the most important function of TGF beta 1 & TGF beta 2seems to be chemotaxis &mitogenesis of osteoblast precursors, & they also have the ability to stimulate osteoblast deposition of the collagen matrix of wound healing & of the bone. In addition they also inhibit osteoclast formation & also bone resorption, thus favoring bone formation than resorption. It directly inhibits both proliferation & differentiation of the osteoclast precursor cells & inhibits the function of the mature osteoclasts with reduction in reactive e oxygen radicals. Insulin like growth factor: It consists of two proteins-IGF 1(somatomedin c) & IGF2 (skeletal growth factor) which are secreted by osteoblasts,both the factors induce preosteoblast proliferation & differentiation, osteoblast collagen synthesis,& inhibit collagen breakdown.IGF bound to the protein in the matrix may be released in the active form following osteoclastic
resoeption.locaaly produced IGF1secreted by fibroblast& cells in the bone & cartilage is controlled by variety of factors. Corticosteroids reduces IGF1 synthesis.
Fibroblast growth factor: The matrix proteins, acidic FGF & basic FGFare produced by osteoblast,bind heparin & are angiogenic factors. But there effects on bone invivo are not known.in vitro they cause proliferation of osteoblast progenitor cells but inhibit differentiation, & do not appear to effect the osteoclast.FGFs stimute new bone formation.
BONE MORPHOGENETIC PROTEIN: One group of cytokines, bone morphogenetic proteins (BMPs), has been demonstrated to have true osteoinductive properties. BMPs have been proven to stimulate new bone formation in vitro and in vivo. In addition, they play critical roles in regulating cell growth, differentiation, and apoptosis a variety of cells during development, particularly in osteoblasts and chondrocytes. There are currently 16 identified BMPs, although only a subset have been found to be expressed in fracture healing. BMPs were initially characterized by Urist; their identification was based on the capacity of demineralized bone powder to induce de novo bone formation in an intramuscular pouch, demonstrating the ability to directly induce mesenchymal connective tissue to become boneforming osteoprogenitor cell.
During fracture repair & graft healing, BMP-2, BMP-3 (also known as osteogenin), BMP-4, and BMP-7 (OP-1) have been found to be expressed to varying degrees. BMPs are initially released in low levels from the extracellular matrix (ECM) of fractured bone. Osteoprogenitor cells in the cambium layer of the periosteum may respond to this initial BMP presence by differentiating into osteoblast. Immunolocalization demonstrates an increase in detectable BMP2=4 in the cambium region of the periosteum. BMP receptor IA and IB expression is dramatically increased in osteogenic cells of the periosteum near the ends of the fracture in the early postfracture period or post grafting period. Approximately 1–2 weeks postfracture or graft placement, BMP-2=4 expression is maximal in chondroid precursors, while hypertrophic chondrocytes and osteoblasts show moderate levels of expression. It is hypothesized that the role of BMPs in fracture repair is to stimulate differentiation in osteoprogenitor and mesenchymal cells that will result in osteoblasts and chondrocyte. As these primitive cells mature, BMP expression decreases rapidly. BMP expression temporarily recurs in chondrocytes and osteoblasts during matrix formation, and eventually decreases during callus remodeling. TYPES OF BMPs THEIR PROPERTIES, LOCATION & ROLES: BMP-1:
functions as procollagen C- proteinase responsible for removing
carboxyl propeptides from procollagen I, 2 ,3 . It activates bmp but not osteoinductive BMP-2:
osteoinductive , embryogenesis, differentiation of osteoblasts ,
adipocytes, chondrocytes & also may influence osteoclast activity , may inhibit bone healing It is located in the bone, spleen, liver, brain, kidney, heart, placenta.
BMP-3:(osteogenin)-
osteoinductive , promotes chondrogenic phenotype
It is located in the lung, kidney, brain, intestine. BMP-4:
osteoinductive, embryogenesis, fracture repair, gastrulation &
mesoderm formation (mouse). It is located in the apical ectodermal ridge, meninges, lung, kidney, liver. BMP-5:-osteoinductive, embryogenesis. It is located in lung, kidney, and liver. BMP-6:- not osteoinductive, embryogenesis, neuronal maturation, regulates chondrocyte differentiation. It is located in the lung, brain, kidney, uterus, muscle, skin. BMP-7:-(osteogenic protein-1) osteoinductive, embryogenesis, repair of long bones, alveolar bone, differentiation of osteoblasts,chondroblasts & adipocytes. It is located in the adrenal glands, bladder, brain, eye, heart, kidney, lung, placenta, spleen & skeletal muscles. BMP-8(osteogenic protein-2) osteoinductive, embrogenesis, spermatogenesis(mouse). BMP-8B(osteogenic protein-3)initiation & maintainance of spermatogenesis(mouse). BMP-9:-osteoinductive, stimulates hepatocyte proliferation, hepatocyte growth & function. BMP-12 & BMP-13:-inhibition of terminal differentiation of myoblasts.
Cells as building blocks
Tissue engineering utilizes living cells as engineering materials. Examples include using living fibroblasts in skin replacement or repair, cartilage repaired with livingchondrocytes, or other types of cells used in other ways. Cells became available as engineering materials when scientists at Geron Corp. discovered how to extend telomeres in 1998, producing immortalized cell lines Before this, laboratory cultures of healthy, noncancerous mammalian cells would only divide a fixed number of times, up to the Hayflick limit. Extraction From fluid tissues such as blood, cells are extracted by bulk methods, usually centrifugation or apheresis. From solid tissues, extraction is more difficult. Usually the tissue is minced, and then digested with the enzymes trypsin or collagenase to remove the extracellular matrix that holds the cells. After that, the cells are free floating, and extracted using centrifugation or apheresis. Digestion with trypsin is very dependent on temperature. Higher temperatures digest the matrix faster, but create more damage. Collagenase is less temperature dependent, and damages fewer cells, but takes longer and is a more expensive reagent. Types of cells Cells are often categorized by their source:
Autologous cells are obtained from the same individual to which they
will be reimplanted. Autologous cells have the fewest problems with rejection and pathogen transmission, however in some cases might not be available. For example in genetic disease suitable autologous cells are not available. Also very ill or elderly persons, as well as patients suffering from severe burns, may not have sufficient quantities of autologous cells to establish useful cell lines. Moreover since this category of cells needs to be harvested from the patient, there are also some concerns related to the necessity of performing such surgical operations that might lead to donor site infection or chronic pain. Autologous cells also must be cultured from samples before they can be used: this takes time, so autologous solutions may not be very quick. Recently there has been a trend towards the use of mesenchymal stem cells frombone marrow and fat. These cells can differentiate into a variety of tissue types, including bone, cartilage, fat, and nerve. A large number of cells can be easily and quickly isolated from fat, thus opening the potential for large numbers of cells to be quickly and easily obtained.
Allogeneic cells come from the body of a donor of the same species.
While there are some ethical constraints to the use of human cells for in vitro studies, the employment of dermal fibroblasts from human foreskin has been demonstrated to be immunologically safe and thus a viable choice for tissue engineering of skin.
Xenogenic cells are these isolated from individuals of another species. In
particular animal cells have been used quite extensively in experiments aimed at the construction of cardiovascular implants.
Syngenic or isogenic cells are isolated from genetically identical
organisms, such as twins, clones, or highly inbred research animal models.
Primary cells are from an organism.
Secondary cells are from a cell bank.
. Scaffolds Cells are often implanted or 'seeded' into an artificial structure capable of supporting three-dimensional tissue formation. These structures, typically called scaffolds, are often critical, bothex vivo as well as in vivo, to recapitulating the in vivo milieu and allowing cells to influence their own microenvironments. Scaffolds usually serve at least one of the following purposes:
Allow cell attachment and migration
Deliver and retain cells and biochemical factors
Enable diffusion of vital cell nutrients and expressed products
Exert certain mechanical and biological influences to modify the
behaviour of the cell phase
This animation of a rotating Carbon nanotubeshows its 3D structure. Carbon nanotubes are among the numerous candidates for tissue engineering scaffolds since they arebiocompatible, resistant to biodegradation and can be functionalized with biomolecules. However, the possibility of toxicity with nonbiodegradable nano-materials is not fully understood.
To achieve the goal of tissue reconstruction, scaffolds must meet some specific requirements. A high porosity and an adequate pore size are necessary to facilitate cell seeding and diffusion throughout the whole structure of both cells and nutrients. Biodegradability is often an essential factor since scaffolds should preferably be absorbed by the surrounding tissues without the necessity of a surgical removal. The rate at which degradation occurs has to coincide as much as possible with the rate of tissue formation: this means that while cells are fabricating their own natural matrix structure around themselves, the scaffold is able to provide structural integrity within the body and eventually it will break down leaving the neotissue, newly formed tissue which will take over the mechanical load. Injectability is also important for clinical uses. Materials Many different materials (natural and synthetic, biodegradable and permanent) have been investigated. Most of these materials have been known in the medical field before the advent of tissue engineering as a research topic, being already employed as bioresorbable sutures. Examples of these materials are collagen and some polyesters. New biomaterials have been engineered to have ideal properties and functional customization: injectability, synthetic manufacture,biocompatibility, nonimmunogenicity, transparency, nano-scale fibers, low concentration, resorption rates, etc. PuraMatrix, originating from the MIT labs of Zhang, Rich, Grodzinsky and Langer is one of these new biomimetic scaffold families which has now been commercialized and is impacting clinical tissue engineering. A commonly used synthetic material is PLA - polylactic acid. This is a polyester which degrades within the human body to form lactic acid, a naturally occurring chemical which is easily removed from the body. Similar materials are polyglycolic acid (PGA) and polycaprolactone(PCL): their degradation
mechanism is similar to that of PLA, but they exhibit respectively a faster and a slower rate of degradation compared to PLA. Scaffolds may also be constructed from natural materials: in particular different derivatives of the extracellular matrix have been studied to evaluate their ability to support cell growth. Proteic materials, such as collagen or fibrin, and polysaccharidic materials, like chitosan or glycosaminoglycans (GAGs), have all proved suitable in terms of cell compatibility, but some issues with potential immunogenicity still remains. Among GAGs hyaluronic acid, possibly in combination with cross linking agents (e.g.glutaraldehyde, water soluble carbodiimide, etc...), is one of the possible choices as scaffold material. Functionalized groups of scaffolds may be useful in the delivery of small molecules (drugs) to specific tissues. Synthesis A number of different methods has been described in literature for preparing porous structures to be employed as tissue engineering scaffolds. Each of these techniques presents its own advantages, but none is devoid of drawbacks. 
Nanofiber Self-Assembly: Molecular self-assembly is one of the few
methods to create biomaterials with properties similar in scale and chemistry to that of the natural in vivo extracellular matrix (ECM). Moreover, these hydrogel scaffolds have shown superior in vivo toxicology and biocompatibility compared with traditional macroscaffolds and animalderived materials. 
Textile technologies: these techniques include all the approaches that
have been successfully employed for the preparation of non-woven meshes of different polymers. In particular nonwoven polyglycolide structures have been tested for tissue engineering applications: such fibrous structures have been found useful to grow
different types of cells. The principal drawbacks are related to the difficulties of obtaining high porosity and regular pore size. 
Solvent Casting & Particulate Leaching (SCPL): this approach allows
the preparation of porous structures with regular porosity, but with a limited thickness. First the polymer is dissolved into a suitable organic solvent (e.g. polylactic acid could be dissolved into dichloromethane), then the solution is cast into a mold filled with porogen particles. Such porogen can be an inorganic salt like sodium chloride, crystals of saccharose, gelatin spheres or paraffin spheres. The size of the porogen particles will affect the size of the scaffold pores, while the polymer to porogen ratio is directly correlated to the amount of porosity of the final structure. After the polymer solution has been cast the solvent is allowed to fully evaporate, then the composite structure in the mold is immersed in a bath of a liquid suitable for dissolving the porogen: water in case of sodium chloride, saccharose and gelatin or an aliphatic solvent like hexane for paraffin. Once the porogen has been fully dissolved a porous structure is obtained. Other than the small thickness range that can be obtained, another drawback of SCPL lies in its use of organic solvents which must be fully removed to avoid any possible damage to the cells seeded on the scaffold. 
Gas Foaming: to overcome the necessity to use organic solvents and
solid porogens a technique using gas as a porogen has been developed. First disc shaped structures made of the desired polymer are prepared by means of compression molding using a heated mold. The discs are then placed in a chamber where are exposed to high pressure CO2 for several days. The pressure inside the chamber is gradually restored to atmospheric levels. During this procedure the pores are formed by the carbon dioxide molecules that abandon the polymer, resulting in a sponge like structure. The main problems related to such a technique are caused by the excessive heat used during compression molding (which prohibits the incorporation of any
temperature labile material into the polymer matrix) and by the fact that the pores do not form an interconnected structure. 
Emulsification/Freeze-drying: this technique does not require the use of
a solid porogen like SCPL. First a synthetic polymer is dissolved into a suitable solvent (e.g. polylactic acid in dichloromethane) then water is added to the polymeric solution and the two liquids are mixed in order to obtain an emulsion. Before the two phases can separate, the emulsion is cast into a mold and quickly frozen by means of immersion into liquid nitrogen. The frozen emulsion is subsequently freeze-dried to remove the dispersed water and the solvent, thus leaving a solidified, porous polymeric structure. While emulsification and freeze-drying allows a faster preparation if compared to SCPL, since it does not require a time consuming leaching step, it still requires the use of solvents, moreover pore size is relatively small and porosity is often irregular. Freeze-drying by itself is also a commonly employed technique for the fabrication of scaffolds. In particular it is used to prepare collagen sponges: collagen is dissolved into acidic solutions of acetic acid or hydrochloric acid that are cast into a mold, frozen with liquid nitrogen then lyophilized. 
Thermally Induced Phase Separation (TIPS): similar to the previous
technique, this phase separation procedure requires the use of a solvent with a low melting point that is easy to sublime. For example dioxane could be used to dissolve polylactic acid, then phase separation is induced through the addition of a small quantity of water: a polymer-rich and a polymer-poor phase are formed. Following cooling below the solvent melting point and some days of vacuum-drying to sublime the solvent a porous scaffold is obtained. Liquid-liquid phase separation presents the same drawbacks of emulsification/freeze-drying. 
CAD/CAM Technologies: since most of the above described approaches
are limited when it comes to the control of porosity and pore size, computer
assisted design and manufacturing techniques have been introduced to tissue engineering. First a three-dimensional structure is designed using CAD software, then the scaffold is realized by using ink-jet printing of polymer powders or through Fused Deposition Modeling of a polymer melt.
ORGANOTYPIC AND HISTIOTYPIC MODELS OF ENGINEERED TISSUESTHE COLLAGEN GEL MODELThe model uses a collagen gel scaffold prepared by combining, in the cold, a neutralized 0.3–1.0 mg/ml solution of acid extracted collagen with medium, serum, and mesenchymal cells (Bell et al., 1979)—dermal fibroblasts, for example, if the goal is to fabricate a skin equivalent. The bacteriological petri plate or other vessel, to which cells do not attach, into which the mix is poured, is incubated at 37"C in a 5% CO2 incubator. The collagen polymerizes, when neutralized and warmed, forming a lattice of fine fibrils, 10–20 nm in size, which trap fluid. The result is a gel in which the previously added cells are distributed. The mesenchymal cells in the gel extend and attach podial processes to the collagen fibrils and withdraw the processes with the attached fibrils toward the cell body. As the fibrils are bundled by the cells, fluid is squeezed out of the lattice. The process of gel contraction, known as syneresis, can reduce the size of the collagen lattice 30- to 40- fold, depending on the cell and collagen concentration used. The condensed gel is tissuelike in its consistency, providing a substrate on which epithelial, endothelial, or mesothelial cells may be plated.
THE SKIN EQUIVALENT AS A DEVELOPMENTAL MODELThe first skin consisting of “living� dermis and epidermis reconstituted from cultivated cells and collagen scaffolding (Bell et al., 1979, 1981, 1983) was shown to undergo virtually complete differentiation in vitro, lacking, however, pigment, sweat glands, neurogenic elements, a micro-circulation, and hair follicles. The model can be reproduced faithfully and be kept alive in vitro for months, at least. Although collagenolytic activity is high in young dermal equivalents (Nusgens et al., 1984; Rowling et al., 1990), possibly associated with tissue remodeling, it has been observed that the resistance of dermal equivalents to breakdown by collagenase is greatly enhanced by 30 days of cultivation in vitro, suggesting that extensive cross-linking (probably by cellsecreted lysyl oxidase) of the collagen fibrils has occurred, as shown by Rowling et al. (1990). Continued differentiation of the model in vitro and the resemblance of cells in the matrix to their in vivo counterparts, rather than to cells grown on plastic in two dimensions, are distinguishing feature. THE SKIN EQUIVALENT AS AN IMMUNOLOGICAL MODELThe skin equivalent can be constituted with cultivated parenchymal cells free of any subsets of immune cells normally found in the dermis and epidermis. Using the X chromosome as a genetic marker, female cells are used to make up skin equivalents, which are then transplanted to male hosts across a major histocompatibility barrier, e.g., from Brown Norway to Fisher rats. Sher et al. (1983) demonstrated in the rat model, by karyotyping cells grafted in skin equivalents, that allogeneic fibroblasts were not rejected. It has been reported that clinical trials of skin equivalents made up with human allogeneic keratinocytes as well as allogeneic fibroblasts do not provoke an immune reaction in recipients (Parenteau et al., 1994). The model should be a valuable tool for determining the roles played by cells of the immune system
and the microcirculation in allograft rejection of actual skin. It should allow use of cells of any genotype and of human origin to study genetic abnormalities, as well as the contribution of specific genetic loci to skin development by transplanting skin equivalents to immunodeficient rodents.
THE SKIN EQUIVALENT AS A DISEASE MODELA psoriasis model was fabricated to test the contribution of psoriatic dermal fibroblasts to the expression of features of the disease in vitro (Saiag et al., 1985). A button of normal keratinocytes suspended in medium was plated in the centers of dermal equivalent disks constituted with normal human or psoriatic dermal fibroblasts, and the rate of spreading of keratinocytes over the dermal substrate was measured. It was observed that the psoriatic fibroblasts induced hyperproliferation and greater spreading of keratinocytes compared with the growth and spreading induced by control fibroblasts, suggesting that dermal cells may play a role in the progress of the disease. In addition to the study of psoriasis, and other epidermal diseases such as epidermolysis bullous, the model should provide an in vitro basis for studying dermal connective tissue disorders, including dermatosparaxiis and sclerosis. It is obvious that any pair of populations of mesenchymal cells and epithelial cells, of which one or both is diseased or aberrant, can be used in the threedimensional coculturing system for studying the expression of features of the disease and testing modalities of treatment. THE SKIN EQUIVALENT AS A WOUND HEALING MODELTwo-tissue skin models can be used in vitro to analyze the role of dermis in epidermal wound healing (Bell and Scott, 1992). After constituting a differentiated skin equivalent in a 24-mm multiwall well-plate insert, a central
disk of the skin is removed with a punch. The acellular layer of collagen in contact with the membrane of the insert is replaced and the remainder of the gap is filled with a collagen scaffold to the level of the interface between the dermis and epidermis. The rate of overgrowth of the neodermis by keratinocytes and the development of the epidermis can be taken as measures of the effectiveness of the design of neodermis as an interacting substrate. The wound healing model can accommodate acellular dermal scaffolds with or without signals to test their effectiveness in attracting dermal fibroblasts from the surrounding matrix. VASCULAR MODELS WITH CELLS ADDEDVascular models that examine the effect of shear and other forces on monolayers of endothelial cells in vitro have been developed by Nerem et al. (1993), who have shown that the rate of endothelial cell proliferation is decreased by flow and that entry of cells into a cycling state is inhibited. They suggest that a coculture system in which the endothelium is supported on smooth muscle tissue would be superior for providing a more physiologic environment. Such a system was developed by Weinberg and Bell (1986), who showed that a basal lamina was laid down between the endothelium and the contiguous smooth muscle tissue, cast in the form of a small-caliber tube in vitro. Fabricating the vessel was a three-step procedure. The first tissue layer cast around a small caliber mandrel was a smooth muscle cell media, whose ends were anchored in a velcro cuff or held fast by ridges and grooves in the mandrel until radial contraction made space for bands that were secured around the ends. Hence the mechanical restraint imposed on the contracting tube allowed contraction to occur radially but not longitudinally, because each end of the vessel equivalent was held fast. The second tissue layer cast was the adventitia surrounding the smooth muscle media. To make room for it, the fluid expressed from the collagen gel scaffold was drawn out of the casting tube, and the mixture of adventitial fibroblasts
suspended in medium containing neutralized collagen was introduced into the space between the media and the wall of the cylindrical casting chamber. After the adventitia had contracted radially but not longitudinally, because the media provided a frictional surface that prevented it, the mandrel was extracted, leaving a lumen of the tissue tube that was filled with a suspension of endothelial cells. The cells came to rest on the inner surface of the media as the tube was rotated. VASCULAR MODELS WITHOUT CELLSVascular prostheses constructed from Dacron and other synthetics have been in use for many years but are known to elicit persistent inflammatory reactions and to become occluded. The thermosetting polymers are not biodegradable and do not integrate with host tissues, but some successes have been reported under limited conditions in experimental animals. For the foregoing reasons, other acellular materials have been proposed and tried as arterial substitutes. Animal tissues that resemble arteries have been used with some success, in particular, the porcine small intestine (Sandusky et al., 1992). The mucosal cells are scraped off the luminal side and the muscular layers are removed from the abluminal side, leaving the stratum compactum, a dense, highly organized fibrillar collagen matrix and the looser connective tissue of the mucosa. The material can be used as a scaffold for cells in vitro and has been used in animal experiments. In vivo, it is invaded by capillaries that contribute cells that provide an intima, whereas smooth muscle cells migrating from the anastomoses provide a media.
SOFT LITHOGRAPHY As the need of biologists to control and manipulate materials on the micrometer scale has increased, so has the need for new microfabrication techniques. Our
laboratory has developed a set of microfabrication techniques that are useful for patterning on the scale of 0.5 m and larger. We call these techniques “soft lithography” because they use elastomeric (that is, soft) stamps, molds, membranes, or channels (Xia and Whitesides, 1998). Many other techniques can and have been used to pattern cells and their environment (Hammarback et al., 1985; López et al., 1993a; Park et al., 1998; Vaidya et al., 1998). The most commonly used method has been photolithography. This technique has, of course, been highly developed for the microelectronics industry; it has also been adapted, with varying degrees of success, for biological studies (Hammarback et al., 1985; Kleinfeld et al., 1988; Ravencroft et al., 1998). As useful and powerful as photolithography is (it is capable of mass production at 200-nm resolution of multilevel, registered structures), it is not always the best or only option for biological studies. It is an expensive technology; it is poorly suited for patterning nonplanar surfaces; it provides almost no control over the chemistry of the surface and hence is not very flexible in generating patterns of specific chemical functionalities or proteins on surfaces; it can generate only two dimensional microstructures; and it is directly applicable to patterning only a limited set of photosensitive materials (e.g., photoresists). Soft lithographic techniques are inexpensive, are procedurally simple, are applicable to the complex and delicate molecules often required in biochemistry and biology, can be used to pattern a variety of different materials, are applicable to both planar and nonplanar substrates ( Jackman et al., 1995), and do not require stringent control (such as a clean room environment) over the laboratory environment beyond that required for routine cell culture (Xia and Whitesides, 1998). Access to photolithographic technology is required only to create a master for casting the elastomeric stamps or membranes, and even then, the requirement for chrome masks—the preparation of which is one of the slowest and most expensive steps in conventional photolithography—can often be bypassed (Deng et al., 1999; Duffy et al., 1998; Grzybowski et al., 1998; Qin
et al., 1996). Soft lithography offers special advantages for biological applications, in that the elastomer most often used (PDMS) is optically transparent and permeable to gases, is flexible and seals conformally to a variety of surfaces (including petri dishes), is biocompatible, and can be implanted if desired. The soft lithographic techniques that we will discuss include microcontact printing, patterning with microfluidic channels, and laminar flow patterning. SELF-ASSEMBLED MONOLAYERS Because many of the studies involving the patterning of proteins and cells using soft lithography have been carried out on self-assembled monolayers (SAMs) of alkane thiolates on gold, we give a brief discussion of SAMs (Bain and Whitesides, 1988b; Bishop and Nuzzo, 1996; Delamarche and Michel, 1996; Dubois and Nuzzo, 1992; Merritt et al., 1997; Ostuni et al., 1999; Prime and Whitesides, 1993; Ulman, 1996). SAMs are organized organic monolayer films (Fig. 18.1A) that allow control at the molecular level over the chemical properties of the interface by judicious design and fabrication of derivatized alkane thiol(s) adsorbed to the surface of films of gold or silver. The ease of formation of SAMs, and their ability to present a range of chemical functionality at their interface with aqueous solution, make them particularly useful as model surfaces in studies involving biological components. Furthermore, SAMs can be easily patterned by simple methods such as microcontact printing ( CP) with features down to 500 nm in size and smaller (Xia and Whitesides, 1998). These features of SAMs make them the best structurally defined substrates for use in patterning proteins and cells. SAMs on gold are used for the majority of experiments requiring the patterning of proteins and cells, because they are biocompatible, easily handled, and chemically stable [for example, silver oxidizes relatively rapidly, and Ag(I) ions are cytotoxic].
POLYMER SCAFFOLD PROCESSINGRestoration of organ function by utilizing tissue engineering technologies often requires the use of a temporary porous scaffold. The function of the scaffold is to direct the growth of cells migrating from surrounding tissue (tissue conduction) or the growth of cells seeded within the porous structure of the scaffold. The scaffold must therefore provide a suitable substrate for cell attachment, proliferation, differentiated function, and, in certain cases, cell migration. These critical requirements can be met by the selection of an appropriate material from which to construct the scaffold, although the suitability of the scaffold may also be affected by the processing technique. Many biocompatible materials can be potentially used to construct scaffolds. However, a biodegradable material is normally desired because the role of the scaffold is usually only a temporary one. Many natural and synthetic biodegradable polymers, such as collagen, poly(2-hydroxyesters), and poly(anhydrides), have been widely and successfully used as scaffold materials due to their versatility and ease of processing (Thomson et al., 1995). Many researchers have used poly(2-hydroxyesters) as starting materials from which to fabricate scaffolds using a wide variety of processing techniques. These polymers have proved successful as temporary substrates for a number of cell types, allowing cell attachment, proliferation, and maintenance of differentiated function. Poly(2-hydroxyesters), such as poly(L-lactic acid) (PLLA), poly lactic窶田o-glycolic acid(PLGA) copolymers, and poly(glycolic acid) (PGA), are linear, uncross-linked polymers. These materials are biocompatible, degradable by simple hydrolysis, and are Food and Drug Administration (FDA) approved for certain clinical applications. The mechanical properties of the scaffold are often of critical importance especially when regenerating hard tissues such as cartilage and bone. Although the properties of the solid polymer and the porosity of the scaffold have a profound effect on its mechanical properties, polymer processing can also be influential in
this respect. The tensile strength may, for example, be enhanced due to the crystallization of polymer chains. Alternatively, the manufacturing process may cause a reduction in the molecular weight of the polymer, resulting in a deleterious effect on mechanical properties. The shape of a hard tissue is often important to its function and in such cases the processing technique must allow the preparation of scaffolds with irregular three-dimensional geometries. FIBER BONDINGFibers provide a large surface area: volume ratio and are therefore desirable as scaffold materials. One of the first biomedical uses of PGA was as a degradable suture material, which is why it is commercially available in the form of long fibers. PGA fibers in the form of tassels and felts were utilized as scaffolds in organ regeneration feasibility studies (Cima et al., 1991). However, these scaffolds lacked the structural stability necessary for in vivo use. A fiber bonding technique was therefore developed to prepare interconnecting fiber networks with different shapes for use as scaffolds in organ regeneration (Mikos et al., 1993a). PLLA is dissolved in methylene chloride, a nonsolvent for PGA, and the resulting polymer solution is cast over a nonwoven mesh of PGA fibers in a glass container. The solvent is allowed to evaporate and residual amounts are removed by vacuum drying. A composite material is thus produced consisting of nonbonded PGA fibers embedded in a PLLA matrix. The PLLA– PGA composite is then heated to a temperature above the melting point of PGA for a given time period. During heating the PGA fibers join at their cross-points as melting commences, but the two polymers do not join due to their immiscibility in the melt state. The composite is quenched to prevent any further melting of the PGA fibers during cooling. After heat treatment, the PLLA matrix of a PLLA–PGA composite membrane is selectively dissolved in methylene chloride and the resulting bonded PGA fibers are vacuum dried.
Using this technique, the fibers are physically joined without any surface or bulk modification and retain their initial diameter. The PLLA matrix is required to prevent collapse of the PGA mesh and to confine the melted PGA to a fiberlike shape (Fig. 21.1). The heating time is also of critical importance because prolonged exposure to the elevated temperature results in the gradual transformation of the PGA fibers into spherical domains.
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