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. Extraction
b. Types of cells
3. 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 tissue-
engineered 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 collagen-
based 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 proteins-
THROMBOSPONDIN-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 bone-
forming 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 BMP-
2=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 non-
biodegradable 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, non-
immunogenicity, 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 animal-
derived 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 non-woven 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 TISSUES-
THE COLLAGEN GEL MODEL-The 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 MODEL-
The 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 cell-
secreted 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 MODEL-
The 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 MODEL-
A 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 MODEL-
Two-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 ADDED-
Vascular 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 CELLS-
Vascular 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 PROCESSING-
Restoration 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–co-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 BONDING-
Fibers 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.
References
1. Text books of Principles of tissue engineering. (Second edition). Robert . P. Lanza, Robert Langer, Joseph Vacanti.
2. Accuracy Of Three Techniques To Determine Cell Viability In 3D Tissues Or Scaffolds. Benjamin Gantenbein-Ritter, Ph.D.,1 Esther Potier, Ph.D.,1,2 Stephan Zeiter, D.V.M.,1 Marije Van Der Werf, M.Sc.,1,2 Christoph M. Sprecher, Dipl-Ing.,1 And Keita Ito, M.D., Sc.D. TISSUE ENGINEERING: Part C Volume 14, Number 4, 2008.
3. REPAIR OF MANDIBLE DEFECT WITH TISSUE ENGINEERING BONE IN RABBITS. ANZ J. Surg.2005;75: 1017–1021.
4. Proteins and Their Peptide Motifs in Acellular Apatite Mineralization of Scaffolds for Tissue Engineering. TISSUE ENGINEERING: Part B Volume 14, Number 4, 2008.
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