Materials Science and Engineering R 34 (2001) 147±230
Review
Polymeric biomaterials for tissue and organ regeneration
B.L. Seala, T.C. Oterob, A. Panitcha,*
a
Department of Bioengineering, Arizona State University, Tempe, AZ 85287-9709, USA
Department of Chemical and Materials Engineering, Arizona State University, Tempe, AZ 85287-6006, USA
b
Accepted 25 June 2001
Abstract
This paper reviews recent work involving polymeric biomaterials used for skin, cartilage, bone, vascular,
nerve and liver regeneration. Skin trauma involves damage to the epidermal, dermal and/or subdermal tissues.
Epithelial, dermal and full-thickness replacements are considered. Cartilage research is mainly focused on replacing
hyaline cartilage. Researchers investigate both nondegradable polymers, which must provide mechanical stability,
and degradable polymers, which must support cartilage regeneration. Natural healing in large bone defects often
fails. Materials for bone reconstruction must be biocompatible, offer mechanical properties similar to those of bone
and support tissue regeneration. The area of vascular grafts draws attention as improvements to the patency of
existing materials are needed. Studies to improve current vascular graft polymers as well as develop new polymers
are reviewed below. The design and testing of materials for nerve regeneration, to repair damage caused by illness or
accident, is an active area of research. Directional nerve guidance via tubulation is discussed, as are matrix materials
to enhance axonal extension. Finally, liver transplantation remains one of the only options for chronic liver disease
and the demand for liver transplants far exceeds the number of available organs. The complexity of parameters
involved in liver regeneration is presented here. # 2001 Elsevier Science B.V. All rights reserved.
Keywords: Polymeric biomaterials; Tissue engineering
1. Introduction
This paper reviews recent work in polymeric biomaterials for skin, cartilage, bone, vascular,
nerve, and liver regeneration. There are other tissue and organ systems under current investigation,
however, we have chosen to limit ourselves to some of the applications involving tissue
reconstruction of less complex systems with the exception of liver. Liver regeneration is included as
an example of the tremendous materials and biological challenges that lie before us as we consider
how to recreate complex organs. Since a review of all known or currently used polymeric materials
would be quite cumbersome, we have also chosen to limit ourselves to more current work within
skin, cartilage, bone, vascular, nerve, and liver research. As a whole, the use of polymeric materials
in tissue and organ regeneration has enjoyed widespread use in many diverse applications. Other
applications equally as important to the topics reviewed below involve the use of polymeric
biomaterials for drug delivery, cell encapsulation, striated and cardiac muscle regeneration, heart
valve regeneration, and tendon reconstruction to name a few.
Polymeric biomaterials have several important uses in addition to tissue reconstruction.
Examples include poly(methyl methacrylate) bone cement, poly(glycolic acid) degradable sutures,
poly(glycolic-co-lactic acid) bone screws, and poly(vinyl siloxane) dental impression materials.
*
Corresponding author. Tel.: ‡1-480-965-3028; fax: ‡1-480-965-7624.
0927-796X/01/$ ± see front matter # 2001 Elsevier Science B.V. All rights reserved.
PII: S 0 9 2 7 - 7 9 6 X ( 0 1 ) 0 0 0 3 5 - 3
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Fig. 1. Illustration of how some material, biological, medical, and engineering properties must be integrated to achieve
successful biomaterials for tissue regeneration.
Polymers such as poly(ethylene glycol) often are used to extend the circulation half-life of some
drugs. Poly(hydroxyethyl methacrylate) is used to create soft contact lenses. Although, they have
drastically improved the quality of life of many people, these additional applications are beyond the
scope of this review.
The field of biomaterials resulted from a marriage of disciplines including the life sciences,
medicine, materials science, and engineering. This field is not new; throughout history there are
references to glass eyes and wooden and gold-filled teeth [1]. A relatively new development,
however, is the close working relationships between the above-mentioned disciplines to ensure that
the appropriate parameters, as we know them, are being addressed. Fig. 1 illustrates how some of
these factors integrate into the appropriate design of new materials. Figs. 3±9 depict the structure of
the biological systems reviewed. They suggest a complexity, which well exceeds that of typical
polymeric structures. An in-depth understanding of this structure and the biology behind it may be
required for the optimal design of biomaterials. The use of biomaterials for tissue and organ
regeneration is called tissue engineering. Tissue engineers study other materials in addition to
polymers, and many of these materials, e.g. hydroxyapatite, will be crucial for the success of the
field.
The value of polymers is tremendous. The quantity of different materials that polymer chemists
can synthesize numbers well into the billions. The chemical structures of many natural and synthetic
polymers used in medicine are presented in Fig. 2a±c. Polymers with similar chemical characteristics
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
Fig. 2. (a) Chemical structures of some commonly used nondegradable polymers in tissue engineering. These polymers
include (1) polyethylene, (2) poly(vinylidene fluoride), (3) poly(tetrafluoroethylene), (4) poly(vinyl alcohol), (5)
poly(hydroxyalkanoate), (6) poly(ethylene terephthalate), (7) poly(butylene terephthalate), (8) poly(methyl methacrylate),
(9) poly(hydroxyethyl methacrylate) (10) poly(N-isopropylacrylamide), (11) poly(dimethyl siloxane), (12) polydioxanone,
and (13) polypyrrole. (b) Chemical structures of some commonly used degradable polymers in tissue engineering. These
polymers include (1) poly(glycolic acid), (2) poly(lactic acid), (3) poly(ethylene oxide), (4) poly(lactide-co-glycolide), (5)
poly(e-caprolactone), (6) polyanhydride, (7) polyphosphazene, (8) poly(ortho-ester), and (9) polyimide. (c) Chemical
structures of some commonly used biologically-derived polymers in tissue engineering. These polymers include (1)
alginate, (2) chondroitin-6-sulfate, (3) chitosan, (4) hyaluronan, (5) collagen, (6) polylysine, (7) dextran, and (8) heparin.
behave differently in certain situations. For example, polyethylene and ultrahigh molecular weight
polyethylene behave differently as orthopedic biomaterials for knee and hip replacement. Ultrahigh
molecular weight polyethylene exhibits reduced wearing and debris formation compared with
polyethylene. This broad spectrum of materials provides scientists and engineers with not only many
choices of existing polymers, but also the opportunity to design polymers better suited to the tissue
of interest. Although, for instance, skin is a relatively soft tissue, it must withstand large shear
stresses. Bone, on the other hand, is a relatively hard material with high compressive strength that
must support the weight of the body. One can imagine that a specific material would contain
properties more favorable for one application than for another. We certainly would not engineer a
brittle plastic for use as a soda bottle; likewise, we would not want to design one material for all
applications in medicine. In addition to the issues of mechanical properties of the materials, we must
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Fig. 2 (Continued ).
concern ourselves with the surface morphology, porosity, degradation, and chemistry of the
materials. As we will see below, these parameters all play a significant role in cell attachment,
proliferation, differentiation and secretion of the proper ratios of extracellular matrix molecules.
We also must consider biological parameters as we design new polymers. Regardless of other
inherent properties, the material must be biocompatible. This requirement means that the material
must not be immediately attacked or encapsulated by the body. As a result, there are many questions
that need to be answered in the search for new and improved materials. Will the body recognize the
polymer as grossly foreign and respond by walling off the foreign body? Will the material elicit an
immune response? Will cells that are able to regenerate the endothelium of the vasculature adhere,
proliferate and remain differentiated once they are in contact with the polymer? Do we need to
include bioactive factors that are covalently attached to the backbone of the polymeric material? On
the other hand, do we need to consider controlled release of a bioactive factor or factors to coax the
seeded cells to remain differentiated or secrete proper extracellular matrix molecules? All of these
questions are design parameters we must take into account as we develop new scaffolding for tissue
regeneration. We must also understand both the material properties of the extracellular matrix we are
attempting to temporarily replace and the biology of the system, so that we can induce the cells of
the body to replace the temporary scaffolding with the appropriate native tissue components.
To this end, material scientists and engineers have a great challenge ahead of them. We must
embrace not only traditional materials issues, but also biology. Knowledge in the life science fields
continues to develop rapidly. We have now sequenced the entire human genome. Stem cell research
has given us a new angle to approach tissue engineering. As a result, we may be able to harness the
potential of stem cells to develop tissues and organs as nature does during fetal development. This
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
Fig. 2 (Continued ).
approach, however, will take considerable effort by biologists, material scientists and engineers. We
will need to create materials that are biocompatible, support cell adhesion, growth, and
differentiation. These materials will need to temporarily replace mechanical function, and degrade
at rates appropriate to tissue regeneration. One example of the marriage between biology and
polymer synthesis involves the use of molecular biology to incorporate biological signals into the
backbone of polymers [2,3]. Meeting this challenge requires a formidable undertaking, but we live in
an exciting time during which there exists great potential for significantly advancing polymeric
biomaterial design.
The section on artificial skin begins by discussing epithelial, dermal, and full-thickness
replacements. It then goes on to discuss the effects of surface chemistry and morphology on cell
adhesion, morphology and metabolism. Finally, it discusses some novel polymers for use in skin
regeneration. The section describing polymeric materials for cartilage regeneration begins by
assessing nondegradable polymers. A discussion of degradable synthetic polymers follows, and a
review of degradable biopolymers concludes the section. The examination of polymeric materials
utilized for bone regeneration initially considers examples involving the control of surface chemistry
and morphology to guide cell adhesion, morphology, and differentiation. It then looks at research
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that attempts to adapt materials approved by the US Food and Drug Administration (FDA) before
concluding by exploring some novel materials with mechanical and erosion properties favorable for
bone regeneration. Following the discussion of bone regeneration, we present a review of materials
for vascular reconstruction; this section outlines the importance of improving artificial vascular
grafts. It explores endothelial cell seeding as a way of rendering current grafts more biocompatible.
Surface coatings, plasma treatments and biological sealants are also discussed. Polymeric materials
for use in peripheral nerve regeneration are discussed. We begin by looking at matrices for enhanced
neurite extension followed by an evaluation of nondegradable and degradable nerve guide tubes.
Tube-fill matrices are then addressed followed by brief descriptions of neurotrophic factor use and
micropatterning of polymers for axonal guidance. Finally, a brief section overviews materials in use
for liver regeneration including those for use in hepatocyte culture and hepatocyte and endothelial
cell co-culture. Microcarriers for hepatocyte culture are then discussed as possible devices for in
vitro acute liver failure treatments.
2. Skin regeneration
The development of new materials and the improvement of existing materials to promote skin
regeneration are large areas of research in polymeric biomaterials. Annually, several thousand people
need skin grafts due to dermal wounds. Trauma to the skin can be caused by heat, chemicals,
electricity, ultraviolet or nuclear energy, and can result in several degrees of skin damage. Trauma
can cause several degrees of injury. The least damaging traumas tend to wound only the epithelium,
the top layer of the skin. Wounded epithelium generally is healed by the body via re-epithelialization
and does not require skin grafting. More serious trauma can lead to partial or complete damage to
both dermal and subdermal tissues. Unfortunately, the body cannot heal dermal injuries properly.
Since skin forms a protective barrier around the body, damage to the dermis poses several immediate
threats including rapid, severe dehydration and infection. According the National Institute of General
Medical Sciences, 1.25 million burn-related injuries require medical attention annually in the US
alone. Roughly 50,000 of these patients require hospitalization and 25,000 are admitted to special
burn units. Infection leads to approximately 10,000 deaths. Fortunately, advances in wound care,
such as skin grafts, have increased the survival rate of burn victims and especially improved the
chances of survival for people with greater than 50% burn coverage. Today, people suffering from
surface area burns covering 90% of their body can survive, however, these patients often experience
some loss of function.
Skin is considered to be one of the more basic organs for regeneration. As detailed below,
some artificial skin systems are already in use. As shown in Fig. 3, skin does not consist of a simple
structure. The skin, with the integumentary system, forms the external coverage of the body. The
two main layers of the skin include the epidermis, composed of stratified squamous epithelium, and
the dermis, made up of dense connective tissue and fibroblasts. The epidermis is home to epithelial
cells and the matrix producing keratinocytes as well as the melanin producing melanocytes. The
hypodermis, comprised of a looser connective tissue, lies beneath the dermal layer. The skin
possesses sensory receptors, hair follicles for the production and growth of a hair, and two types of
sweat glands: the eccrine sweat glands, which primarily regulate body temperature, and apocrine
sweat glands [4]. As with other organs, the skin is perfused via capillaries and innervated by
the nervous system. Current artificial skins seek to recreate the dermis and the epidermis but do
not attempt to regenerate the intricate structures and functions of the hair follicles or the sweat
glands.
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Fig. 3. A pen drawing of the complex structure of skin adapted from [4]. The two main layers of the skin, as shown, are the
epidermis, composed of stratified squamous epithelium, and the dermis, made up of dense connective tissue. The
hypodermis, lies beneath dermis and is a looser connective tissue. The skin possesses sensory receptors, hair follicles for
the production and growth of hair and two types of sweat glands: the eccrine sweat glands, whose primary function body
temperature regulation and apocrine sweat glands.
Several approaches are taken in materials design for skin regeneration. These approaches vary
from development of full-thickness skin replacements to subdermal and epidermal replacements.
The body is capable of regenerating its own epidermis provided that the dermis remains intact.
However, severely burned victims and patients with diabetic ulcers, among others, often do not have
an intact dermis. Current solutions involve grafting of cadaver skin and xenografting; both
approaches are plagued with disease transmission and immune response problems. Scientists,
clinicians and engineers are working hard to develop therapies that regenerate, or help the body to
regenerate the dermis, epidermis or full-thickness skin. Although autografts are currently the
preferred materials, many patients do not have enough remaining healthy skin; in addition, the donor
site for the graft becomes another site of injury.
Ideally, graft materials would adhere to the wound site, be porous enough to allow diffusion of
wastes, nutrients and excess water, and prevent dehydration. Some water evaporation is essential,
however, since a complete lack of water removal would result in pooling, which in turn would cause
delamination of the grafted polymer. The material would also allow cell migration and have
mechanical properties similar to those of native skin. Lastly, the graft would need to promote dermal
and epidermal regeneration, to prevent bacterial infection, and degrade at an appropriate rate equal
but opposite to that of regenerating tissue. These parameters pose a formidable challenge for any
type of material.
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Current research focuses on three major areas, all of which start with rapid wound coverage.
The first type of graft consists of cultured epidermal cells with no dermal layer. The cells form
sheets, which can be applied to the burned area. The second approach contains only a dermal layer; it
is anticipated that once the dermal layer regenerates, an autograft of epidermal tissue can be applied.
The last approach involves a full-thickness bilayer containing both an epidermis and a dermis [5].
2.1. Epithelial replacements
Genzyme developed one approach to epithelial skin replacement and introduced it as the
product Epicel. Epicel consists of a polyurethane sheet upon which keratinocytes, the cells that
compose the living portion of the epidermis, can be cultured to form small colonies or fully
confluent sheets. The sheets can then be placed cell side down on top of wounds that have already
been treated with a dermal graft. These grafts promote the formation of a new epithelium [6].
Genzyme has solved the problem of thin, difficult to handle sheets of cells by growing the cells
directly on a polyurethane sheet that can then be applied to the wound bed. When the wound closes
or fully heals, the polyurethane sheet then falls off. In this case, the polymer is used as a support for
cell growth and also for application. In other cases, the polymer may actually act as a scaffold for
tissue regeneration as will be seen for dermal replacements. The latter application poses more
difficulties, as the polymer must be compatible for implantation into the body.
Other challenges, such as three-dimensional structure, must also be addressed with other tissues.
Epithelial tissue is one of the few tissues where two-dimensional cell culture may be preferable to
three-dimensional culture. The epithelium grows in a monolayer on top of a thin extracellular matrix
called the basal lamina. The basal lamina has distinct chemical and morphological structure, which
must be understood and mimicked in some way, but here we can likely omit the element of threedimensional design.
2.2. Dermal replacements
Several examples of dermal grafts exist. They range from completely degradable polymers to
layers of degradable and nondegradable materials. In one example of the combined use of degradable
and nondegradable scaffolds, the dermal layer is composed of type I collagen and chondroitin-6sulfate covered with a Silastic sheet. After application to the wound bed, the Silastic sheet acts as a
protective barrier against dehydration and infection while the collagen and chondroitin-6-sulfate act as
templates for tissue regeneration. Once tissue regeneration occurs, the Silastic sheet falls off.
Although the Silastic sheet improves handling and imparts some required properties, concerns still
exist regarding proper water evaporation. This construct is composed of both synthetic polymers and
biologically derived polymers. The synthetic polymers serve as temporary barriers and provide
mechanical stability for product application, and the biologically derived polymers are meant to act as
a template for tissue regeneration. In this instance, all cellular remodeling occurs from cells that
migrate into the new scaffolding from the wound site. As detailed below, other examples rely on
preseeded cells and begin the process of tissue regeneration prior to implantation into the body.
A striking success is the product Dermagraft, a dermal replacement, developed by Advanced
Tissue Sciences. Their approach focuses on growing human foreskin fibroblasts on poly(lactide-coglycolide) (PLGA) degradable scaffolding. In this case, the scaffolding is simply a support for the
fibroblasts, which secrete their own extracellular matrix molecules as well as growth factors and
other metabolically active molecules. These living skin substitutes are then cryopreserved. The
preserved tissue still contains living cells and newly derived biopolymers at the time of implant,
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however, the PLGA has degraded prior to implantation into the body. Recently, the product has been
implanted on pedal diabetic pressure ulcers. Results from clinical trials show that 54% of patients
treated once a week for 8 weeks with the metabolically active skin substitutes were completely
healed at the end of 12 weeks. In contrast, only 31.7% of patients treated with standard therapy were
healed at the end of 12 weeks [7]. Thus, Dermagraft offered a 20% improvement in the healing rate.
Despite the obvious success of this product, we are still a long way from developing materials that
will be effective in 100% of the cases. The PLGA scaffold allows the cells to regenerate an
extracellular matrix, but it is not clear that the PLGA supports proper matrix content, morphology or
assembly. Polymers more conducive to dermal tissue regeneration via chemical signals, growth
factor release, mechanical or morphological properties or some other biological signals may better
induce dermal regeneration and repair.
2.3. Full-thickness replacements
Organogenesis, Inc. has developed a type I bovine collagen matrix with allogenic human
keratinocytes and fibroblasts [8]. The cultured keratinocytes produce the epithelium while the
fibroblasts produce the dermal equivalent. This product is marketed under the name Human Skin
Equivalent (HSE). As a treatment for venous ulcers, HSE is effective 63% of the time while
traditional compression therapy is effective only 49% of the time. Again, this product improves the
life of an additional 14% of the people who suffer from venous ulcers, but it does not improve the
lives of all patients with this malady. In addition, since the scaffolding material consists of bovine
collagen, care must be taken to ensure the product is free of pathogens such as viruses and prions.
Approaches for a full-thickness skin replacement often try to mimic the porous structure and
mechanical properties of the extracellular matrix. The extracellular matrix (ECM) is composed of a
blend of macromolecules, which provide mechanical integrity to the tissues of the body and harbor
several bioactive signals and molecules. From a materials standpoint, the ECM is a highly
crosslinked, water-swollen network. The ECM takes on different characteristics depending on the
tissue. As a result, the density, pore size, crosslink density, and overall composition of ECM remain
tissue specific. For example, the extracellular matrix of cartilage is a highly hydrated, hydrogel-like
material while that of bone is a dense, hard ceramic. Yannas et al. approached the question of how
pore size and degradation rate affect wound healing of full-thickness skin injuries. They found that
there is a minimum and maximum pore size requirement (20±125 mm) in order to limit wound
contraction and scar tissue formation. This size distribution suggests that cells need not only enough
open space through which they can migrate, but also sufficient scaffolding on which they can
establish adhesion contacts and maintain mechanical stability [9].
Another approach to full-thickness dermal replacements uses different formulations of
Polyactive to develop an epidermis and a dermis. Since it is a segmented block copolymer
composed of alternating soft poly(ethylene glycol-terephthalate) and hard poly(butylene terephthalate) segments, Polyactive is biocompatible and degradable. van Dorp et al. made Polyactive
matrices with varying molecular weight (MW) polyethylene glycol (PEG) added to them. A dense
over-layer (10±30 mm thick) was cast on a porous (70±100 mm) under-layer (70±200 mm thick). They
found that MW 300 PEG promoted the best cell adhesion. Higher MW PEG increased water content
and decreased both cell adhesion and proliferation. A fully differentiated epidermis was formed,
however, the material did not allow the formation of either a basal lamina or anchoring zone. The
dense top layer of Polyactive, although necessary as a protective barrier, may hinder the interaction
between the newly forming epidermis and the fibroblasts below by preventing molecular and cellular
migration [10]. Polymers that provide a protective barrier, prevent dehydration and infection, and
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maintain communication between the epidermis and the dermis would greatly improve skin
regeneration.
A novel full-thickness wound healing approach involves the use of PLGA microspheres.
LaFrance and Armstrong cultured dermal fibroblasts and keratinocytes on the surface of PLGA
microspheres (73±140 mm in diameter). This approach contains the following advantages: the
spheres are easier to handle than thin sheets of cells, they can be injected more deeply into the
wound, and the spheres can be designed to release bioactive factors as they degrade [11]. Using the
porcine model, they injected microspheres coated with fibroblasts and epithelial cells into fullthickness wounds and found that coated spheres promoted more rapid and more natural wound
healing than uncoated microspheres. One of the concerns with all biomaterials involves the method
and ease of surgical implantation. These microspheres provide several advantages over other types of
materials since they allow easy handling, do not need to be sutured to the wound site, and do not
have to conform to the wound bed. One major disadvantage of this material, however, is that the
microspheres lack the ability to control fluid loss from the wound surface and to provide a barrier to
infection.
2.4. New approaches to materials design and surface chemistry
It is widely known that the surface chemistry of a material plays a significant role in cell
adhesion to that surface. This phenomenon is likely regulated intermediately by plasma protein
adsorption to the surface. Faced with the challenge of designing materials with specific, biological
interactions, material scientists, biologists and clinicians continue to try to understand the
relationship between surface chemistry and cell adhesion and spreading. To this end, work has been
performed with plasma treated surfaces and functionalized self-assembled monolayers (SAMs)
cultured with keratinocytes. SAMs with hydrophobic end groups or carboxylic end groups and acidfunctionalized or hydrophobic surfaces were seeded with keratinocytes. The cells seeded on acid
functionalized surfaces and SAMs adhered, proliferated and formed confluent monolayers of
keratinocytes over time. The behavior of the cells on acid-functionalized surfaces compared
favorably to keratinocyte behavior on the gold standard type I collagen. Cells did not adhere well to
the hydrophobic SAMs and surfaces [12]. This result indicates that acidic surfaces are more adept at
promoting keratinocyte adhesion and proliferation than hydrophobic surfaces. Another surface
modification approach involves glow-discharge modification. PisËkin and AtacË modified the surface
of activated carbon cloth with either 2-hydroxyethylmethacrylate, dimethyl-aminoethyl methacrylate
or hydroxymethyl disiloxane. They then looked at the material's ability to behave as a protective
barrier to microbial species while maintaining properties that allow toxin and oxygen absorption as
well as water permeability [13]. They found that only the poly(hydroxymethyl disiloxane) decreased
the ability of the carbon sheet to absorb toxins. All of the other surface modified carbon sheets
maintained this ability as well as other favorable characteristics. These surface modifications may
make the activated carbon sheets more amenable to cell culture. As a result, information gained in
surface modification studies may be used to guide us in the appropriate design parameters for new
materials.
Surface chemistry clearly influences a cell's ability to attach and spread. Another factor that
affects cellular behavior is the presence of serum proteins on the surface. Cells inherently adhere to
the proteins via receptors embedded within cellular membranes. When deposited onto a surface,
these serum proteins alter the adhesive characteristics of the surface. Tjia et al. have pre-adsorbed
serum proteins to PLGA to make the polymer more conducive to cell adhesion and migration.
Adsorption of collagen or fibronectin was performed prior to seeding with epithelial cells.
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Exogenous collagen supported greater adhesion and migration and allowed cells to secrete their own
matrix molecules, e.g. fibronectin, which in turn play a role in regulating adhesion and migration.
These results led the group to suggest a paradigm involving induction of cell migration on synthetic
polymers coupled with programmed synthetic matrix remodeling and cell-derived matrix remodeling
[14]. This paradigm should not be taken lightly. It is clear that success in tissue regeneration will
require the design of scaffolds that guide cells through at least the initial stages of development.
Additionally, we emphasize that the native extracellular matrix molecules are designed to promote
cell adhesion, proliferation and remodeling of the extracellular matrix. The key for successful tissue
engineering is to induce the proper remodeling. To date, however, approaches involving extracellular
matrix molecules alone have failed. Nevertheless, we must not forget to take lessons from nature as
we begin to derive new and improved materials for skin and other tissue regeneration.
As alluded to by the work done on full-thickness skin replacement scaffolds generated by
Yannas et al. there exists a notion that the three-dimensional topological structure of materials will
play a large role in the quality of skin generated using new biomaterials designed for artificial skin
[9]. One elegant example is work by Pins et al. They used laser ablation of polyimide to create a
master pattern with varying channel depths and widths. A negative image was created with
poly(dimethylsiloxane) (PDMS), which was in turn used to create a positive image with either
gelatin or a collagen type I Ð glycosaminoglycan mixture. They found a correlation between surface
features and artificial skin structure. Stratification of the artificial epidermis, a natural and necessary
feature of skin, was seen to increase with channel depth, and channels 300 mm deep showed the
greatest stratification [15]. Stratification provides skin with one mechanism for withstanding
everyday shear forces. Consequently, skin located in regions of high shear, such as the skin on the
feet and palms of the hands, shows the greatest degree of stratification.
Not only are topographical features important when designing materials for full-thickness skin
wounds, porosity of the material must also be considered. Porosities of 95±97% allow for structural
stability, provide sufficient surface area to which cells can adhere, and contain enough free space for
adequate nutrient diffusion and cell migration. In one study, porous structures were formed with the
biopolymers chitosin and collagen. Co-cultures of human umbilical vein endothelial cells (HUVECs),
fibroblasts, and keratinocytes were then evaluated. The inclusion of HUVECs attempted to induce the
formation of a capillary network within the regenerating tissue. Although the authors saw capillarylike structures within the material when HUVECs were seeded with fibroblasts, HUVECs seeded
alone tended to die. Thus, the fibroblasts likely secrete factors crucial for HUVEC survival. For coculture, the authors also observed the formation of a dermal-like layer and basal layer [16]. As a
result, the co-culture of fibroblasts with HUVECs appears important for skin regeneration.
One thought as to the reason for the failure of full-thickness skin approaches is the lack of
vascularization of the new tissues. New materials must account for the vascular needs of both new
and old tissues to ensure adequate waste removal and nutrient availability. All cells must be within a
millimeter of a blood supply or they will die due to either the accumulation of waste products or the
lack of nutrients. This constraint imparts another great challenge when designing materials for tissue
regeneration. We must not only consider proper parameters for specific cell types, such as
keratinocytes, but we must also think about how to connect the body's existing blood perfusion
system to the developing tissue.
2.5. Hydrogels
Other approaches used to develop new materials for use as skin substitutes involve the
optimization of properties of existing hydrogel materials. Hydrogels provide a moist wound covering
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conducive to healing, and also protect the wound from infection. A disadvantage of many hydrogel
networks, however, is their lack of strong mechanical properties. A hydrogel must be able to
withstand the stresses and strains of everyday tissue function while the new tissue generates. Young
et al. improved the tensile strength and break point of poly(2-hydroxyethyl methacrylate) (pHEMA).
Their first attempts used spandex fibers, gauze and low lint wipes, and these fiber-reinforced gels
experienced a dramatic increase in mechanical properties. By itself, pHEMA, with an initial water
content of 33%, had a Young's modulus of 1.87 MN m 2. In contrast, the moduli of the fiberreinforced gels were 6.89, 5.31 and 51.92 MN m 2 for the low lint wipe, Spandex and gauze
reinforced gels, respectively [17]. These increases resulted in gels that can withstand greater forces
before tearing, which is a crucial property for any skin substitute. Dvorankova et al. cultured
autologous keratinocytes on pHEMA gels in vitro. The gels were then placed on full-thickness
wounds cell side down. Histology of the tissue replacement 16 days after application showed multilayered, fully differentiated epidermis in the wound bed. In this application, the hydrogel acted as a
scaffold for tissue regeneration and behaved as a protective barrier [18].
When considering a material for medical use, the toxicity of both the polymer itself and the
degradation products must be considered. A polymer would be of little use as a biomaterial if it killed
either the tissue with which it came into contact or any cells that tried to invade the material.
Consequently, the toxicity of materials, especially those resulting from modifications of existing
materials, requires evaluation. One example includes dextran hydrogels, which have been studied as
skin substitutes [19]. Dextran itself is not toxic, but some of the methods used for crosslinking the
polymer result in toxic byproducts. The toxicity of dialdehyde crosslinked hydrogels was compared in
vitro with the toxicity of Tagegerm and OpSite, two semi-occlusive polyurethane dressings, and
DuoDERM, a hydrocolloid wound healing dressing. The toxicity of the modified dextran was found to
be low in fibroblast and endothelial cell cultures and acceptable in keratinocyte cultures. Overall, the
dextran gels compared well with their commercial counterparts, however, no cell culture or clinical
data was reported. Since it is a naturally occurring microbial polysaccharide, is biocompatible, and
inhibits nonspecific interactions, dextran may find many applications as a polymeric biomaterial.
Several factors must be taken into account when developing new materials for skin regeneration.
Toxicity of the materials and of degradation byproducts to cells in culture and to the body is of great
concern. The materials must act as protective barriers while allowing water and oxygen diffusion.
Too little water diffusion can result in swollen, fluid-filled wounds that cannot heal properly; too
much water diffusion will result in severe dehydration. The materials must both support and
influence tissue regeneration; the body cannot replace the dermal tissue without help. The researcher
must not only take care to design the appropriate material for the tissue of interest, but also remain
aware of crucial design parameters that are conducive to adequate blood perfusion. Although great
strides have been taken to improve the quality of care of burn victims and patients with diabetic
ulcers, the current materials are not able to induce healing in close to 100% of the cases. Even in
successful applications, the cosmetic results are not always favorable. The community needs to place
a significant effort in the development of a new class of materials for skin regeneration. In order to
accomplish this daunting task, knowledge from materials science and engineering must be combined
with ideas from molecular and cellular biology, developmental biology, and medicine.
3. Cartilage regeneration
Over the past decade, one of the most active areas of polymeric-biomaterials research has
involved cartilage repair. This activity results in large part because cartilage defects cannot
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adequately heal themselves. As a result, millions of people require treatment to repair damaged
cartilage, and much of this damage involves cartilage within the knee joint. In fact, the American
Academy of Orthopedic Surgeons estimates that more than six million people experience some form
of knee-related problem within the US each year. Thus, any treatment that would successfully aid in
the regeneration of cartilage would greatly benefit the lives of patients either recovering from injury
or suffering from osteoarthritis. Due to its biochemical composition and isolation from neural and
circulatory pathways, cartilage can seem like a deceptively simple tissue to replace or repair. In
reality, however, cartilage regeneration continues to challenge state-of-the-art three-dimensional
scaffold technology.
Although the human body contains three types of cartilage (elastic, fibrous, and hyaline
cartilage), most current research has focused on hyaline cartilage. Hyaline cartilage exists as the
predominant form of cartilage in the body and coats the surfaces of all articulating joints. As a
result, joint cartilage is referred to often as articular cartilage. Fig. 4a depicts the histological and
collagen fiber structure of this type of cartilage. The composition of articular cartilage includes
chondrocytes and several extracellular matrix molecules. Chondrocytes only make up about 1% of
the total volume of cartilage; however, they continually remodel and organize their surrounding
matrix. The extracellular matrix contains two major types of molecules: collagen and proteoglycans.
Over 90% of the collagen found in articular cartilage consists of type II collagen. This matrix protein
provides much of the mechanical integrity of cartilage and comprises 60% of the total dry weight
of cartilage. The right panel of Fig. 4a illustrates the complex organization of collagen fibers
within cartilage. In the tangential zone, the fibers run more parallel to the joint surface. The fibers
begin to weave in a more oblique fashion in zone II and ultimately are arranged in a vertical pattern
orthogonal to the superficial layer. On a cellular level, chondrocyte density remains somewhat
sparse superficially and then tends to increase towards zone III. The arrangement of chondrocytes
in zone IV appears columnar, and the cells within this region begin to calcify and form the boundary
between cartilage and subchondral bone. Proteoglycans, the other major component of cartilaginous
matrix, compose 25±35% of cartilage dry weight. These complex molecules are formed by glycosaminoglycan polysaccharides attached to a central protein. Within cartilage, the major
glycosaminoglycans include hyaluronic acid, chondroitin sulfate, keratan sulfate, and dermatan
sulfate. Large aggregating proteoglycans are often called aggrecans and contain large amounts of
chondroitin sulfate and keratan sulfate. Hyaluronic acid, or hyaluronan, is somewhat unique among
the glycosaminoglycans since it exists as very long chains with large molecular weights (often more
than 1 106 Da) and, as seen in Fig. 4b, can associate with proteoglycan monomers via link
proteins. The resulting macromolecule aggregates can greatly influence compressive properties and
hydration [4,20,21].
Consisting of sugars modified with negatively charged carboxylate or sulfate groups,
glycosaminoglycans (GAGs) are highly polar. When many of these GAGs associate with a core
protein to form a proteoglycan, the polar nature of the proteoglycans allows the molecules to interact
strongly with water. Since cartilage consists of 60±80% by weight water, many of the viscoelastic
and mechanical properties exhibited by cartilage result from this dynamic relationship between
cartilaginous GAGs and water [4,20,21]. In fact, cartilage found within some regions of the body,
such as the articular cartilage of the hip, can withstand compressive stresses up to 20 MPa [22].
Thus, any replacement treatment needs to be sensitive to and supportive of loading. Compared with
other tissues, articular cartilage contains no vasculature, nerves, or lymphatic vessels. Thus, cartilage
must remain relatively thin (with a thickness on the order of about 2 mm) in order for sufficient
nutrient and waste diffusion. Since this review will only focus on cartilage repair strategies
incorporating polymeric devices, additional information regarding other therapies and considerations
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Fig. 4. (a) A pen drawing of the structure of articular cartilage adapted from that given by Frank Netter [263]. The
composition of articular cartilage includes chondrocytes and several extracellular matrix molecules. Chondrocytes only
make up about 1% of the total volume of cartilage. The extracellular matrix contains two major types of molecules:
collagen and proteoglycans. Over 90% of the collagen found in articular cartilage consists of type II collagen. This matrix
protein provides much of the mechanical integrity of cartilage and comprises 60% of the total dry weight of cartilage. It is
organized into fibers that compose four zones: Zone I or tangential, Zone II or oblique, Zone III or vertical, and Zone IV.
Below Zone IV lie the end plate and the trabecular bone. Proteoglycans, on the other hand, compose 25±35% of cartilage
dry weight. (b) A pen drawing of an aggregated proteoglycan adapted from that given by Frank Netter [263]. Large
aggregating proteoglycans are often called aggrecans and contain large amounts of chondroitin sulfate and keratan sulfate.
Chondroitin sulfate and keratan sulfate, called glycosaminoglycans (GAGs), are modified with negatively charged sulfate
groups and are highly polar. When many of these GAGs associate with a core protein to form a proteoglycan, the polar
nature of the proteoglycans allows the molecules to interact strongly with water.
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
(such as cell suspension, bioreactor design, and drug delivery) can be found in several well-written
articles [22±27].
Polymers involved in cartilage repair fall into two major categories: nondegradable and
degradable. Some nondegradable materials include poly(vinyl alcohol), poly(hydroxyethyl
methacrylate), and poly(N-isopropylacrylamide). Although many of these synthetic polymers are
designed to withstand the high mechanical stresses experienced within articulating joints and thereby
replace damaged cartilage, some nondegradable materials attempt to support tissue formation.
Degradable polymers consist of either synthetic materials, such as poly(glycolic acid) or poly(lactic
acid), or naturally-derived materials, e.g. alginate or collagen. Many types of nondegradable and
degradable polymers have been investigated as therapies for articular cartilage repair; however,
current trends have favored degradable materials. One reason for the appeal of degradable polymers
is that they can form a temporary scaffold for mechanical and biochemical support. As new cartilage
begins to form within the defect site, these materials will then degrade and leave behind a
regenerated tissue. Some of the important parameters for designing polymeric devices for cartilage
repair involve providing sufficient mechanical properties for joint loading, incorporating proper
biochemical environments and porosities, accommodating cellular infiltration, maintaining
chondrocyte phenotype, allowing desired degradation rates, and providing suitable mechanisms
for surgical handling and implantation. Additionally, successful polymer candidates should carefully
avoid any potential problems related to compliance mismatch between the device and the
surrounding native tissue. As we examine research involving cartilage repair, we will first review
nondegradable polymer matrices and then focus on degradable systems.
3.1. Nondegradable polymers
3.1.1. Poly(vinyl alcohol)
When designing a nondegradable polymeric replacement for cartilage, several parameters need
to be addressed. Some of these design considerations include providing adequate lubrication and
mechanical properties, having good biocompatibility, allowing the material to adhere well to
surrounding tissue, and resisting wear and fatigue. With the exception of mechanical properties,
many synthetic hydrogels can meet these prerequisites [28]. Poly(vinyl alcohol) (PVA) was one of
the first synthetic polymers tested as an artificial cartilage. PVA is a hydrogel and contains a water
content similar to that of cartilage; however, like many hydrogels, it lacks sufficient mechanical
stability for use as a cartilage replacement in its unmodified form [29]. PVA does offer several
beneficial properties as an artificial cartilage. It is relatively biocompatible, it can swell to
accommodate a large water content, it can be sterilized, and it can be molded into desired shapes
[30]. Several groups have attempted to modify the process of synthesizing PVA in order to create a
hydrogel with improved mechanical properties [28±31]. The result of one of these efforts led to a
commercially available PVA hydrogel known as SalubriaTM (Salumedica, Atlanta, GA). SalubriaTM
is created by completing a series of freeze/thaw cycles with the PVA polymers and 0.9% saline
solution. By altering the ratio of PVA and water, the molecular weight of PVA, and the quantity and
duration of freeze/thaw cycles, the material can be synthesized with varying mechanical attributes.
Recently, Stammen et al. noted that any nondegradable material intended for use as a cartilage
replacement should be able to withstand enormous loading consistently over millions of cycles.
Consequently, they examined the shear and compressive characteristics of SalubriaTM. They selected
two formulations of PVA: one with 75% water and the other containing 80% water. These gels were
then molded into cylinders measuring 6 mm in height and 6 mm in diameter. By analyzing stress±
strain relationships, shear modulus, and compressive modulus, the authors observed that the
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compressive modulus increased with greater strain levels, with the PVA containing 80% water
yielded slightly higher moduli than the gels with lower water content. At 30% strain, the
compressive tangent modulus measured about 2.5 MPa for both gel types. When the strain level
advanced to 60%, however, the moduli of the gels reached values between 15 and 21 MPa. The shear
modulus results did not differ between gel type. Tests measuring the compressive failure of the gels
showed that the 75% water PVA had an ultimate stress near 2.1 MPa at 60% strain, while the 80%
water PVA did not perform as well and failed at a 1.4 MPa stress and 45% strain. Although the
hydrogel formulation with less water content resulted in a stronger material, all of the recorded
parameters for both gels fell within the range of human articular cartilage. Further testing is
warranted, however, to examine how the SalubriaTM biomaterial fatigues and wears. The authors also
briefly mention that growth factors or other bioactive materials could be incorporated into the PVA
hydrogel system in order to promote favorable biological responses in vivo [30].
Oka et al. have also worked with developing an artificial articular cartilage based on PVA. They
created PVA hydrogels with an average degree of polymerization of 7000 and a molecular weight of
308,000 Da. After drying the polymer in a vacuum, the PVA was rehydrated until it contained 20%
water. When the polymer was subjected to an impact load test, the maximum recorded stress reached
4.5 MPa and lasted for several milliseconds. When the same experiment was performed with
subchondral bone with articular cartilage, the peak stress was higher (7.5 MPa) but the duration of
the stress was much lower. Thus, the PVA had better damping properties than the natural tissue. To
test PVA biocompatibility, cylindrical hydrogels (4 mm in height and 4 mm in diameter) were
prepared and implanted into osteochondral defects in rabbits. The animals were then observed over a
period of 52 weeks. Upon examination, the PVA hydrogels did not elicit an inflammatory response
and only induced a mild tissue reaction. To determine whether or not wear particles from PVA
influence biocompatibility of the material, small pieces of PVA (ranging from 50 to 300 m in size)
were implanted within the knee of rats. After 2 months, the PVA fragments showed slight
inflammatory reaction. In order to improve some of the wear properties of PVA, gels with higher
molecular weights were created and exposed to gamma radiation. Improvements to wear were also
made by introducing hyaluronic acid into the gel. Even with these improvements, however, the
authors concluded that PVA would not be suitable for total joint replacement but may prove
beneficial for smaller scale cartilage replacement or joint resurfacing [28].
3.1.2. Polyacrylates
Another attempt to create an artificial articular cartilage involved poly(2-hydroxyethyl
methacrylate) (pHEMA). Malmonge and Arruda thermally copolymerized 2-hydroxyethyl
methacrylate with acrylic acid in different ratios. The resulting hydrogels were then washed and
evaluated mechanically. Gels containing more acrylic acid had higher negative charge densities and
as a result, experienced larger osmotic pressures. A plot of complex modulus versus acrylic acid
concentration revealed a somewhat flat but linear relationship with higher amounts of acrylic acid
contributing, generally, to a higher complex modulus. For example, a gel without acrylic acid had a
modulus of around 1.2 MPa, and the modulus for a gel containing 15% acrylic acid measured
1.3 MPa [32]. In another study, Malmonge et al. examined the mechanical behavior of newly formed
cartilage tissue within a defect treated with a pHEMA implant. They created cartilage defects in
Wistar rats and mechanically tested both the polymer device within the defect and the surrounding
normal articular cartilage for up to 16 weeks post-implantation. Throughout the experiment, the
pHEMA material remained attached to the adjacent tissues and did not induce an inflammatory
response. All of the defects left empty as controls were filled with fibrous tissue. By comparing the
properties of the hydrogel with those of the repair tissue within untreated defects, the authors found
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that the mechanical characteristics of the pHEMA implant differed significantly. Throughout the
duration of the study, nontreated defects were much stiffer than native cartilage, whereas the
pHEMA structure proved more compliant than adjacent tissue [33].
Attempting to improve the quality of neocartilage, Sawtell et al. proposed a poly(ethyl
methacrylate)/tetrahydrofurfuryl methacrylate (PEMA/THFMA) system. Although they do not
biodegrade, these polymer networks provide beneficial properties for mechanical uniformity and
implantation. Some of these properties include room temperature, low exothermic polymerization
and minimal shrinkage during polymerization. Initial in vitro experiments examined both
chondrocytes and agarose-suspended chondrocytes seeded on the PEMA/THFMA material. Over
14 days, the cells cultured on PEMA/THFMA showed an increase in glycosaminoglycan synthesis, a
slow rate of cellular proliferation, and a rounded morphology [34]. In vivo rabbit studies compared
PEMA/THFMA with poly(methyl methacrylate)/methyl methacrylate (PMMA/MMA). Both of these
polymer systems polymerized in situ within a 4.5 mm subchondral defect. Even though the materials
did not include a cell suspension, cartilage formation still occurred after 6 weeks. The defects with
PMMA/MMA contained mostly fibrous cartilage, had little proteoglycan production, and were
devoid of zonal differentiation. In fact, 77% of these defects were not fully resurfaced. In contrast,
80% of the PEMA/THFMA defects experienced complete resurfacing with hyaline-like cartilage,
good distribution of collagen type II and proteoglycans, and zonal organization [35]. More recent
work returned to an in vitro analysis of the polymer system. Over a period of 28 days, the PEMA/
THFMA material was compared with Thermanox, a polyethylene terephthalate film. Results
indicated that chondrocytes seeded on the Thermanox control experienced a greater degree of
proliferation and higher glycosaminoglycan production than those seeded on PEMA/THFMA did. In
addition, the ratio of glycosaminoglycan content remaining in the matrix to that diffusing into the
media was higher for the Thermanox culture than the PEMA/THFMA culture. Neither material
induced the chondrocytes to express a full hyaline phenotype since both collagen type I and type II
were present within the cultured matrix [36]. Although the polymeric properties of PEMA/THFMA
are conducive to implantation and in situ polymerization, the performance of the PEMA/THFMA
scaffold with respect to cartilage formation seems similar to the other polymer systems described in
this review.
3.1.3. Poly(N-isopropylacrylamide)
Stile et al. investigated the use of poly(N-isopropylacrylamide) (pNIPAAm) as an injectable
hydrogel for cartilage tissue applications. Since it has a lower critical solution temperature (LCST)
of 328C, an aqueous solution of pNIPAAm can be injected in situ directly into a cartilage defect. As
the polymer solution temperature increases to 378C, a reversible hydrogel will form. The authors
created loosely crosslinked polymers of pNIPAAm and a copolymer of pNIPAAm and acrylic acid
(pNIPAAm-co-AAc). These polymers had LCSTs of around 32 and 348C, respectively. Both of these
polymers had very similar water contents at room temperature (around 93%). At 378C, however,
pNIPAAm expelled almost half of its water, but after 6 days in PBS, the water content increased
from 44 to 68%. The pNIPAAm-co-AAc did not experience the same dramatic water loss. In fact,
the initial water content at physiological temperature actually increased slightly but decreased to
about 74% after 6 days in PBS. These values become important in the context of in situ gelation
since a polymeric plug needs to adequately fill the cartilage defect void. If too much water separates
from the hydrogel, the polymer might not yield a tight or secure fit and consequently, would decrease
the potential for tissue regeneration. In vitro experiments with bovine chondrocytes demonstrated
that both polymers supported cell viability and allowed cells to produce extracellular matrix
molecules. Although pNIPAAm-based hydrogels are typically considered nondegradable polymers,
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this group has successfully synthesized biodegradable pNIPAAm using ethylene glycol, lactide, and
e-caprolactone crosslinkers [37]. This system requires further investigation, however, these initial
results showed that hydrogels derived from pNIPAAm can provide a biocompatible matrix for
applications such as cartilage repair.
3.1.4. Polyethylene
Another class of nondegradable materials includes polyethylene-based polymers. By melting
low-density linear polyethylene, Hasegawa et al. were able to coat a three-dimensional fabric
consisting of 400 mm diameter ultrahigh molecular weight polyethylene fibers. The resulting
polymer had a compressive behavior similar to that of natural cartilage. After fabricating a
5 mm 10 mm 5 mm three-dimensional structure, the majority of the surface area was coated
with hydroxyapatite (HA) powder. The authors implanted these fabrics into defects within rabbit
articular cartilage; one knee received an implant while the other remained empty. Hyaline-like
cartilage started forming around the implants after 2 weeks, partially covered the articular surface
after 8 weeks, and remained at the 8-week level after 24 weeks. In contrast, either fibrous tissue or
bone filled the empty control defects. When tested mechanically, the polyethylene fabrics implanted
for 12 weeks had a compressive stress±strain relationship very similar to that of native tissue. Very
little cartilage infiltrated the pores of the construct over 24 weeks; almost 50% of the tissue within
the fabric was fibrous in nature. After 24 weeks, bone had grown around more than 70% of the
area surrounding the polymer. Regarding the articular surface, the polyethylene material did induce
a significantly greater amount of cartilage formation; however, only 15% of the fabric surface
was actually covered by cartilage from weeks 8 to 24 [38]. A material designed to have a layer of
HA on a scaffold for cartilage regeneration may, when juxtaposed to bone, induce the integration
of the neocartilage with the adjacent bone; this material would also be useful for osteochondral
defects.
3.2. Synthetic degradable polymers
3.2.1. Polyesters
One of the earliest and most widely investigated groups of materials studied for cartilage repair
includes the synthetic, degradable poly(a-hydroxy acids). The two most well known of these
polymeric biomaterials include poly(glycolic acid) (PGA) and poly(lactic acid) (PLA). These
polymers remain popular for a variety of reasons including the fact that both of these materials have
properties that allow hydrolytic degradation through de-esterification. Once degraded, the
monomeric components of each polymer are removed by natural pathways; glycolic acid can be
converted to other metabolites or eliminated by other mechanisms, and lactic acid can be cleared
through the tricarboxylic acid cycle [39]. Thus, the body already contains highly regulated
mechanisms for completely removing residual traces of the polymers. Due to these properties, PGA
and PLA have been used in products such as degradable sutures for several decades and both have
received approval for use by the US Food and Drug Administration (FDA). Some studies have even
indicated that these materials may not actively provoke an immune response since they lack peptide
bonds [39]. One potential concern, however, arises from the local pH changes that can occur upon
degradation due to the acidic nature of both the glycolide and lactide monomers. With its additional
methyl group, PLA is much more hydrophobic than the highly crystalline PGA. As a result, PLA has
a much slower degradation rate. In fact, PGA can degrade in a few weeks [40]; whereas PLA can
remain stable for over 1 year [41]. These mechanical properties can be tailored to a degree by
creating copolymers of PGA and PLA. However, the characteristics of the resulting copolymer are
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
not always a linear combination of the properties of the pure polymers. For example, materials made
with equal ratios of PGA and PLA degrade more quickly than PGA scaffolds. Due to its chiral
structure, PLA can exist in two forms: D-PLA and L-PLA. However, L-PLA has experienced more
widespread use as a biomaterial since L-lactic acid occurs naturally. The racemic mixture of D,L-PLA
has amorphous characteristics, whereas a pure polymer of either of the stereoisomers tends to be
more crystalline [42].
Early work with polyester materials held much promise for tissue engineered cartilage. Cima
et al. developed the first protocols for seeding chondrocytes onto degradable polymer meshes. These
meshes would act as scaffolding to support cell attachment and allow the chondrocytes to begin
synthesizing cartilaginous extracellular matrix molecules. Initial in vitro work showed that bovine
chondrocytes adhered to Vicryl fibers (90:10 PLGA coated with 70:30 PLGA and calcium stearate,
Ethicon), retained a rounded morphology, and produced glycosaminoglycans [43]. Cell-polymer
constructs were then implanted subcutaneously in nude mice for corresponding in vivo work. These
studies revealed that after 9 months, more than 90% of the constructs contained cartilage similar to
human fetal cartilage complete with sulfated glycosaminoglycans and type II collagen. In contrast,
no cartilage was formed in any of the control matrices or by chondrocytes not supported by a
polymer [44]. To further investigate the potential of polyesters in repairing cartilage defects, Vacanti
et al. performed an expanded in vivo study on rabbits. Chondrocytes were extracted from rabbit
hyaline cartilage, seeded onto 100 mm thick, nonwoven PGA meshes with 75±100 mm interfiber
spacing, and allowed to incubate in tissue culture conditions for 1 week. Cartilage defects
(5 mm 8 mm 2 mm) were then created in the knees of rabbits, and the cell-polymer structures
were implanted in the defect site. After 7 weeks, neocartilage was observed in almost all of the
experimental grafts. Within the control groups, only small amounts of fibrocartilage could be
detected. Therefore, the thin PGA sheet was successful as a matrix for allowing chondrocytes an
opportunity to resurface a cartilage defect in an articulating joint [45].
Freed et al. greatly expanded the role of polyester materials in neocartilage research. Noting that
the previous work used two-dimensional materials, they characterized the type of cartilage formed
on three-dimensional scaffolds in order to investigate joint resurfacing of larger defects. Specifically,
they investigated fibrous poly(glycolic acid) (PGA) felts and porous poly(L-lactic acid) (PLLA)
foams measuring 1 cm 1 cm 0:15 cm and 1 cm 0:5 cm 0:3 cm, respectively. The nonwoven
PGA materials had 14 mm diameter fibers with a density of 8.5 mg/cm2, and PLLA was formed from
chloroform solvent casting using salt particles ranging from 100 to 500 mm in diameter. These
scaffolds were then seeded with bovine chondrocytes, cultured in vitro for 2 weeks, and inserted
subcutaneously in the backs of athymic mice. In vitro evaluation indicated cellular adhesion reached
54% and that PGA supported chondrocyte morphology and phenotype much better than PLLA. After
5 weeks in vivo, PGA felts without cells had degraded completely, but nonseeded PLLA was present
up to 6 months after implantation. Whereas chondrocytes injected without a polymer matrix formed
small, nonhomogenous cartilage and polymers without cells did not induce cartilage formation, all
cell-polymer constructs developed cartilage in the same dimensions as the original scaffold.
Neocartilage formed from the scaffolds was rich in sulfated glycosaminoglycans and contained type
I and type II collagen. Although PGA constructs experienced a cellular growth rate twice that of
PLLA and seemed to work much better for short-term applications, both scaffolds showed
remarkable promise as long-term in vivo materials [46]. One possible explanation for the slower
growth of chondrocytes on PLLA is that the solvent extraction technique may have inhibited
favorable interactions between the cells and the materials. As time passed, more of the residual
solvent or salts could diffuse from the site and allow PLLA to become a much more favorable
environment for chondrogenesis [47].
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Due to the faster degradation rate and success of PGA relative to PLLA in many of the shortterm in vitro studies, Freed et al. continued their work by applying large-scale polymer extrusion
techniques to create uniform PGA felts that were used to grow large pieces of cartilage in vitro. To
create optimal spacing for cellular attachment and proliferation, these scaffolds consisted of a
meshwork of 13 mm diameter fibers and a porosity greater than 95% [48]. Other work has confirmed
that PGA felts with 14±15 mm fiber diameters, high porosity, and 75±100 mm interfiber distances
can successfully support the chondrocyte phenotype in vitro and in vivo [49]. This processing
technique, which allows material reproducibility and structural control, provides the possibility to
gain greater insight into how the parameters of scaffold design influence cellular behavior and tissue
formation. These experiments also showed that PGA felts seeded with chondrocytes degraded at a
rate complimentary to that of neocartilage formation [48]. After succeeding in growing relatively
large pieces of cartilage in vitro, Freed et al. had to evaluate the therapy within an articulating joint.
Using a rabbit model, full-thickness cartilage defects measuring 3 mm in diameter and 1±2 mm deep
were created in each animal. These holes were then filled by either empty PGA disks or PGAchondrocyte constructs that had been growing in vitro for several weeks. After 6 months in vivo, the
defects with empty PGA controls contained nonhomogenous regions of extracellular matrix
characteristic of both fibrocartilage and hyaline cartilage. The PGA-chondrocyte grafts, however,
allowed the formation of smooth, hyaline cartilage complete with physiologically correct
chondrocyte orientation and reconstruction of the subchondral plate. Although only a pilot study,
these results showed that PGA can support cartilage formation in a region of increased mechanical
loading [50]. Another recent study, also using a rabbit model, supported these results. Schreiber et al.
noticed that neocartilage grown in vitro on PGA scaffolds induced significantly more hyaline
cartilage after 9 months in vivo than untreated defects. In addition, the quality of repair tissue for the
defects receiving neocartilage plugs remained somewhat stable for 24 months post-implantation
[51].
Several of the above studies have concluded that chondrocytes passaged several times in tissue
culture did not experience compromised activity when seeded in PGA felts [48,50,51]. In fact, some
experiments have shown that glycosaminoglycan and collagen production can increase [50]. By
culturing chondrocytes in monolayer after isolation from cartilage tissue, the effective cellular yield
can dramatically increase. Since the number of chondrocytes harvested from human hyaline cartilage
is often small compared with the amount needed to seed a polymer scaffold, and donor tissue can be
scarce, expanding primary chondrocytes could lead to a more cost-effective therapy by creating
several constructs from cells isolated from a small biopsy [50,52,53].
Another advantage of culturing cells and scaffolds prior to surgical implantation is an increase
in the mechanical integrity of the graft. Since constructs grown in vitro have time to develop
cartilaginous tissue, their increased firmness allows easier implantation [50]. Also, in vitro culturing
allows much of the PGA to degrade resulting in a reduction of potential inflammatory or bulk release
interactions [48,50].
Additional work with PGA materials attempted to develop a model for in vitro cartilage
development by growing constructs in a reduced gravity bioreactor. Over 40 days, chondrocytes
proliferated three-fold, glycosaminoglycan concentrations reached 4.8% of construct wet weight or
68% as much as natural cartilage. Although the neocartilage contained a third of the amount of
collagen found in native tissue, over 90% of the formed collagen consisted of type II [53]. With the
recent excitement generated by the possibilities offered by stem cells, similar experiments have
evaluated the performance of chick bone marrow stromal cell differentiation on PGA disks. By
culturing the cells with fibroblast growth factor 2, constructs incubated in vitro for 4 weeks showed
fairly homogenous tissue formation. This tissue tested positive for several cartilage markers
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including proteoglycans and type II collagen, and the levels of these markers were comparable to
chicken epiphyseal cartilage. Neither biochemical nor histological results of the neocartilage were
affected by the amount of stromal cells used to seed the scaffolds. Results from this study
demonstrate that bone marrow samples may provide an alternative cell source for creating
autologous or allogeneic cartilaginous grafts [54]. Therefore, PGA remains a candidate polymeric
material as a cellular delivery vehicle for research involving long-term, large animal joint
resurfacing. One hindrance to many of the polyester scaffolds, however, is that repeatedly successful
results from research involving full joint resurfacing remains elusive since a feasible method for
attaching polymeric materials to an entire joint surface has not yet been developed [50].
Although several studies discuss histological and biochemical properties of neocartilage, few
papers quantify biomechanical parameters. Using the PGA scaffold protocols developed in Langer's
laboratory, Ma et al. performed compression and stress relaxation testing on constructs grown in
vitro. PGA felts without cells experienced a decrease in compressive modulus and aggregate
modulus as culture time increased. These results were due to PGA degradation. In contrast, the
compressive modulus of PGA-chondrocyte constructs increased and even surpassed that of normal
bovine cartilage after 9 weeks in vitro. Although the aggregate modulus of the polymer-cell
structures increased steadily over 12 weeks, the corresponding modulus for native tissue remained 10
times larger [55]. One reason that the aggregate modulus for neocartilage does not match that of
natural tissue may result from the lack of mechanical loading during in vitro culturing. Much of the
mechanical properties of cartilage result from interaction between the negatively charged
glycosaminoglycans and water. Without proper loading, chondrocytes may not produce a sufficient
amount of proteoglycans, and the resulting cartilage would lack the tremendous compressive
strength of normal cartilage. The third mechanical parameter measured the apparent permeability of
the scaffold materials. The permeability in PGA without cells quickly became too high due to
degradation of the scaffold. Seeded matrices, however, had permeabilities comparable to normal
cartilage after 3 weeks. Interestingly, cell morphology was not homogenous within the PGA mesh.
Cells covering the perimeter of the scaffold were much flatter than chondrocytes located more
centrally. As a result, the removal of the surface layer caused a dramatic increase in apparent
permeability, a 25% reduction in the value of the compressive modulus, but had no effect on the
aggregate modulus [55]. Long-term in vitro studies (lasting 25 weeks) indicated that several
properties including compressive modulus and permeability tended to stabilize after 12 weeks and
the recorded values reached levels typical of natural cartilage. Although it increased until week 20,
the measured aggregate moduli for the constructs still remained 40% lower than those for native
cartilage. These results suggested that longer in vitro incubation periods may not prove sufficient to
produce neocartilage with mechanical properties identical to those of hyaline cartilage [56].
However, cartilage grown first in vitro and subsequently implanted may not need to be completely
identical to that of natural tissue. Although compliance mismatch remains a concern, many
approaches hope to create a matrix with a homogenous disbursement of chondrocytes and
extracellular molecules that will become modified in situ to form hyaline cartilage [48].
Work performed by Chu et al. has tested some of the biological parameters of PLA foams. By
seeding rabbit costal chondrocytes on cylindrical cores of a D,D±L,L-poly(lactic acid) porous matrix,
the group was able to assess chondrocyte viability and conclude that the material could be used as a
high density carrier for chondrocytes [57]. Using the same highly porous scaffold, additional studies
were performed. After seeding D,D±L,L-PLA scaffolds with rabbit perichondrocytes, the constructs
were press fit into drilled articular cartilage defects in rabbits measuring 3.7 mm in diameter and 4±
5 mm in depth. Gross examination following 6 weeks in vivo showed a flat yet cartilaginous-like
tissue in all but one specimen. Biochemical analysis revealed that 80% of the collagen content of the
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neocartilage was type I and the remainder was type II. Both of these values as well as the
glycosaminoglycan concentrations were significantly different than normal hyaline cartilage, but
close to values for normal perichondrium [58]. These results indicate that the PLA scaffold did
support chondrogenesis; however, the characteristics of the newly formed cartilage may have been
determined more by the perichondrocyte phenotype than by PLA chemistry.
Sittinger et al. also probed the effectiveness of a-hydroxy materials as cellular scaffolds.
For their experiments, materials selection consisted of a manufactured nonwoven poly(L-lactide)
(PLLA) and a commercially available nonwoven 90:10 blend of PGA/PLLA (Ethisorb 210;
Ethicon). Adult human articular chondrocytes were harvested and used to evaluate polymer
biocompatibility and cell attachment. Degradation studies showed that the PLLA polymer needed
more than 1 year for complete degradation. In a different protocol, however, the faster degrading
PGA/PLLA (PGLA) material strongly lowered the pH below physiological range after just 3 weeks.
In contrast, the PLLA material maintained reasonable pH for periods up to 28 weeks. Since low pH
can accelerate degradation (look for reference), chondrocytes were exposed to various concentrations of L-lactic acid and glycolic acid to determine any biocompatibility issues. Results of
cellular activity after 24 h as a function of monomer concentration were similar up to 1 mg/ml. At
10 mg/ml, glycolic acid caused an almost complete reduction in chondrocyte activity; lactic acid did
not cause a significant decrease in activity. Over a period of almost 2 weeks, each monomer in
concentrations of 2 mg/ml greatly inhibited chondrocyte activity in media not adjusted for pH
[59]. These results could be important in experimental or treatment design in situations involving
the bulk release of monomeric substituents or where pH buffering or local metabolism rates are
compromised. Another part of the research involved coating the scaffolds with either polylysine
or collagen type II since these materials can support the chondrocyte phenotype [60]. For both
polymers, polylysine and collagen treatments significantly increased cellular adhesion relative to
untreated materials. Collagen-coated structures were intermediate in cell adhesion compared with
polylysine and controls [52,59]. Although neither polymer allowed complete uniformity in cellular
distribution, the greater hydrophobicity of the PLLA may have contributed to a higher cell seeding
density than that observed for PGLA [59]. In other experiments, controlled perfusion chambers
were used to grow seeded constructs. After 14 days, the resulting cartilage stained positively for
proteoglycans and collagen type II. Thus, chondrocytes retained phenotype and remained active
when maintained through automated equipment. The use of perfusion chambers offers a method
of cell culture that could allow the production of more uniform and reproducible neocartilage
[52,61].
One of the concerns regarding polymer processing for biological application is inability to
completely remove all residual solvents or salts that may be cytotoxic. Since poly(lactides) are
frequently made using chloroform or methylene chloride solvents, Gugala et al. wanted to develop a
more biocompatible PLA material using a process devoid of chlorine. To this end, they employed
salt leaching techniques using nonchlorinated solvents to create a 80:20 blend of poly(L-lactide) and
poly(D,L-lactide). The resulting cylindrical PLA scaffolds measured 20 mm in diameter, 2 mm depth,
had pore sizes ranging from 350 to 500 mm, and had porosities of 95%. The group then harvested
chondrocytes from sheep articular cartilage and seeded the disks with 250,000 cells each. After 9
weeks in vitro, chondrocytes not only proliferated, but also maintained their characteristic rounded
morphology. Histology sections showed that proteoglycan concentrations increased during the
culture time, but the proteoglycans were not homogeneously distributed throughout the matrix.
Although more work is needed to analyze the quality of neocartilage produced, the results of this
initial experiment showed that the PLA materials cast from nonchlorinated solvents performed in a
manner similar to that of traditionally-processed polyesters [47].
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Some modifications have been made to polyester matrices in order to alter the properties of
newly formed cartilage. Lohmann et al. evaluated two types of poly(D,L-lactide-co-glycolide)
(PLGA) scaffolds. The first was denoted PLG-H and consisted of 80% 75:25 PLGA and 20% 75:25
PLGA modified with a carboxyl end group. This carboxyl group increased the hydrophilicity of the
polymer, which can theoretically lead to better cell adhesion during seeding. The second scaffold,
PLG-FR, was 90% 75:25 PLGA with the remaining 10% containing reinforcing poly(glycolic) acid
fibers with diameters of 15 mm and lengths of 2.6 mm [62]. By orienting these fibers in a preferred
direction, an increase in compressive modulus and yield strength has been observed [63]. For this
study, PLG-FR had a Young's modulus of 32 MPa and a yield stress of 2 MPa. In comparison, the
values for PLG-H were 1.7 and 0.16 MPa, respectively. Resting zone cartilage chondrocytes were
isolated from the ribs of rats and seeded onto the two different polymers. The constructs were then
surgically inserted into the calf muscle of nude mice. After 8 weeks in vivo, implants, composed of
either polymer, contained varying degrees of cartilage and showed evidence of cellular hypertrophy.
Although the amount of cartilage increased for each type of polymeric construct over time, no
significant difference could be detected between the hydrophilic or reinforced fiber treatments. In
addition, all of the explanted constructs contained fibrous connective tissue. Since the PLG-H
material had a faster degradation rate and lower stiffness than PLG-FR, constructs composed of
PLG-FR were much better at maintaining the original implant shape [62]. Although biochemical
testing was not performed to evaluate the quality of cartilage developed in each construct, these
results involving modified forms of PLGA offer additional strategies in polyester matrix design and
demonstrate how chemical or mechanical properties can be altered in an attempt to provide a better
scaffold for cartilage repair.
A recent study by Ishaug-Riley et al. focused on characterizing the effect of different kinds of
commercially available polymers as substrates for human chondrocytes. Several polymers and
copolymers were investigated including poly(glycolide), poly(L-lactide), poly(D,L-lactide), 85:15
poly(D,L-lactide-co-glycolide), poly(e-caprolactone), 90:10 poly(D,L-lactide-co-caprolactone), 40:60
poly(L-lactide-co-caprolactone), 9:91 poly(D,L-lactide-co-caprolactone), 67:33 poly(glycolide-cotrimethylene carbone), and poly(dioxanone) [64]. Two properties these polymers share are that all
are considered degradable and all are approved by the FDA [65]. With exception to the
poly(glycolide-co-trimethylene carbone) copolymer (PGTMC), all of the other tested polymers are
classified as polyesters. Since these materials have chemical structures that differ in the location and
number of oxygen and methyl groups, these selected polymers could offer insight into how
variations in hydrophobicity could influence cellular behavior. Spin-casting techniques were
employed to create thin polymer films on glass coverslips measuring 22 mm in diameter. Tissue
culture polystyrene (TCPS) was used for all controls. Cell adhesion studies showed that PGA and
PGTMC surfaces contained the highest number of cells, whereas poly(e-caprolactone) and L-PLA
had the fewest cells. Statistical analysis, however, revealed similar attachment patterns for all
polymers relative to TCPS, and no trend was observed relating cell adhesion to polymer
hydrophobicity as measured by contact angles. Although an interesting study for attempting to
choose a scaffold material with optimal cell adhesion properties, the experimental results do not
indicate significant differences in two-dimensional cell seeding between the materials. In addition,
care must be taken when interpreting two-dimensional findings since both cells and polymeric
materials can assume drastically different characteristics in three-dimensional models [64]. Although
chondrocyte phenotype was not evaluated, this study successfully demonstrates that chondrocytes
can adhere to many of the polyesters approved by the FDA. As a result, further investigation of these
polymers is warranted; however, novel materials should also be evaluated as the degradable
polyesters have not yet achieved complete success.
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3.2.2. Polyethylene oxide
Both the location and varied geometries of most chondral defects cause difficulties when trying
to design a less-invasive polymeric device for cartilage repair. To help remedy this problem, several
groups have investigated the use of polymers that can gel in situ. One approach involving an
injectable polymer uses polyethylene oxide (PEO). Unlike alginate and PEMA, PEO can readily
degrade hydrolytically and can be cleared by the kidneys [66]. In addition, the high degree of
hydration within PEO gels acts as a barrier to macromolecules involved in the immune response. In
fact, PEO has often found use as a biomaterial due to its well-documented biocompatibility [67].
Previous work has succeeded in modifying the structure of PEO with various acrylates that allow the
PEO to be chemically crosslinked when exposed to ultraviolet light [66]. This process adds the
benefit of structural stability to the polymer system. As a result, the properties of PEO can be altered
in a way that allows more control over mechanical integrity and deterioration.
Some of the foundation for work with injectable polyethylene oxides for use in cartilage
regeneration came from the research of Sims et al. They developed a system to deliver a mixture of
bovine chondrocytes and PEO subcutaneously within nude mice. Initial experiments indicated that a
20% concentration (weight to volume) of 100,000 MW PEO created a viscous gel that could be
delivered through a 22-gauge needle. Constructs were created by suspending 10 million
chondrocytes/ml PEO gel. These gels were then injected into mice and observed for 3 months.
During this time interval, the glycosaminoglycan content of the neocartilage steadily increased and
reached 3.54% of wet weight at 12 weeks. In comparison, native adult bovine cartilage was
composed of 4.38% (of wet weight) glycosaminoglycans. DNA content also increased with time and
exceeded normal values by 60% at the conclusion of the experiment. Although present after 6 weeks,
the neocartilage at 12 weeks was statistically similar to native tissue with respect to these two
parameters. As a result, PEO can provide a biocompatible scaffold in which chondrocytes can
proliferate and secrete extracellular matrix components [68]. Despite these encouraging results, the
PEO gels formed by linear chain association experienced rapid diffusion and poor mechanical
stability [69]. In response, Elisseeff et al. sought to show that PEO gel integrity could be enhanced
by creating semi-interpenetrating networks through photopolymerization. Twenty percent of PEO
gels were created in phosphate buffered saline by combining PEO dimethacrylate (3400 Da MW)
with PEO (100,000 Da MW) in a 2:3 ratio. After adding chondrocytes and photoinitiator, the
suspension was injected into mice and polymerized using ultraviolet light. Results showed that the
chondrocytes survived the photopolymerization process [70]. Additional studies aimed to further
investigate the potential of PEO as a scaffold for cartilage repair. In one of these experiments,
various ratios of PEO and PEO dimethacrylate were investigated to determine how both the polymer
matrix and the polymer crosslink density would influence neocartilage formation. Throughout the
study, 20% polymer solutions were prepared from PEO and PEO dimethacrylate; however, the actual
formulations for these solutions contained 10, 20, 30, or 40% PEO dimethacrylate (PEODM).
Bovine chondrocytes were added to these solution to give a final density of 50 million cells/ml, and
1-hydroxycyclohexyl phenyl ketone was used as the initiator. After injecting athymic mice with the
different gels, the mice were placed under an ultraviolet A light (2 mW/cm2, 3 min exposure) to
allow the solutions to polymerize and then evaluated for up to 6 weeks. Biochemical analysis of the
constructs showed an inverse relationship between the amount of total glycosaminoglycan content
and the ratio of PEODM within the gel. Total collagen content, however, increased more as a
function of time than any other parameter. After 6 weeks in vivo, type II collagen only comprised
35% of the total collagen content within constructs made with 40% PEODM. Therefore, the
neocartilage developed in this experiment failed to match native cartilage composition. From a
mechanical viewpoint, both the 10 and 20% PEODM gels failed to create enough stabilization to
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truly differentiate them from noncrosslinked PEO. Only the 40% PEODM gels exhibited good
mechanical integrity following photopolymerization. In fact, this gel had a larger equilibrium
swelling volume than pure PEODM gels [69]. These results indicate that the mechanical properties
of PEO can be successfully modified to increase construct stability without sacrificing
biocompatibility. Recently, Elisseeff et al. performed another set of experiments to study in more
detail the biochemical and mechanical properties of neocartilage formed within PEO gels. These in
vitro studies showed the results of bovine and ovine chondrocytes seeded within 20% PEO gels (3:2
ratio of 100,000 MW PEO to 3400 MW PEODM) at densities of 50 and 40 million cells/ml,
respectively. Although biochemical analysis revealed a time dependent increase in both the total
collagen and glycosaminoglycan content for both types of chondrocytes, collagen concentrations
were much lower than native cartilage and even below levels found in neocartilage from some of the
other polymer systems discussed. Also, bovine constructs contained higher amounts of
glycosaminoglycans than ovine constructs even though the number of bovine chondrocytes
decreased with time and the number of ovine chondrocytes greatly increased with time. These results
demonstrate that chondrocytes obtained from various species and sources can react differently under
similar conditions. Thus, a successful polymer system must be sensitive or adaptable to the type of
cell they carry. Biomechanical testing showed that over an incubation time of 6 weeks, constructs
experienced continual increases in their equilibrium modulus, dynamic stiffness, and streaming
potential. Even with these increases, however, these measurements were still 10 times less than those
for natural cartilage [71]. The quality of neocartilage from experiments with PEO resembles the type
of cartilage produced from PGA scaffolds. In many ways, however, PEO offers more flexibility with
respect to cartilage repair. For example, as a viscous liquid, PEO can be processed into various
shapes for either in vitro or in vivo incubation. Furthermore, PEO can be injected and crosslinked in
vivo, which decreases both surgical time and risk. One limitation of the minimally invasive
photopolymerization process, however, is that the pre-polymer solution must be situated at a depth
that allows enough light energy to penetrate the skin [69,70]. One potential way to circumvent this
issue would involve the use of light sources, such as fiberoptics, in arthroscopic surgery. Future
developments for PEO hydrogels include experimenting with different crosslinking chemistry,
photopolymerization through visible light initiators, and crosslinking based on changes in ionic
concentrations [68]. As a result of these future endeavors, research on polyethylene oxide hydrogel
scaffolds for application in repairing cartilage defects continues to expand.
Many current investigations have sought to create composite materials for repairing
osteochondral defects. Since subchondral lesions affect bone as well as cartilage, some researchers
have attempted to accommodate a more complex tissue engineering matrix by combining materials
suited for both tissues. Since many of these approaches combine materials already discussed and a
specific discussion regarding many of the polymers used for bone reconstruction will follow, a
complete examination of dualistic polymer-based osteochondral plugs will not be presented here.
3.3. Natural degradable polymers
Other approaches to cartilage repair have focused on the use of natural materials as vehicles to
maintain chondrocyte phenotype or mimic the cartilagenous extracellular environment. Although
several naturally occurring materials can act as potential scaffolds, the most commonly investigated
materials have included alginate, collagen, hyaluronic acid, and fibrin glue. Additionally, one less
well studied biopolymer included in this review is chitosan. Much of the interest in these natural
polymers stems from their biocompatibility, relative abundance (i.e. commercial availability), ease
of processing, and/or possible ability to mimic the microenvironment found within cartilage.
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3.3.1. Alginate
Alginate has long been looked at as a polymeric biomaterial. Alginate is composed of two
repeating monosaccharides: L-guluronic acid, and D-mannuronic acid. Repeating strands of these
subunits form linear, water-soluble polysaccharides. Once these polysaccharides are exposed to
calcium ions, a three-dimensional gel quickly forms. Therefore, various pharmaceuticals, growth
factors, or cultured cells can be easily encapsulated. Calcium alginate gels do not degrade
hydrolytically, but can degrade enzymatically or in the presence of a chelating agent such as EDTA
[72]. In addition, the diffusion of calcium out of an alginate gel either in vitro or in vivo can cause
dissociation between alginate chains and can lead to a decrease in mechanical integrity over time
[73]. Since alginate is not native to the human body, however, concerns have been raised concerning
both a potential immune response and the lack of complete degradation [68,72]. Both guluronic and
manuronic acid contain negatively charged functional groups. As a result, these negatively charged
regions within the alginate scaffold can parallel the electrostatic conditions created by sulfated
proteoglycans in native tissue and might offer chondrocytes a more conducive environment [74].
Several groups have focused on the interaction between alginate and chondrocytes. For
example, Guo et al. seeded chondrocytes within alginate gels and characterized cellular morphology.
They found that chondrocytes within alginate retained their round shape compared with the more
fibroblastic appearing chondrocytes seeded in monolayer culture [75]. In vivo work has also
illustrated that cartilaginous tissue can form when alginate-cell suspensions are crosslinked in situ
[76]. Other work has since shown that dedifferentiated chondrocytes can redifferentiate when placed
in an alginate environment. Liu et al. encapsulated human articular chondrocytes within alginate
beads and examined DNA and proteoglycan synthesis for 70 days. Before seeding the chondrocytes
onto the alginate scaffolds, the chondrocytes were cultured in monolayer and lost their rounded
morphology. However, once embedded in an alginate matrix, the cells became rounded and began
aggressive production of proteoglycans. Both the ratio of chondroitin sulfate-6 to chondroitin
sulfate-4 and type II collagen mRNA increased with time, while type I collagen mRNA was
downregulated [77]. Although the mechanism for the alginate-chondrocyte interaction is not yet
fully understood, these results suggest that alginate contains some properties that allow chondrocytes
to regain correct morphology and phenotype. Furthermore, chondrocytes have survived within
alginate gels for up to 8 months [78].
Current alginate-cell gel preparations often contain nonhomogenous regions that result not only
from concentration differences, but also from radial channels formed from the presence of a cell
suspension [79]. Research by Aydelotte et al. prepared alginate beads with various channels by using
different solvents (water, sucrose, sodium chloride, and sucrose and sodium chloride) and by
changing the gelling solution. The number of formed channels depended on both the amount of cells
suspended in the beads and the concentration of sodium chloride in the solvent. As cellular density
increased and sodium chloride concentration decreased, more channels were observed. For example,
cellular densities of 0:5 106 or 1 106 cells/ml resulted in 10±15 and 20±30 channels per bead,
respectively. An increased concentration of sodium chloride in the gelling solution decreased the
number of channels. In vitro experiments showed that chondrocytes aligned along these channels and
proliferated in columns in a manner similar to chondrocytes along an in vivo growth plate. This
preliminary work shows that alginate can be modified with few channels (better suited for articular
chondrocytes) or many channels in order to mimic in vivo microstructure (e.g. growth plate
chondrocytes) [79].
Although alginate has shown encouraging results for chondrocyte proliferation, alginate itself
does not exist either within the cartilage matrix or within the body. Therefore, alginate has the
potential to induce an immune response. To this end, more recent work has aimed at modifying
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alginate gels with materials indigenous to the body. These new composites can then be used to
optimize the advantages of each component material. Lindenhayn et al. encapsulated articular
chondrocytes within alginate beads containing hyaluronic acid and fibrin. Hyaluronic acid was
examined because it exists readily within cartilaginous extracellular matrix, can bind with
chondrocytes through a plasma membrane receptor, and can form aggregates with proteoglycans
secreted by the chondrocytes. Thus, transplanted chondrocytes might better recognize this
hyaluronan environment and upregulate matrix production. At the same time, the ability of
hyaluronic acid to form aggregates would create a high density matrix by preventing proteoglycans
from diffusing out of the alginate beads [80]. Fibrin gels promoted chondrocyte cluster formation
and cartilaginous extracellular matrix deposition, but tend to degrade before producing a stable tissue
[72]. During in vitro experimentation, cultures containing hyaluronic acid and alginate showed a
greater increase in cell proliferation during the first 15 days compared with alginate beads alone.
Following day 15, alginate beads with both hyaluronic acid and fibrin experienced significantly
greater cellular growth than the alginate-hyaluronan composites. Additionally, the concentration of
alginate and presence of cells influenced the diffusion of hyaluronic acid. Decreased diffusion rates
of hyaluronan resulted from higher alginate densities that created smaller pores and the ability of
chondrocytes to bind with hyaluronic acid [80]. In further studies investigators seeded chondrocytes
within both fibrin-alginate and porous fibrin gels. The fibrin-alginate gels contained 4.5% fibrinogen
and 0.6% alginate and were supplemented with aprotinin. These gels could then be altered to form
porous fibrin materials by introducing a sodium citrate chelating agent to remove the alginate.
Results showed that these gels as well as alginate gel controls remained stable over a 30-day period.
Although gels containing fibrin allowed greater cellular proliferation than alginate gels, the cells
contained within the porous fibrin beads lost their phenotype and assumed a more fibroblastic
appearance after only 5 days. Chondrocytes within fibrin-alginate gels followed this trend but to a
much lesser degree. This dedifferentiation can account for the rapid rise in cell number.
Immunohistochemistry showed that the porous fibrin cultures did not contain any type II collagen. In
contrast, all of the gels containing alginate stained positively for type II collagen. Even though the
porous fibrin beads were designed to give chondrocytes more space for matrix deposition, these gels
could not give the cells enough structural support. The alginate-fibrin gels showed more promise and
remained stable for over 60 days. As a result, this hybrid material can offer insight into how the
microenvironments of composite materials affect chondrogenesis [72].
Results from recent investigations by Gregory et al. suggest that although the biochemical
composition of cartilage cultured in alginate gels can resemble that of native tissue, the structural
and spatial relationships between extracellular molecules can differ. By studying the effect of
alginate on the in vitro development of embryonic chick chondrocytes, this group saw evidence of
collagen segment-long-spacing crystallite-like formation instead of characteristic banded fibrils.
However, both collagen content and phenotype of the chondrocytes were comparable to the results of
other researchers. One offered explanation suggested that the polyanionic characteristics of the
alginate induced crystallite formation without affecting chondrocyte phenotype [81]. Even though
alginate gel did not seem to alter the biochemical composition of the extracellular matrix, the
abnormal organization of a matrix molecule, such as collagen, could result in different or even
compromised mechanical properties.
Much of the in vivo work focused on alginate-chondrocyte suspensions developed from
progress made by Vacanti et al. Since mannuronic acid can increase inflammatory responses, Paige
et al. used alginate gels consisting of more than 70% guluronic acid. Experiments showed that
cartilage could develop in gels formed from a wide range of both calcium chloride and alginate
concentrations. However, 30±50 mM calcium chloride and 1.0±1.5% alginate solutions produced
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optimal gels [49]. The groups further described a system in which bovine chondrocytes were first
seeded in an alginate solution and then injected subcutaneously into mice following the addition of
CaSO4. By using calcium sulfate instead of calcium chloride, the rate of the polymerization process
decreased to allow time for administration of the scaffold. Following 12 weeks, the alginatechondrocyte matrices contained hyaline-like cartilage formation characterized by increased
proteoglycan and collagen type II content [76]. Although their aim centered on craniofacial
cartilage reconstruction, both the material and process could be used for any type of cartilage. Later
work better characterized the alginate scaffold by comparing implanted constructs of equal
dimensions. Again, after 12 weeks of subcutaneous implantation, histological examination showed a
hyaline-like architecture containing many chondrocytes as well as increased amounts of
proteoglycans and collagen. Biomechanical compression testing confirmed cartilage formation
since constructs seeded with 10 million cells/ml had a mean force at failure almost ten times as large
as the controls [82]. Even though this mean force almost reached 5 MPa, it still fell short of some of
the loads experienced by human cartilage [22].
Building upon the work of Paige et al., Fragonas et al. used an in vivo rabbit model to evaluate
the performance of an alginate-cell suspension gelled in situ. They created 3 mm diameter defects in
the femoral condyle that extended to the subchondral bone and filled the cavity with either an
alginate solution (control) or an alginate-cell suspension. Gelation was induced quickly by adding
calcium chloride. After 6 months, the defects with the alginate controls contained fibrous tissue and
little glycosaminoglycans. In contrast, the repair mediated by the alginate-cell matrix was not
fibrous, blended well with surrounding cartilage, allowed glycosaminoglycan content to increase
over time, and showed evidence of cellular recruitment from subchondral bone. Although future
work needs to further characterize the biochemical composition and mechanical properties of
neocartilage formed from alginate gels as well as evaluate additional animal models, alginate
appears to provide a conducive environment for cartilage repair [83].
van Susante et al. explored the possibility of both alginate and collagen gels. They presented
results that showed bovine articular chondrocytes proliferated over 100% over 12 days when seeded
in type I collagen gels. Over the same time period, the number of chondrocytes in alginate gels
decreased. Additionally, proteoglycan synthesis in the alginate constructs continually increased
throughout the experimental time frame while proteoglycan production in collagen samples began to
decrease after 1 week [84]. Therefore, in contrast to the alginate system, the type I collagen structure
did not allow bovine chondrocytes to maintain their phenotype.
3.3.2. Collagen
Collagen gels have also been investigated as polymeric devices for supporting chondrogenesis.
Collagen is present in most types of connective tissue within the body. Although there are 16
different types of collagen, the most abundant kind of collagen is type I; type II collagen is most
prevalent in hyaline cartilage. Both type I and type II collagen fibers consist of an intricately
arranged network of fibrils. Molecularly, the primary amino acid sequence of collagen contains a
three peptide repeat; the first amino acid is glycine, and the next two positions are generally
occupied by proline and hydroxyproline, respectively. Three of these repeating protein chains
intertwine in a triple helix arrangement, associate with more helices, and ultimately form a
macromolecular fiber structure with properties resistant to pressure. As a result, collagen acts as a
very good structural matrix protein [4]. Research focused on cartilage repair has examined the use of
both type I and type II collagen scaffolds.
Schuman et al. investigated rat tail and calf skin as different sources for type I collagen. After
dissolving the collagen in acetic acid, bovine chondrocytes were added to the suspension at a density
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of 1 106 cells/ml. The suspension was then allowed to gel for 30 min at 378C after which an
overlay of media was added. During the first week of in vitro cell culture, chondrocytes within
bovine collagen gels assumed a more favorable morphology than cells within the rat collagen matrix.
After 14 days, however, 70% of the cells in the rat tail collagen and 60% of the cells in the calf skin
collagen had dedifferentiated to a fibroblastic state. Histological staining resulted in very modest
areas of possible proteoglycan synthesis. In addition, some of the gels were not mechanically stable
enough to transfer from the tissue culture wells. Due to these results, collagen did not seem like the
best candidate as a carrier device for cell transplantation, but modifications with growth factors,
higher seeding densities, higher collagen concentrations within gels, or longer culturing times could
improve the system [85].
Since few chondrocytes are harvested from donor tissue, Frondoza et al. wanted to look at ways
of using polymeric devices for increasing cellular yields. Recognizing that many cell types
proliferate at faster rates when cultured within microcarrier suspensions [86], they proposed the use
of 180 mm dextran beads, 155 mm dextran beads, 175 mm type I collagen-coated dextran beads, and
100±400 mm type I collagen beads to study the behavior of human chondrocytes. Cells were obtained
from the knees of adult human donors, cultured for several weeks, and then seeded following both
microcarrier and monolayer protocols. After 2 weeks, cells had attached to all of the dextran-based
beads; however, the chondrocytes had adhered so tightly that attempts to release them through
enzymatic cleavage methods resulted in cellular destruction. Cells attached to collagen beads easily
detached when exposed to trypsin. The microgravity reactors did provide an environment conducive
to greater cellular proliferation than standard monolayer culture. In addition, all of the microbeads
supported the chondrocyte phenotype as evidenced by the expression of type II collagen and the
downregulation of type I collagen [87]. Conclusions from these experiments show that collagen
microspheres can act as a tool to increase cell density without sacrificing viability.
The composition and structure of biomaterial scaffolds play a large role in influencing the
biological activity of seeded cells. Using collagen matrices, Nehrer et al. examined how collagen
type and pore size affect chondrocyte morphology and activity. To develop type I collagen materials
of different porosities, they altered the freeze drying procedure after coprecipitating bovine type I
collagen and shark chondroitin-6-sulfate. Analysis of the resulting scaffolds revealed uniform wall
thicknesses, pore sizes of either 20 or 83 mm, and porosities of 83 or 87%. Type II collagen sponges
were created by crosslinking porcine cartilage since noncrosslinked type II collagen matrices
degraded within 2 days. The crosslinking step helped stabilize mechanical integrity and decrease the
rate of degradation. All of the type II collaginous materials had pore diameters of 86 mm and a
corresponding porosity of 85%. In addition, matrices made from either type of collagen contained
about 2% glycosaminoglycans. During in vitro cell seeding studies, the various scaffolds were each
seeded with adult canine articular chondrocytes (previously cultured in monolayer) and examined
acutely (after 3 h) and over several weeks (up to 3 weeks). Due to their slower degradation rate, type
I sponges retained mechanical integrity for 3 weeks. The type II materials had almost completely
degraded after this time, but the cell containing constructs contained enough stability for handling,
implantation, and suturing. Within the type I collagen materials, scaffolds of either pore size
experienced 30±35% shrinkage after 7 days. Although acute findings showed that a significant
amount of chondrocytes seeded on matrices with smaller pore sizes expressed a more typical,
spherical morphology than cells exposed to larger pore sizes, results after 1 week revealed that only
27% of the chondrocytes maintained a rounded appearance regardless of matrix pore size. In
contrast, type II collagen materials supported rounded morphology for roughly 70% of the seeded
chondrocytes at both 3 and 7 days. Over time, however, the number of spherical cells on these
constructs decreased. Since phenotype often parallels cell morphology, these morphological
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observations accurately reflected the biochemical activity of chondrocytes as measured by construct
weight, glycosaminoglycan synthesis, and DNA synthesis. At 1 week, cells seeded on type I collagen
constructs with small pore sizes were more active biochemically than those cells in contact with
larger pores. Unfortunately, immunohistochemistry did not reveal the presence of any synthesized
type II collagen in any of the type I matrices. Type II constructs did stain positively for type II
collagen, however, no definite conclusions were made due to the inability to distinguish between
new collagen synthesis and the collagen content of the degrading pore walls. Overall, type II
materials experienced not only the highest increase in dry weight, but also the highest rate of GAG
synthesis per cell. This result suggested that type II collagen may prove a better scaffold for
neocartilage formation than type I collagen. Biologically, this conclusion seems reasonable since
most of native cartilage consists of type II collagen. Consequently, chondrocytes seeded in a type II
collagen-rich environment can more easily redifferentiate into the correct phenotype [88,89]. Nehrer
et al. expanded this work by evaluating the polymer systems in an in vivo canine model. They
created 4 mm diameter holes in the articular cartilage in the hind leg of each dog and then implanted
plugs comprised of type I or type II collagen with and without seeded autologous chondrocytes.
After 15 weeks, all of the dogs were sacrificed and examined grossly and histologically. Defects
filled with seeded type II constructs contained the most amount of reparative tissue, whereas half of
the type I matrices experienced incomplete filling. Both of the seeded collagen substrates induced
more tissue healing than either nonseeded scaffolds or empty defects. This repair tissue, however,
consisted of primarily transition tissue and fibrocartilage. More quantitative results showed that type
I and type II treatments contained 4 and 2% hyaline cartilage, respectively. Of all the experimental
groups, the untreated controls actually contained the highest percentage of hyaline cartilage (12% of
the reparative tissue). Although the collagen matrices did not perform well in this study, the authors
suggested that their implantation of freshly seeded scaffolds may not have allowed sufficient stability
and that a procedure incorporating long-term in vitro culturing may improve the quality of
subsequent tissue repair in vivo [90]. Research from other groups has supported this hypothesis. For
example, Kawamura et al. showed that when chondrocytes were seeded within collagen gels (0.15%
type I collagen) and subsequently cultured in vitro for several weeks, the cells became surrounded by
cartilaginous extracellular matrix molecules and could then receive signals from a more natural
environment. As a result, the success rate of these constructs increased relative to both controls and
freshly-seeded gels [91].
Using a rabbit model, Frenkel et al. had some success with collagen scaffolds. They used a twolayered bovine type I collagen matrix (4 mm diameter) marketed by Integra LifeSciences. The deep
layer consisted of a dense matrix designed to inhibit fibroblast infiltration from the subchondral
bone, and the more superficial layer had greater porosity in order to support chondrocyte phenotype.
After extracting cells from the knees of rabbits, the authors seeded the collagen scaffolds with the
chondrocytes and cultured the constructs for 2 weeks prior to implantation. At all time points from 6
to 24 weeks post-implantation, the seeded scaffolds induced a more hyaline-like cartilage repair
compared with controls containing unseeded matrices or no treatment. Biochemically, the seeded
implants contained normal levels of type II collagen (more than 90% of total collagen) after 6 weeks,
which continued to increase up to week 24. In contrast, unseeded implants and empty defects showed
a decreasing trend in the amount of type II collagen as time progressed. All defects experienced an
increase in glycosaminoglycan production over time; however, only the defects filled with a seeded
or unseeded scaffold contained normal levels by 24 weeks. The authors measured but did not report
the aggregate moduli for the implants. They also indicated the need to conduct in vivo studies
beyond 6 months to characterize the long-term effects of this method before looking at the repair of
larger defects [92].
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Although not an inherent parameter for polymer design, the incorporation or release of growth
factors can improve the biological performance of a polymeric cell carrier. Regarding the
upregulation of chondrocyte activity within collagen scaffolds, many groups have studied the effect
of basic fibroblast growth factor (bFGF). Fujisato et al. harvested rat costal chondrocytes and seeded
the cells on collagen sponges with or without bFGF. They then implanted the materials
subcutaneously in the backs of nude mice. Implants containing bFGF experienced an accelerated
rate of cartilage regeneration compared to those without the growth factor [93]. To investigate the
effect of bFGF incubation time on chondrocyte culture, Toolan et al. seeded lapine chondrocytes on a
type I collagen matrix and cultured the constructs in the presence of bFGF and insulin. Biochemical
analysis did not indicate any difference in the aggregate modulus of elasticity between constructs
bathed in media alone and those bathed in growth factor supplemented media. After 9 weeks, this
modulus measured roughly 0.68 MPa. The addition of the bFGF and insulin did contribute to a
significantly higher glycosaminoglycan content in the scaffolds after 6 weeks of culture. Even
though the values at 9 weeks decreased to levels observed after 2 weeks, the addition of growth
factors still doubled glycosaminoglycan content [94]. As found in these other studies, Matsusaki et al.
confirmed that bFGF induced chondrocyte proliferation. In contrast to the study performed by
Toolan et al., however, they found that bFGF inhibited the synthesis of proteoglycans. This result
may have occurred due to the lack of insulin in their cultures or due to other differences in the
culturing environment. One important observation from Matsusaki et al., is that although it allowed
greater cellular proliferation, the presence of bFGF did not seem to negatively influence the
chondrocyte phenotype. Instead, it seemed that the material to which the cells attached played a
more important role in both morphology and phenotype [95]. Thus, the addition of biochemicals,
such as bFGF, may offer a way to augment cellular activity while the polymeric device supports the
desired phenotype.
3.3.3. Hyaluronan
As mentioned previously, one attractive characteristic of polyester materials is the ability to
allow biologically compatible hydrolytic degradation. As mentioned previously, these materials have
received widespread attention; however, by themselves, these a-hydroxy polymers have not proven
completely successful materials for tissue regeneration. One recent biomaterial for cartilage repair
incorporates the modifiable degradation rates offered by the polyesters into hyaluronic acid, a natural
glycosaminoglycan polymer found abundantly within cartilaginous extracellular matrix. By altering
the chemistry of an esterification reaction, new materials can be produced that allow for increased
biocompatibility, tailored degradation rates, and a microenvironment more similar to that of natural
cartilage.
In its natural form, however, hyaluronan lacks several desirable characteristics for use as a
polymeric scaffold material. Some properties of HyA that are not conducive for biomaterials
applications include high water solubility and fast resorption and tissue clearance times [96]. The
molecular structure of hyaluronan, however, offers many options for increasing stability through
chemical modification. Some research groups have considered various mechanisms for crosslinking
and coupling HyA chains within synthetic matrices or to other HyA chains. The review by
Campoccia et al. offers a good introduction on the specifics of such HyA modifications [96]. For
applications involving cartilage repair, many recent papers have focused on HYAFF1 11. Made by
Fidia Advanced Biopolymers (Abano Terme, Italy), HYAFF1 11 is an ester derivative of natural
HyA. By esterifying the carboxylic group on the glucuronic acid resides of 80±200 kDa sodium
hyaluronate with various alcohols, the aqueous solubility of HyA can be altered from completely
water soluble materials to insoluble hydrogels. For example, an increase in either the degree of
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esterification or the hydrocarbon content of the added alcohol will increase hydrophobicity. Thus,
HYAFF1 11 can be created with tailored degradation rates since hydrophobicity directly influences
hydration and the de-esterification reaction [96]. In addition, these ester derivatives can be easily
processed by extrusion, lyophilization, or spray drying to form sponges, microspheres, meshes,
fibers, or thin membranes [96,97]. Unfortunately, when hydrated, these HyA derivatives lose much
of their mechanical strength [96].
Recently, HYAFF1 11, the derivative formed from the complete esterification of all free
carboxylic groups with benzyl alcohol, has been examined as a degradable scaffold biomaterial for
tissue regeneration. By incorporating the aromatic benzyl group into the polysaccharide backbone,
HyA becomes insoluble in aqueous solutions, but still retains an X-ray diffraction pattern similar to
natural HyA [96]. Although insoluble in protic solvents, HYAFF1 11 will hydrate to a 40% increase
in weight and spontaneously and completely hydrolyze in artificial plasma over a period of 2
months. As the benzyl group hydrolyzes off the glucuronic acid residues, the resulting carboxylic
function group restores the HyA derivative to its more natural, water soluble configuration. As a
result, the relatively short HyA chains can dissociate from the mesh and become more subject to
enzymatic degradation. Thus, HYAFF1 11 can provide several weeks of mechanical support for
tissue infiltration before deteriorating. Since hydrolysis yields HyA, the byproduct of degradation
not only remains biocompatible, but may also help establish a more natural extracellular matrix [96].
HYAFF1 11 membranes can contain 20 or 40 mm thick fibers, 10±400 mm pores, a 10 m2/cm3
surface area, an 80% porosity, and a specific weight of 100 g/m2 [96±100]. Upon hydration, the
extruded fibers experience swelling that can roughly double fiber diameter [98].
Published results indicate that HYAFF1 11 shows considerable promise for cartilage repair
[96±98]. The experiments performed by Aigner et al. demonstrated that the material not only offered
a biocompatible matrix for cellular proliferation, but also seemed to support the chondrocyte
phenotype. These results have since been reported by others [96,97]. When tested in vitro with
human nasoseptal chondrocytes initially cultured in monolayer, HYAFF1 11 scaffolds seeded with
cells seemed more stable compared with scaffolds without cells. Also, the chondrocytes proliferated
within the porous regions of the material and adhered to the fibers themselves. To investigate the
degree of redifferentiation, type I and type II collagen content was examined qualitatively using
immunocytochemistry. Although the cells followed a trend of increasing redifferentiation as
suggested by an upregulation of type II collagen, experimental results revealed a continual presence
of type I collagen. In vivo experiments followed this trend. Seeded scaffolds subcutaneously inserted
in the backs of mice did show evidence of cartilage formation after 26 days. However, results of
immunohistochemistry indicated the presence of both type I and type II collagen. While the ratio of
type II collagen to type I collagen needs further investigation, these studies do suggest that the
hyaluronan material appeared to positively influence chondrocyte phenotype since none of the cells
used to seed the scaffolds expressed type II collagen while in monolayer culture [98].
The potential of the HYAFF1 11 system was further investigated when Solchaga et al. seeded
the matrices with mesenchymal progenitor cells. Due to their ability to differentiate into various cell
types such as osteoblasts and chondrocytes, mesenchymal progenitor cells continue to receive
considerable attention by researchers in the field of osteochondral repair. For their studies, Solchaga
et al. harvested these cells from the bone marrow of rabbits and seeded them onto HYAFF1 11
sponges, ACP sponges, and calcium phosphate ceramic. Also created by Fidia Advanced
Biopolymers, ACP sponges are formed by condensing the carboxylic groups of unmodified
hyaluronan molecules with the hydroxyl groups on other HyA chains. The resulting crosslinked
material has similar porosity (85%) but slightly smaller pore sizes (10±300 mm) and overall surface
area (7.34 m2/cm3) compared with HYAFF1 11. Unlike HYAFF1 11, ACP remains highly unstable
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and degrades in solution after 1 week. The calcium phosphate ceramics were used as control
materials since their interaction with mesenchymal progenitor cells has been well characterized. Cell
seeding experiments were performed with and without presoaking the materials in a solution
containing fibronectin. Preconditioning the scaffolds with fibronectin caused an increase in cell
adhesion for all materials. In general, twice as many cells adhered to both the coated and uncoated
HYAFF1 11 (fibronectin-coated HYAFF1 11 samples incorporated the most cells) and coated ACP
materials than the fibronectin-coated ceramic controls. Untreated ACP had the same number cell
density as the controls. ACP materials, however, degraded in little over a week. As a result, in vivo
studies could only compare the ceramic controls with HYAFF1 11 since the ACP sponges
completely degraded and left no evidence of bone or cartilage formation at the site of surgical
implantation. Three weeks after implanting the materials subcutaneously into the backs of mice,
examination of the devices revealed that the cell-HYAFF1 11 scaffold induced significantly more
cartilage or bone formation than the cell-ceramic material. Without the presence of the mesenchymal
progenitor cells, however, only fibrous tissue was found. Additionally, results from animals
sacrificed 6 weeks after implantation did not differ from the 3-week data. Although mesenchymal
progenitor cell preparations can vary widely in their ability to differentiate into chondrocytes or
osteoblasts, these results indicate that HYAFF1 11 can successfully support mesenchymal cell
proliferation and differentiation for osteochondral applications [99].
Also acknowledging the promise of stem cell research, Radice et al. analyzed the extracellular
matrix components formed from mesenchymal progenitor cells seeded on HYAFF1 11. For both in
vitro and in vivo studies, the group harvested cells from the bone marrow of rabbits and from the
iliac crest of human patients. Results from the in vitro cell seeding experiments indicated that after 3
weeks of static culture, the cells from both species were well dispersed within the HYAFF1 11
scaffolds. Immunohistochemistry of the cultures confirmed the presence of fibronectin, laminin, and
types I, II, IIA, III, and IV collagens. In general, type IIA collagen was found in all examined
cultures. Since type IIA collagen represents the precursor for type II collagen, tissue sections rich in
type IIA collagen indicate the majority of the cultured cells were expressing a prechondrogenic
phenotype with the potential for further differentiation. In vivo experiments were performed by
inserting naked scaffolds and cell-scaffold materials into 3 mm diameter cylindrical osteochondral
defects created in the knees of rabbits. Four months after the surgical implantation, the HYAFF1 11
treatments (both with and without cell seeding) promoted faster repair filled with more hyaline-like
cartilage than the empty defects used as controls. Comparisons between treatments revealed no
significant differences between scaffolds seeded with cells and those without cells. This result
suggested that the HYAFF1 11 material may support the recruitment of the host organism's own
mesenchymal stem cells or other cells involved in osteochondral repair. Each of the control defects
developed fibrous tissue, but only 37% of the defects filled with the hyaluronan matrices contained
fibrous tissue. Of the defects filled with cartilage-like tissue, histological analysis showed that
fibrocartilage formed less than 30% of the time. Although the scaffold took about 4 months to
degrade, the experiments showed evidence of zonal formation 2 months following implantation.
Cells located superficially differentiated into articular-cartilage-producing chondrocytes, and deeper
cells showed signs of ossification and incorporation into subchondral bone [100]. Thus, the
HYAFF1 11 material remains a promising material even though it cannot guarantee a successful
cartilage defect repair. While long-term in vivo studies are necessary for further evaluation,
HYAFF1 11 shows significant potential as a viable osteochondral biomaterial.
An additional modification to hyaluronan gels involves sulfating free hydroxyl groups. Barbucci
et al. created a variety of sulfated derivates ranging from one to four sulfate groups per disaccharide
subunit. Using diamines, the group controllably crosslinked individual hyaluronic acid chains
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together to form hydrogels. When ovine chondrocytes were seeded onto patterned surfaces
containing regions of sulfated hyaluronan and regions of poly(ethylene terephthalate), the
chondrocytes tended to migrate towards the hyaluronan and appeared well spread compared to
cells on the PET. Studies involving this alteration are still preliminary, however, this work
demonstrates a different approach for derivatizing hyaluronan to create a new polymeric network for
potential use in cartilage repair [101].
3.3.4. Fibrin gels
Another class of naturally occurring polymers used for cartilage repair involves fibrin glues.
Fibrinogen self-assembles to become fibrin upon injury to a blood vessel. Fibrin plays a major role
during wound healing within the body. The actual polymerization process occurs as fibrinogen
molecules become active. Each fibrinogen molecule is composed of an a-, b-, and g-subunit. The aand b-chains contain ligands that have affinity for their respective binding pockets within the fibrin
molecule. In the inactivated form of fibrinogen, however, each ligand is flanked by a short peptide
sequence that acts as a protective cap. In the presence of the enzyme thrombin, this peptide sequence
is cleaved, and the ligand becomes active. The ligands of one fibrinogen molecule will then
associate, through hydrogen bonds and electrostatic interactions, with the binding pockets on other
fibrin molecules. As a result, a random, three-dimensional network will form. Physiologically,
thrombin also activates a transglutaminase enzyme called factor XIIIa. Factor XIIIa will chemically
crosslink lysine and glutamine residues found on the a-, b-, and g-chains of different fibrin
molecules. This enzymatic reaction converts the three-dimensional network into a covalent hydrogel.
Fibrin gels can degrade either through hydrolytic or proteolytic means. An example of an enzyme
that can actively dissociate fibrin is plasmin. Often isolated from blood plasma, fibrinogen is
commercially available from various manufacturers. Thus, the cost of making fibrin gels remains
relatively low. In addition, fibrinogen can be obtained from a patient's own blood which limits the
potential for disease transmission or immunogenic reactions. Although dependent upon thrombin
concentrations, the gelation of fibrinogen into fibrin occurs quickly (within 30±60 s) following the
addition of thrombin. Due to its natural role in wound healing, fibrin has been investigated and used
as a clinical fixative. More recently, many studies have investigated fibrin as a potential tissue
engineering scaffold.
Some of the early work involving both fibrin and chondrocytes aimed to evaluate the ability of
fibrin gels to support chondrocyte viability. Homminga et al. created several fibrin gels seeded with
different concentrations of cells and incubated the gels at physiological conditions for 1 week. They
found that fibrin was not cytotoxic to chondrocytes; rather, fibrin allowed the cells to maintain
phenotype. The gels, however, began degrading in as little as 3±4 days [102]. Building upon this
work, Sims et al. made in vivo observations regarding fibrin as a scaffold for neocartilage formation.
In their study, bovine chondrocytes were suspended within a fibrin gel at a density of
12:5 106 cells/ml. These constructs were then subcutaneously implanted in the backs of athymic
mice. After 12 weeks, the resulting tissue had a macroscopic and physical similarity to natural
cartilage. Histology confirmed the formation of cartilaginous extracellular matrix, and biochemical
analysis revealed that the neocartilage had glycosaminoglycan and DNA concentrations slightly
lower but comparable to those of native bovine cartilage [103]. These positive preliminary
investigations led to further work on the fibrin system. Ting et al. modified the previous work in two
ways in a study to evaluate new options for craniofacial reconstruction. First, in order to evaluate
constructs containing human chondrocytes, they harvested cartilage from a donor undergoing
surgical treatment for a costal disorder. Second, the group sought to culture cartilage in prefabricated
molds. Specifically, a human nasal construct made from ethylene-vinyl acetate copolymer was used.
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An amount of 25 106 chondrocytes/ml were suspended in a fibrinogen solution, injected into the
mold, and polymerized with thrombin. The newly formed construct then developed in vitro under
static tissue culture conditions for 4 weeks. After 4 weeks, the construct was subcutaneously
implanted into a mouse and allowed to grow for an additional 4 weeks. Unfortunately, the limited
number of viable chondrocytes only permitted the creation of one fibrin scaffold with human cells.
Although the construct retained some nasal features, upon extraction from the mouse, examination
showed that the construct experienced a 75% reduction in total volume during in vitro incubation and
only contained 12% of the original volume after in vivo testing. No residual fibrin glue remained
within the construct. Histology revealed the presence of glycosaminoglycans, collagen, and viable
chondrocytes. More quantitative results indicated that the glycosaminoglycan concentration fell
within the range of published values for bovine cartilage; DNA content was twice that of bovine
cartilage. The study also included a biomechanical test for the neocartilage that identified an
apparent modulus (slope of the stress±strain curve). After subjecting the construct to a 200 mN static
stress and a 80 mN, 5 Hz dynamic stress, the apparent modulus was calculated to be 0.4 MPa. As
with most tissue engineered cartilage, this value falls short of the 3.1 MPa modulus recorded for
native cartilage [104]. Although further testing of human chondrocyte cultures is necessary, these
results support a hypothesis that the fibrin gel scaffold may offer better success for cartilaginous
cosmetic reconstruction than for articular joint resurfacing.
Some attempts have been made to modify fibrin gels. Recently, the work by Meinhart et al.
attempted to stabilize fibrin so that gels would remain intact for longer periods of time. In culture
and within the body, fibrin remains subject to hydrolytic or enzymatic degradation known as
fibrinolysis. Commercially-available fibrin gel systems, such as Tissucol1 (Immuno AG, Vienna,
Austria), contain certain concentrations of aprotinin, a basic polypeptide that inhibits several serine
proteases. During their experiment, Meinhart et al. suspended various densities of human
chondrocytes (1:25 105 to 20 106 cells/ml) in each of two different fibrin gel preparations.
The first preparation contained a standard amount of 3000 kIU (kilo international units)/ml aprotinin.
The second, modified treatment had 8500 kIU/ml aprotinin and 15 mg/ml tranexamic acid. Each of
the resulting constructs was incubated in vitro for 4 weeks. Standard fibrin gels seeded with at least
4 106 cells/ml showed initial signs of degradation after 3 days in culture, and all of these gels had
completely degraded after 4 weeks. When these standard gels were seeded with low cell densities
(no more than 1 106 cells/ml), a solid construct remained throughout the culture interval; however,
no cartilaginous extracellular matrix components were present after histological staining. All of the
gels with increased resistance to fibrinolysis remained solid during the 4 weeks, although the gels
seeded at high densities were not as thick as the lower density gels. In contrast, histology showed that
high-density gels had much larger areas of proteoglycan and type II collagen deposition than the
lower density gels. The preliminary results presented by this study suggested that the addition of
aprotinin and tranexamic acid to the fibrin gels did inhibit fibrinolysis without compromising
chondrocyte viability or phenotype [105]. Thus, fibrin gels can be manipulated to provide some
additional mechanical stability while chondrocytes develop an extracellular matrix. Furthermore, the
experiment illustrates how factors not intrinsic to the polymeric vehicle, such as cell-seeding density,
can greatly affect overall material performance.
To help identify methods for creating neocartilage of high quality, similar investigations have
probed some of the parameters involved in fibrin gel cell suspensions. For example, Silverman et al.
studied how fibrinogen and chondrocyte concentrations influenced neocartilage formation in
athymic mice. By decreasing the concentration of thrombin, the gelation process slowed to allow in
situ polymerization following subcutaneous injection of the fibrinogen/chondrocyte suspension. In
the first part of the experiment, mice received injections with various fibrinogen concentrations (20,
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50, 80 mg/ml) but without chondrocytes. The effect of aprotinin was also tested by injecting another
group of mice with 3000 units of aprotinin/ml mixed with a fibrinogen concentration of 80 mg/ml.
As expected, gels composed of lower starting fibrinogen concentrations weighed less than higher
density gels after 4 weeks in vivo. Gels containing aprotinin had significantly higher weights than
gels without the fibrinolytic inhibitor. Therefore, either an increase in fibrinogen concentration or the
addition of fibrinolytic inhibitors will stabilize fibrin gels. Due to its more gradual degradation rate, a
fibrinogen concentration of 80 mg/ml was used to determine the optimal cell seeding density for
fibrin gels. In this second part of the experiment, constructs containing 10, 25, or 40 million porcine
chondrocytes/ml were evaluated up to 12 weeks in vivo. Although injections of chondrocytes
without a fibrin matrix developed some cartilaginous tissue, all of the fibrin-cell constructs resulted
in larger and more well defined cartilage nodules. However, the gels with 40 106 cells/ml
contained the most regions of homogenous cartilage formation. Collagen typing showed a direct
correlation between the cell seeding density and the total percentage of type II collagen produced.
Only the gels containing the most chondrocytes approached native tissue with respect to type II
collagen expression; collagen type II made up less than 50% of total collagen in the other gels [106].
Overall, these experiments demonstrate the capacity of fibrin both to support the chondrocyte
phenotype in vivo and to provide a minimally invasive approach for repairing cartilage defects.
Apparent within the study remains an inherent tradeoff between cell seeding density and scaffold
stability. Although higher initial chondrocyte seeding densities allow a more even and homogenous
distribution of cartilaginous extracellular matrix, an increased number of cells can lead to a faster
breakdown of the fibrin gel. If gels degrade too quickly, construct stability decreases, chondrocytes
can migrate away from the area of treatment, and a more fibrous tissue may develop within the
cartilage defect.
Many of the previous fibrin studies presented or cited data obtained from in vivo
experimentation on athymic mice. Since they lack an appropriate immune response against
xenografts, athymic mice provide a good animal model for autologous transplantation research and
initial, proof-of-concept results. However, successful treatments in these animals may not prove
effective in other, nonimmunocompromised animal models [107]. Although fibrinogen purified from
the blood of individual patients can offer an autologous cell-delivery vehicle, many groups are trying
to develop more general allogeneic solutions for hyaline cartilage restoration. In many of these cases,
however, fibrin has not proven to be an effective matrix material.
One group, van Susante et al., has conducted several studies on fibrin as a cellular delivery
vehicle for repairing articular cartilage defects. For their experiments, van Susante et al. used
Tissucol1, a two-component fibrin glue system developed by Immuno AG (Vienna, Austria). In one
study, this group evaluated the effectiveness of the material in large, subchondral defects in goat
articular cartilage. Harvested rabbit chondrocytes were suspended within the fibrinogen precursor
solution at a density of 10 106 cells/ml. The cell suspension was then injected into pre-drilled
cylindrical defects (10 mm diameter and 4 mm depth) created in a load-bearing region on the goat's
knee and allowed to polymerize. One advantage of using fibrin in this fashion was the ability to
completely fill the defect by gelling in situ. The defects used as controls were left empty, and the
quality of tissue formation for each treatment was examined over a 1 year period. Fifty two weeks
following surgery, neither gross observation nor histological evaluation revealed a difference
between the fibrin-cell suspension and the controls. In fact, after 13 weeks, no viable chondrocytes
were present within the fibrin grafts, but inflammatory cells were actively degrading the fibrin
matrix; no residual fibrin was found after 26 weeks. Tissue that did develop within the defects treated
with fibrin had fibrocartilaginous characteristics indistinguishable from that of the tissue within the
controls. One encouraging trend was the continual increase in the ratio of type II collagen to total
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collagen content. Normal articular cartilage isolated from goat consists of 98% type II collagen. In
this study, the repair tissue contained 17% type II collagen at week 2 and steadily increased to 75%
by week 52. Unfortunately, this trend cannot be attributed to the polymeric delivery system since the
fibrin and control groups yielded the same results throughout most of the experiment. One
hypothesis for the failure of the fibrin system was that the poor mechanical properties of fibrin gels
could not provide enough stability for the cell system in large defects. As a result, this lack of
support allowed chondrocytes to migrate out of the gel and away from the defect [108]. Although the
fibrin system did not prove a successful matrix for large cartilage defects, van Susante et al.
hypothesized that a chondrocyte-fibrin suspension might provide a scaffold for hyaline cartilage
production if supported by a stronger material. They chose hydroxyapatite (HA) as a subchondral
bone substitute and formed composite materials by layering 1 ml of the cell-fibrin suspension on a
10 mm HA cylinder. Following the same procedure as the previous study, they created 10 mm
diameter cylindrical defects in the knees of goats and filled the space with the HA composites.
Results were less than promising since small regions of hyaline-like cartilage seen at week 4 were
completely replaced by fibrous tissue by week 12. Again, inflammatory cells actively degraded the
fibrin matrix, and after 24 weeks the fibrin matrix had completely degraded. In addition, the HA did
not induce proper subchondral bone formation. These results suggest that although chondrocytes can
actively produce extracellular matrix components within fibrin, the fibrin gel system does not offer
enough mechanical integrity to allow articular cartilage formation in areas of high loading. van
Susante et al., keenly acknowledged that larger animal models, such as goat, may lack the more
robust intrinsic cartilage repair mechanisms observed in smaller animal models, such as mice and
rabbits. As a result, the quality of the repair tissue generated during articular cartilage experiments
performed in larger or mature animals depends much more upon the physical and chemical
characteristics of the biomaterial scaffold or cellular delivery vehicle [109].
3.3.5. Chitosan
Another biopolymer material of interest to cartilage research is known as chitosan. Chitosan, a
polysaccharide derived from chitin, consists of a relatively simple glucosamine monomer. In fact, the
chemical structure of chitosan produces properties similar to those of many glycosaminoglycans.
Past research has shown that chitosan is relatively biocompatible and biodegradable, does not evoke
a strong immune response, and is low in cost due to its abundance and the diverse methods of
chemically processing the polymer [110]. Sechriest et al. developed a hydrogel based on chitosan.
Recognizing the importance of mimicking the extracellular environment, they crosslinked
chondroitin sulfate with chitosan with the hope of providing an environment conducive for
maintaining chondrocyte phenotype and inducing cartilage formation. Their initial study did not
attempt to create three-dimensional scaffolds; rather, they prepared thin, composite membranes to
investigate the effect of the novel polymer on chondrocyte culture. When chondrocytes were seeded
onto the films, the authors observed that chondrocytes on the chondroitin sulfate-chitosan (CSAchitosan) maintained a more rounded morphology than cells on polystyrene controls. Chitosan
without chondroitin sulfate could not serve as a control because pure chitosan could not support
adequate cell adhesion. Thus, the presence of chondroitin sulfate provided a signal for chondrocyte
anchorage. Correlating with the morphological results, CSA-chitosan chondrocytes proliferated five
times less than their polystyrene counterparts. Biochemically, the synthesis of proteoglycans was not
different from controls. In contrast, however, the polystyrene control cultures contained virtually all
collagen types found in fibrous cartilage and primarily produced type I collagen, whereas the CSAchitosan material induced a much higher level of type II collagen. Quantitatively, the type I-to-type
II ratio for polystyrene surfaces was 2:1; the ratio was 2:3 for CSA-chitosan surfaces. Since their
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experiments only lasted 1 week, the authors concluded that future work needs to examine longer
cell culture time and the possibility of fabricating a more three-dimensional material in order to
more clearly evaluate the biochemical composition of neocartilage grown on a CSA-chitosan
construct [111].
4. Bone regeneration
It is common knowledge that the body can efficiently repair fractures in bone. These small
defects are often replaced by natural bone, which is even stronger than the original bone. Larger
defects, however, generally cause more trouble since the body is incapable of repairing them. Bone
defects can occur as a result of congenital abnormalities, trauma, or disease. Traditional methods for
filling these defects have used acellular cadaver bone or autologous bone. Both have serious
problems associated with them; the former method can result in disease transmission while the latter
is associated with two surgical/wound sites and morbidity at the donor site. Alternative methods have
been developed, but no method has yet provided a satisfactory solution. As a result, researchers and
the medical community are turning toward the promising field of tissue engineering to develop new
methods of bone regeneration.
Bone is a very dense, specialized form of connective tissue. The bone matrix consists of type I
collagen and calcium phosphate in the form of hydroxyapatite. A compact, dense cortical layer
(compact bone) comprises the outer region of long bones, while trabecular bone (cancellous bone)
fills the interior. As seen in Fig. 5, the major structural component of compact bone is called an
osteon. Composed of a concentric lamellar matrix, osteons create cylindrical conduits known as
Haversian canals, which provide access for the circulatory and nervous systems. The capillaries
within the Haversian canals originate from arteries and veins within the marrow cavity. The
Volkmann's canals, which are transverse to the Haversian system, provide the pathway for bone
nourishment. With the exception of articulating surfaces, cortical bone is surrounded by a thin
connective tissue, the periosteum, which consists primarily of a collagen-rich fibrous layer and
osteoprogenitor cells. Fig. 6a illustrates the structure of trabecular or cancellous bone. As the name
implies, cancellous bone consists of a lattice-like structure comprised of trabeculae, which are small
spicules of bone surrounded by marrow. A more detailed drawing of trabecula can be seen in Fig. 6b.
Within each trabecula, osteocytes maintain bone matrix, osteoclasts degrade regions of existing
structure, and osteoblasts (especially those associated with the osteoid region) actively build new
sections of bone into the interconnecting marrow spaces [4,112,113]. The regeneration of bone tissue
involves not only the synthesis of hydroxyapatite rich collagen scaffolding, but also the recreation of
an intricate structure that lends itself to sensation and mechanical stability.
As with all materials implanted into the body, polymers for bone regeneration must be
biocompatible. In addition, these materials should be either moldable, shapeable, or polymerizable in
situ to ensure a good fit in the defect area. They should support cellular adhesion and growth,
maintain cellular differentiation, provide a porous matrix through which nutrients and wastes can
easily diffuse, and degrade into biocompatible byproducts. It is also crucial that these materials have
mechanical properties similar to native bone. Human trabecular bone typically has a compressive
strength of 5 MPa and a modulus of 50 MPa [114]. Table 1 shows the relationship between these
results and the measured moduli of commonly used polymers. The material must maintain its
mechanical properties as it degrades until the newly regenerating tissue can adequately support
loading. If it fails mechanically, the material may lead to the failure of a patient's arm or leg. On the
other hand, if the material is too strong, it may cause stress shielding of the remaining natural bone
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
Fig. 5. A pen drawing of the structure of cortical bone adapted from that given by Frank Netter [263]. The bone matrix
consists of type I collagen and calcium phosphate in the form of hydroxyapatite. A compact, dense cortical layer (compact
bone) comprises the outer region of long bones. The major structural component of compact bone is called an osteon.
Composed of a concentric lamellar matrix, osteons create cylindrical conduits known as Haversian canals, which provide
access for the circulatory and nervous systems. The capillaries within the Haversian canals originate from arteries and
veins within the marrow cavity. The Volkmann's canals, which are transverse to the Haversian system, provide the pathway
for bone nourishment. With the exception of articulating surfaces, cortical bone is surrounded by a thin connective tissue,
the periosteum, which consists primarily of a collagen-rich fibrous layer and osteoprogenitor cells.
and result in bone erosion. Finally, the biomaterial should withstand standard sterilization procedures
and have a long shelf life.
Several approaches, including nonpolymeric ones, have been investigated to replace injured or
degenerated bone. Titanium knee and hip implants and poly(lactide-co-glycolide) (PLGA) screws
have both experienced good success. These materials, however, do not have the same mechanical
properties as bone and as a result, cannot be used for long-term implants. For example, since they are
much stronger than bone, metallic implants tend to stress shield bone, which leads to the erosion of
the surrounding native bone. Over time, the implant loosens due in part to mechanical mismatch and
stress shielding, and causes not only pain and suffering for the patient, but also the need for
additional surgical procedures to replace the malfunctioning implant. Materials that have mechanical
properties closely matching those of native bone would eliminate the problems associated with stress
shielding and would provide enough strength to support the mechanical strains of weight bearing
while new tissue develops. In the search for such materials, polymers remain attractive candidates
since they can be designed to have a great variety of elastic moduli, including those close to cortical
bone (3±50 MPa).
While designing these materials, scientists and engineers cannot forget that upon implantation,
the surface of the polymers will be immediately exposed to a physiological environment. In vivo,
proteins will rapidly cover the surface of any material and tend to denature on the surface due to
thermodynamic driving forces. Following this phenomenon, we then have little control over the
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Fig. 6. (a) A pen drawing of the structure of trabecular or cancellous bone adapted from that given by Frank Netter [263].
This type of bone lies interior to the cortical layer in long bones bone. As the name implies, cancellous bone consists of a
lattice-like structure comprised of trabeculae, which are small spicules of bone surrounded by marrow. Within each
trabecula, osteocytes maintain bone matrix, osteoclasts degrade regions of existing structure, and osteoblasts (especially
those associated with the osteoid region) actively build new sections of bone into the interconnecting marrow spaces. (b) A
more detailed pen drawing of trabecula adapted from that given by Frank Netter [263]. The location of the osteocytes,
osteoblasts and osteoclasts can be seen within the trabecular structure.
properties of a surface. At the same time, however, we must be sure that the materials still are
conducive to cell binding, proliferation, and maintenance of cellular phenotype. Several studies
based on the chemistry of the surface of materials and cell adhesion have shown that surface
chemistry does play a role in cell adhesion. This interaction indicates that as we attempt to optimize
other design criteria, we also must properly design surface chemistry and/or incorporate bioactive
signals to influence the behavior of cells. Detailed results of surface modification studies will follow
below.
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187
Table 1
Some bone regeneration polymers and their propertiesa
Material
Degradable
Cancellous bone
Trabecular bone
PLA
PLGA
Poly(ortho-ester)
Polyphosphazene
Polyanhydride
PET
PET/HA
PLGA/Ca phosphate
PLA/Ca phosphate
PLA/HA
PolyactiveTM
DegraPolTM
Yes
Yes
Yes
Yes
Yes
Yes
Yes
No
No
Yes
Yes
Yes
Yes
Yes
a
Compressive Modulus
strength (MPa) (MPa)
5
NR
(bulk)
NR
(bulk)
60 20
(surface) 4±16
(surface) NR
(surface) NR
(bulk)
(bulk)
(bulk)
(bulk)
(bulk)
50 [114] (compressive)
50±100 [114] (Young's)
NR
0.5 (tensile) 2.4 [128] (Young's)
NR
NR
140±1400 [115] (tensile)
NR
NR
0.25
5 [116] (Young's)
NR
NR
30±1200 [117] (elastic)
320 60
NR
NR
6±9
NR
NR
Porous
(mm)
Support cell Processable
adhesion
(Moldable)
Yes
Yes
100±500
150±710
NR
160±200
NR
NR
NR
100±500
100±500
NR
NR
NR
Yes
Yes
Yes
Yes
Yes
Yes
Yes
No
Yes
Yes
Yes
NR
Yes
Yes
No
No
Yes
Yes
NR
NR
Yes
Yes
Yes
Yes
Yes
NR
Yes
Yes
NR indicates not reported.
This section will examine polymeric materials that are being utilized for bone regeneration.
Initially, we will examine examples involving the control of surface chemistry and morphology to
guide cell adhesion, morphology and differentiation. We will then look at research that attempts to
adapt materials approved by the US Food and Drug Administration (FDA). Finally, we will explore
some novel materials with mechanical and erosion properties favorable for bone regeneration.
4.1. Control of surface chemistry and morphology
Surface modification techniques have been actively studied for all aspects of tissue
regeneration. It may be possible to modify the surface of a material to force cell specific adhesion
while designing the bulk material characteristics not only to guide adherent cells along a path of
tissue regeneration, but also to provide the mechanical stability necessary for continued tissue use. In
one example, metal oxide surfaces were modified with the cell adhesive peptide RGD and
FHRRIKA, a heparin binding domain (HBD) from bone sialoprotein [118]. Table 2 lists the names
of the amino acids and their corresponding one-letter abbreviations. The surface was first modified
with an organosilane followed by coupling of maleiimide functionality for direct peptide attachment
through cysteine thiols. The group then coupled various ratios of RGD:HBD to the surface and
seeded rat calivary osteoblast-like cells on the materials. An RGD:HBD of 75:25 was ideal for
mineralization of the surface. Regarding cell adhesion, the a2b1 integrin receptor played a crucial
role in cell adhesion for seeding times lasting less than 30 min; avb1 was involved for longer time
Table 2
A list of amino acids and their one-letter abbreviations
Amino acid
Abbreviation Amino acid
Abbreviation Amino acid
Abbreviation Amino acid
Abbreviation
Alanine
Arginine
Asparagine
Aspartic acid
Cysteine
A
R
N
D
C
Q
E
G
H
I
L
K
M
F
P
S
T
W
Y
V
Glutamine
Glutamic acid
Glycine
Histidine
Isoleucine
Leucine
Lysine
Methionine
Phenylalanine
Proline
Serine
Threonine
Tryptophan
Tyrosine
Valine
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points. This difference in integrin involvement may signify a way to design surfaces for osteoblast
specific adhesion over controlled times. Many polymers contain functional groups along the
backbone that can be readily functionalized with small bioactive peptides, such as RGD. The
polymer backbone can then provide mechanical stability and the properties governing shape
designation, while the peptides supply bioactive signals for bone regeneration.
Researchers have found that the surface plays a large role not only in the initial adhesion, but
also in the activity and differentiation of osteoblasts. Yamamoto et al. looked at polyethylene
terapthalate (PET), acrylic acid/methacryloyloxyethyl phosphate modified PET surfaces, and the
latter surface coated with either collagen or hydroxyapatite (HA) [119]. Rat bone marrow cells were
seeded on each of these surfaces and cultured in vitro for 2 weeks. Imaging with TEM, the authors
saw little cell adhesion to the PET alone. However, both the phosphate containing polymer surface
and the collagen surface supported osteoblast adhesion and the formation of thin, dense layers of HA
juxtaposed to the substrate surface. The HA-coated surface also supported cell adhesion with about
10 times more HA deposition; in addition, the cells were cuboidal in morphology and appeared more
osteoclast like. HA is a crystalline form of calcium phosphate that is known to enhance bone
formation, and collagen is an adhesive protein. In this case, collagen did support adhesion of cells,
but enhanced bone mineral deposition required HA. Simply varying the chemical nature of a surface
affects adhesion and spreading of many cell types. Webb et al. coated glass surfaces with five
different functionalized silanes to determine the functional effect on cell adhesion, spreading, and
migration [120]. They chose thiol, oxidized thiol, quaternary amine, and methyl functionalized
surfaces. When MC3T3-E1 osteoblast-like cells were cultured on the surfaces, the thiol surface
supported the greatest amount of cell adhesion and spreading followed by the oxidized thiol
materials. Both materials, however, allowed the least amount of migration. This result shows a clear
inverse relationship between spreading and migration.
When designing new materials, we must consider the required degree of spreading for cell
function and the desired amount of migration to completely fill the materials with cells or
extracellular matrix molecules. The time period for cell spreading and migration also should be
compatible with the ingrowth of new blood vessels to support tissue survival. Although they may yet
succeed in achieving these objectives, surface chemistry modifications alone have not yet offered
completely satisfactory solutions. As a result, we must consider cell function and how to use material
design to influence cellular behavior. Deposition of HA demonstrates one way to ensure that some
osteoblasts synthesize mineralized bone. However, other features of the material, such as the
morphology of the surface and the overall structure, remain important factors that are also crucial
design considerations in any endeavor to regenerate bone.
By using surface roughened polystyrene strips, Hatano et al. isolated surface roughness as a
factor which influences osteoblastic differentiation markers [121]. They found that rough surfaces
were more conducive to osteoblastic differentiation than smooth surfaces. They also found that as the
roughness increased to 0.8 mm, the levels of several differentiation markers, including alkaline
phosphatase and osteocalcin, reached a maximum. At surface roughness values greater than 0.8 mm,
marker synthesis declined. Matsuzaka et al. made micro-patterned materials of polystyrene and
poly(lactide). The patterns contained ridges and groves of the same width with quadrants having 1, 2,
5, or 10 mm ridges and groves; each groove measured either 1 or 1.5 mm in depth. Rat bone marrow
cells were then cultured on these materials. The authors found better cell adhesion and higher HA
and alkaline phosphatase synthesis when cells were seeded on poly(lactide) substrates. In addition,
cells attached only to the ridges and not to the grooves when the ridges and grooves were 1±2 mm in
width. In contrast, both 5±10 mm ridges and grooves supported cell adhesion. Cells attached to PLA
surfaces with 1±2 mm width ridges and 1 mm deep wells produced the most mineralized extracellular
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
matrix. The depth of the ridges/grooves played no role in adhesion or cellular metabolic activity.
This work indicates that both substrate material and morphology can control cellular behavior.
Certainly, the native extracellular matrix has a specifically defined morphology and chemistry.
4.2. FDA approved materials
The FDA approved material poly(lactic-co-glycolic acid) (PLGA) can be readily processed into
three-dimensional porous structures. The processing involves formation of the PLGA foam in the
presence of salt which leave behind pores when it is leached from the system. Pore sizes of at least
100 mm are necessary for cell penetration. PLGA (75:25) scaffolds with average pore sizes of 150±
300 or 500±710 mm were seeded with rat calvarial osteoblasts. Osteoblasts continued to proliferate
for 56 days. At this time point, the cells had invaded the scaffolds to a depth of approximately
200 mm. The cells stained positively for alkaline phosphatase activity, and the scaffolds had been
partially mineralized. Similar results have been found with rat bone marrow stromal cells [122]. This
data indicates that for small defects, porous PLGA foams may be suitable for bone regeneration.
However, research has not yet shown if this material will be suitable for defects greater than 200 mm.
Brekke and Toth tried to consider several factors as they designed a new material for bone
regeneration [123]. They used a porous PLGA filled with 750,000 molecular weight hyaluronic acid
(HyA) and bone morphogenic protein-2 (BMP-2). PLGA filled many of the architectural
requirements of the material by mimicking cancellous bone microstructure, while the HyA provided
a viscous scaffold for cell ingrowth, vascular ingrowth and stimulation, and growth factor storage.
The BMP-2 acted as a stimulant for bone regeneration and was osteoconductive; however, pockets of
HyA did not contain adequate amounts of growth factor. As a result, adipose tissue formed in these
regions instead of bone tissue. Improved processing techniques to uniformly distribute growth
factors within HyA will be necessary to fully evaluate this material as a viable candidate for bone
regeneration.
Several synthetic and natural polymers have been studied in connection with bone regeneration.
Collagen-polyvinyl alcohol films were coated onto coverslips and crosslinked either hydrothermally
or via gluteraldehyde [124]. These films were then seeded with osteoblasts. Surprisingly, increased
collagen content caused decreased cell attachment and spreading. This effect likely resulted from the
conformation of the collagen during crosslinking. Free collagen is mobile and can expose different
regions of the collagen molecule. In contrast, crosslinked collagen suffers from decreased mobility
and potentially cryptic cell-binding motifs. It would be interesting to evaluate similar scaffolding
that had been crosslinked to a lesser degree and also to evaluate the scaffolding in three dimensions
since these materials may behave differently in a three-dimensional configuration. Several studies
have shown different biological behavior of the same cell types when studied on chemically similar
three-dimensional and two-dimensional materials.
Other possibilities of materials that attempt to induce bone formation include blends of
polymers with ceramic materials such as tricalcium phosphate (TCP). Tricalcium phosphate has
been known to enhance bone formation both in vivo and in vitro. Kikuchi et al. combined poly(lactic
acid) (PLA) with TCP. They found mechanical properties reasonable for bone regeneration (fracture
strength of approximately 50 MPa; Young's modulus of approximately 5.18 MPa; and three-point
bending of approximately 50 MPa). They also observed that the material was nontoxic in the
presence of MC3T3-E1 cells obtained from mouse osteoblasts [116].
Zhang and Ma have tried to improve the bone bonding ability of porous PLA scaffolds by
growing a layer of hydroxyapatite (HA) on the surface of the three-dimensional foams [125]. They
created a PLA scaffolding with a porosity of up to 95% by using a liquid±liquid extraction technique
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with DMSO. The scaffolds were then dipped into simulated body fluid and HA crystals formed on
the surface of the scaffolds. Crystals grew within all of the porous surfaces. These crystals may offer
an advantage to the material since many experiments have shown that HA improves bone bonding to
surfaces. Ma et al. also produced. PLLA/HA composites and cultured osteoblasts with the materials.
They found that osteoblasts migrated further into the composite materials and had higher survival
rates on the composites than on PLLA alone. In addition, bone specific markers were expressed in
higher amounts in the composite than in PLLA over a 6-week time period [126]. Murphy, Kohn and
Mooney take a similar crystallization approach to combining calcium phosphate with PLGA
scaffolding. They form the porous PLGA by a combined solvent casting and particulate leaching
process. The porous scaffolding is then placed in a simulated body fluid to encourage the deposition
of a mineral layer on the surface. After 16 days in the simulated body fluid, a mineral layer had
formed on the scaffolding, which increased the compressive modulus five fold (250 kPa) [127].
Although significant, this increase in modulus falls short of the known modulus of trabecular bone
[114].
In another method that takes advantage of the bone enhancing capabilities of calcium
phosphate, Daculsi et al. developed a biphasic calcium phosphate (BCP)/methylhydroxypropylcellulose composite as an injectable bone regeneration material. The ratio of BCP:polymer was 60:40.
Short time implantation in rabbits showed a decrease in density of the synthetic material surface with
increased bone formation and no observation of an inflammatory response. After 12 weeks, 79% of
the total surface area was occupied by newly formed bone and residual BCP grains. Due to its porous
nature, this bone substitute allowed cells injected within the composite to invade the scaffolding and
begin the process of bone regeneration. As the methylcellulose diffuses out of the material over time,
it leaves a space for the newly formed bone [128]. Initial studies of this material show promise and
bode well for future use of composite materials in bone regeneration applications.
Kasuga et al. has developed a HA fiber/PLA composite that shows improved modulus of
elasticity over PLA alone while maintaining a similar bending modulus. The fibers were 40±150 mm
long and 2±10 mm in diameter. During the studies, the composite materials showed an increase in the
modulus of elasticity as the total concentration of fibers increased. Beyond blends composed of 60%
fiber, however, the bending modulus begins to decrease, and brittle fractures can occur. The authors
measured a modulus of 10 GPa with 60% fibers in the composite, while PLA alone exhibited a
modulus of elasticity of 2±7 GPa. The composite modulus falls well within the range seen in cortical
bone (3±30 GPa) [129].
While continued work with FDA approved polymers remains critical and necessary as a path
toward rapid approval of improved implant materials, these materials may not contain the optimal
properties for tissue regeneration. In fact, years of research confirm this observation. There exists a
great need for the design and synthesis of new materials, which are not off-the-shelf materials, but
materials that have been engineered to meet the constraints imposed by the tissue of interest. In this
vein, several researchers are engineering novel materials for bone regeneration.
4.3. New materials
Currently, new, degradable polyurethane materials are being engineered for application in bone
regeneration. Saad et al. have synthesized a material called DegraPol-foam. This foam is based on a
a,o-dihydroxy-oligo[((R)-3-hydroxybuterate-co(R)-3-hydroxyvalerate)-block-ethylene glycol]-copolymer. They find that osteoblasts will grow on the polymer, remain differentiated, and proliferate
for up to 2 weeks. Degradation of the polymer results primarily in poly[(R)-3-hydroxybuteric acid]
(PHB) and lysine-methylester. These degradation products are phagocytosed by macrophages and
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
osteoblasts, to some extent, and show no toxicity [116]. Further studies of the degradation products
with osteoblasts confirmed the lack of toxicity. The authors found no change in type I collagen or
osteocalcin synthesis (two osteoblast markers) after 32 days of exposure to the PHB; they did find a
time and dose dependent change in alkaline phosphatase activity. A maximal response of alkaline
phosphatase activity occurred after 4 days exposure to 2 mg PHB [130]. These results not only
suggest the usefulness of this type of material for bone regeneration, but also indicate that the
degradation product PHB may actually stimulate the osteoblasts in a positive manner.
Another promising new material is Polyactive, a polyethylene oxide-co-polybutylene
terephthalate co-polymer with bone bonding properties [131]. This material exhibits hydrogel
properties. Within confined volumes and in the presence of aqueous solution, it exerts a pressure on
the confining walls due to frustrated swelling. This pressure can reach as high as 2 MPa. Large, dry
cylinders of Polyactive were implanted, with a tight fit, into goat femora and allowed to swell; preswollen Polyactive was implanted into the contralateral femor. After 3, 9, and 25 weeks, histological
data showed good material-bone contact for the dry implants, while the pre-swollen implants had a
layer of soft tissue between themselves and the bone. Improved contact between the dry implants and
bone is attributed to the increased swelling pressure exhibited by the polymer, which is more likely
to cause the formation of tight junctions between the swelling polymer and the native tissue [131]. In
addition, this pressure may mimic the pressure exhibited in native bone tissue.
4.4. Novel surface eroding polymers
Materials with surface eroding properties as opposed to bulk degrading polymers have been
characterized frequently as biomaterials. One class of surface degrading polymers consists of
poly(ortho-esters). An advantage of these polymers is that as new tissue grows into the space
occupied by the artificial scaffolding, only the surface of the polymer scaffold degrades, leaving the
bulk of the material with its original mechanical integrity. With respect to bone, this may mean the
difference between catastrophic failure of an implant and successful regeneration of the bone.
Andriano et al. have studied the degradation properties of 50:50 PLGA scaffolding and a poly(orthoester) scaffolding coated with gelatin by bathing the materials within saline over a 6-week period.
The PLGA scaffolding lost only approximately 5% of its weight after 3 weeks, but had lost
approximately 37% of its weight after 6 weeks. The poly(ortho-ester) eroded in a more linear fashion
and had lost about 43% of its weight after 6 weeks. In addition, in vivo results showed that the bone
mineral density of the surface eroding polymers was 25% higher than that of the 50:50 PLGA at 6
and 12 weeks post-implantation in rabbit calvarial defects [132]. Solheim et al. characterized the
inflammatory response and effect of bone regeneration of both poly(D,L-lactic acid) and poly(orthoester). They implanted poly(ortho-ester) or PLA containing demineralized bone particles or
demineralized bone alone into the abdominal muscle of male Wistar rats. The bone particles
contained growth factors for bone regeneration while the polymer supported shape, particle
retention, and sustained release of the bioactive compound. They found that poly(ortho-ester)
showed no inflammation and had little to no effect on bone formation; PLA provoked a chronic
inflammatory response and inhibited bone formation [133].
Polyphosphazenes are also surface degradable materials that have been investigated for bone
regeneration. Ethyl glycinato p-methylohenoxy modified phosphazenes were used to synthesize twodimensional and highly porous (160±200 mm) three-dimensional scaffolds [134]. MC3T3-E1 cells
were cultured on both scaffolds. Cells grown on three-dimensional scaffolds adhered at higher
numbers and proliferated more rapidly than cells grown on two-dimensional surfaces. In fact, by day
21, they had proliferated throughout the matrix. These materials need further characterization,
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especially in the areas of mechanical properties and degradation rates, in order to ascertain whether
or not they may serve as viable candidates for bone regeneration.
Other new surface eroding materials include a class of poly(anhydride-co-imide). These
materials degrade into biocompatible products and provide favorable mechanical properties; the
polyanhydrides have well defined surface eroding properties while the polyimides have high strength
and rigidity. Attawia et al. synthesized poly[pyromellitylimidoalanine (PMA-ala):1,6-bis(carboxyphenoxy)hexane (CPH)] and studied osteoblast adhesion, proliferation and differentiation over a
21-day period. They found that the rat calivary osteoblasts behaved similarly on the PMA-ala-coCPH as they did on tissue culture polystyrene with respect to adhesion, proliferation and
differentiation over the entire 21-day period. The cells stained positively for alkaline phosphatase,
osteocalcin (an extracellular matrix protein of bone) and osteopontin (a cellular adhesion protein).
This new class of materials can be designed to have compressive moduli over the range of 10±
60 MPa, which is well within the range of calcinous bone [135]. Muggli et al. have also worked on
degradable anhydride-based materials for bone regeneration. They synthesized a series of polymers
with varying ratios of sebasic acid (MSA) and 1,6-bis(carboxyphenoxy) hexane (CPH). By varying
the ratio of the monomers they could control both crosslink density and degradation rates. For
example, they synthesized materials with degradation rates ranging from 2 days to 1 year and
mechanical properties ranging from 1.4 to 0.14 GPa. In general, the mechanical properties of the
crosslinked polymer are similar to those of both cortical and trabecular bone. More importantly, due
to their surface eroding nature, these polymers retain 70% of their mechanical integrity when 50%
on the materials has eroded away [115]. As a result of this characteristic, these polymers are likely to
maintain required mechanical properties while degrading at a rate comparable to bone regeneration.
Polyanhydride esters based on siacylic acid monomers also have sparked interest as a potential
material for bone regeneration since the siacylic acid degradation product can be osteoconductive in
vivo [136]. This work demonstrates the possibility of using the degradation products of novel surface
eroding polyanhydrides to further induce bone formation.
4.5. Growth factor encapsulation
In addition to the design and synthesis of new materials and further study of existing materials,
researchers are attempting to alter a material's ability to regenerate bone tissue by including
bioactive molecules. One class of biologically active molecules, which has proven useful in bone
regeneration, involves the bone morphogenic proteins or BMPs. BMPs are active in other tissue
regeneration as well, but some of these compounds, such as BMP-2, are active in bone formation.
Like many biomolecules, BMP-2 exists commercially in recombinant form, making it a potentially
highly useful agent in tissue regeneration. Winn et al. sought to take advantage of the potency of
BMP-2 by combining it with porous PLA coated with type I collagen. These scaffolds were either
preseeded with osteoblast precursor cells or implanted in Harlan nude athalamic rats without cells.
Scaffoldings with BMP-2 significantly induced more bone production than scaffoldings without
BMP-2 [137]. These results are encouraging because they suggest the possibility of regenerating
bone without having to harvest or expand osteoblast populations from the patient. Peter et al. showed
that another growth factor may improve the success of bone regeneration. They entrapped
transforming growth factor-b1 (TGF-b1) in PLGA microspheres containing 5% polyethylene glycol
and studied the effect of the release of the TGF-b1 on osteoblasts seeded on poly(propylene
fumerate). They found that TGF-b1 significantly increased the proliferation of osteoblasts, alkaline
phosphatase activity, and osteocalcin production after 21 days in culture in comparison to cells
seeded on poly(propylene fumerate) in the absence of TGF-b1 [138].
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
It may be possible to use some of the newly designed polymers in combination with bioactive
molecules, e.g. BMP-2 and calcium phosphate derivatives, to produce scaffolding materials that will
provide all of the requirements mentioned at the outset of this section. These requirements include,
among others, mechanical attributes similar to those of native bone, pliability to fit the defect, and
regenerative properties. Certainly, the mechanical properties obtained with some of the surface
eroding polymers show considerable promise, as they are similar to native bone. In addition, using
surface erosion as the primary mode of degradation may allow the synthetic bone scaffolding to
maintain mechanical integrity until sufficient natural bone has replaced the defect.
Additional reviews on polymeric biomaterials for bone regeneration are available. A recent
review by Hutmacher details polymeric biomaterials and discusses in vitro bioreactors for proper
tissue development. The review also contains a table showing time dependent polymer degradation
profiles as well as the amount of time that a polymer will maintain mechanical properties in vivo. The
actual mechanical properties, however, are not given [139]. Middleton and Tipton recently reviewed
the use of poly(lactide), poly(glycolide) and polycaprolactone and their blends as orthopedic
biomaterials. Their review includes mechanical properties, degradation rates and commercial uses of
these polymers [140]. Temenoff and Mikos review injectable polymers and ceramics for bone and
cartilage regeneration. They include mechanical properties, chemistries and modes of crosslinking,
and they also discuss regeneration efficacy of the materials [141]. These reviews were published
subsequently to the writing of this review. This review attempts to compliment the information given
in the above-mentioned articles, and does not attempt to include all information regarding bone
regeneration, especially data concerning ceramic and current commercially available materials.
5. Vascular grafts
The cardiovascular system in humans is comprised of the heart, blood vessels, and blood. The
vascular system consists of arteries, arterioles, veins, venuoles, and capillaries. Arteries transport
blood from the heart to the organs where the arteries branch into arterioles. Arterioles further branch
into capillaries, which infiltrate the tissues and allow exchange of nutrients and waste. The
capillaries then converge to form venules, which further coalesce into veins. The artery remains the
most commonly diseased vessel, and it is the diseased artery that leads to the requirement of bypass
surgery. Fig. 7 shows the structure of an artery. The wall of the artery contains three layers:
adventitia, media, and intima. The outermost layer is the adventitia and is composed of collagen rich
connective tissue containing few elastic fibers. The middle layer is the media, which consists of
smooth muscle, arranged in circumferential layers, and more elastic fibers. Proliferation of these
smooth muscle cells can result in intimal hyperplasia or narrowing of the artery. The innermost layer
(intima) surrounds the lumen and consists of a basal lamina upon which sits a monolayer of
endothelial cells [4,113]. This complex structure provides a blood conduit that, in a healthy state,
inhibits intimal hyperplasia, thrombosis, and atherosclerosis.
Atherosclerotic cardiovascular disease remains the number one cause of death in the western
world. This disease causes localized reduction in blood flow through arteries (stenosis), and
ultimately stops blood flow entirely through the affected vessel. Treatment consists of bypassing the
affected area using an artificial graft or saphenous vein [142]. Currently, most small-diameter graft
procedures, less than 6 mm in diameter, are performed using saphenous vein, an autologous native
blood vessel. The sustained function of these natural grafts is attributed to the presence of viable
endothelial cells on the luminal surface. Due to the nature of vascular disease, however, patients in
need of grafts often do not have veins suitable for grafting.
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Fig. 7. A pen drawing of the structure of a muscular artery adapted from [264]. The wall of the artery is comprised of three
layers: advantitia, media and intima. The outermost is the adventitia, composed of collagen rich connective tissue
containing few elastic fibers. The middle layer is the media consisting of smooth muscle, arranged in circumferential
layers, and more elastic fibers. Proliferation of these smooth muscle cells can result in intimal hyperplasia or narrowing of
the artery. The innermost layer (intima) consists of a basal lamina upon which sits a monolayer of endothelial cells. Interior
to the intima is a layer of elastic tissue called the internal elastic lamina [4,113]. This complex structure provides a blood
conduit that in its healthy state, inhibits intimal hyperplasia, thrombosis and atherosclerosis.
Vascular grafts, both small-diameter (less than 6 mm) and large-diameter, are used as a
treatment for vascular disease. Current artificial vascular grafts are made of Dacron (poly(ethylene
terephthalate)) or expanded polytetrafluoroethylene (ePTFE). Overall, an estimated 550,000 vascular
grafts are implanted annually in the US alone. The patency of large-diameter grafts, while
satisfactory, still leaves room for improvement. Small-diameter grafts, however, remain patent only
15±30% of the time after 5 years. This low patency falls well below the accepted average for medical
procedures and continues to drive research efforts aimed at improving grafts [143]. The decreased
tolerance of small-diameter grafts to thrombosis and impaired flow forces scientists and engineers to
use new materials and make inventive improvements to standard materials. A review of past and
present materials used for vascular grafts is presented by Greenwald and Berry and will not be
repeated here [144]. Moukwa presents a discussion of polymer-based biomaterials and their
development since the 1920s [145]. In addition to new, innovative materials, improvements to
current small-diameter vascular grafts include endothelial cell seeding, surface treatments with
antithrombotic fluoropolymers, and adhesion promoting peptides.
5.1. Inflammatory response and wound healing
Numerous factors can affect the biocompatibility of vascular grafts. One of these involves the
inflammatory response provoked by an implanted material. Polymeric materials used in vascular
graft applications exacerbate this inflammatory response, the degree of which depends on both the
type of tissue surrounding the implant and the properties inherent to a specific implant material.
Stooker et al. presented a good overview of the wound healing process after graft implantation and
concluded that the healing process relied heavily on a well-balanced interaction between endothelial
and smooth muscle cells [146]. Urayama et al. offered an interesting overview of healing with their
analysis of 14 explanted grafts from humans. For Dacron grafts, at 5 and 24 days after implantation,
a thin layer of thrombi, with no smooth muscle cells, covered portions of the luminal surface. After
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11±148 months of implantation, a single layer of endothelial cells covered the graft near the
anastomosis. The rest of the luminal surface was covered by connective tissue matrix containing
collagen fibers (mostly type III). For ePTFE grafts, the luminal surface contained a thin layer of
thrombi at 4 and 7 months following implantation. After 42±86 months, a thin layer of fibrin covered
the luminal surface, and a small amount of collagenous matrix containing mostly type III collagen
was also present. After comparing the grafts, the authors did not observe any significant difference in
healing between Dacron and ePTFE [147]. In addition, none of the grafts indicated regeneration of
healthy vascular tissue. One thought as to how to achieve formation of normal vascular tissue
focuses on the neovascularization of the graft.
Neovascularization of the wound site is important to healing, cell growth, nutrient delivery, and
waste removal. In the early stages after implantation, monocytes migrate from the vasculature and
subsequently develop into macrophages; the mechanisms supporting the macrophage response are
still unclear. Hagerty et al. investigated the types of macrophages in the inflammatory response
provoked by the implantation of vascular grafts in subcutaneous and adipose tissue environments.
This study showed that over 70% of macrophages present in both graft types after 3 and 5 weeks
stained positively for the ED1 antigen, indicating recent arrival from the vasculature. The data also
supported the hypothesis that macrophages actively proliferated [148]. Salzmann et al. analyzed the
inflammatory response and neovascularization of some common polymeric graft materials in a rat
model. Subcutaneous implants suffered from thicker and less cellular fibrous capsules. The material
porosity also influenced encapsulation, with more porous surfaces yielding less encapsulation. In this
experiment, the tested ePTFE materials proved less inflammatory than the Dacron materials.
Furthermore, an inverse relationship seemed to exist between the inflammatory response and
neovascularization, which indicated that the control of inflammation could increase angiogenesis
[149].
Additional research, conducted by Greisler, also demonstrated significant differences in
vascular healing and cell proliferation based on polymer composition. After 4 weeks of implantation,
Dacron prostheses showed an inner capsule of fibrin coagulum containing no mesenchymal cells. In
contrast, poly(glycolic acid) (PGA) vascular prostheses contained a highly cellular inner capsule,
consisting of longitudinally and circumferentially oriented myofibroblasts, beneath a monolayer of
factor VIII positive endothelial cells at the blood-contacting surface. The more extensive tissue
ingrowth observed in PGA grafts was attributed to differential activation of macrophages by
implanted lactide/glycolide copolymers. Capillary infiltration was noted in all lactide/glycolide
copolymers grafts, but was not seen in Dacron prostheses [143,150].
Neointima formation commonly occurs in the midgraft region when endothelial cells are
transplanted onto the luminal surface. Kleinert et al. presented a quantitative morphological analysis
of cell density in the neointima observed after implantation of endothelial cell sodded ePTFE grafts.
Unlike the intima of native artery, the formed neointima consists of an endothelial cell monolayer as
well as a sub-endothelial layer of cells that varies in thickness. The results of the analysis showed
that the neointima experienced a thickening from 3 to 12 weeks post-implantation and then
regressed. After 52 weeks, intimal thickness was approximately equal to values recorded at 3 weeks.
The average cell density within the neointima remained almost uniform over the 52-week
implantation time, and fell within a normal range observed for natural arteries [151].
5.2. Mechanical properties
This work will not attempt to provide a complete overview of vascular mechanics since
Greenwald and Berry have presented a concise review regarding mechanical properties of
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large-diameter vascular prostheses. A number of different materials, including Dacron and PTFE,
were discussed [144]. However, some of the essential parameters will be addressed. Compliance
mismatch between synthetic grafts and natural arteries alters the anastomotic geometry and can lead
to both flow disturbances and separation zones. Diameter mismatch is also important since the
maintenance of wall shear stress levels can only occur within properly sized grafts. This type of
mismatch causes hemodynamic disturbances that are thought to affect intimal hyperplasia formation
near the anastomosis. Weston et al. studied diameter mismatch using a model that allowed alterations
in the phase angle between the pressure and flow waves. Three different end-to-end anastomosis
models were studied: two where the graft was smaller than the artery by 6 and 16%, and one where
the graft diameter exceeded that of the artery by 13%. In the undersized grafts, wall shear rates near
the proximal anastomosis increased slightly and had minimum values near the distal anastomosis.
For the oversized graft, the minimum shear rate occurred near the proximal anastomosis. Intimal
thickening correlated with low mean shear rates at arterial branch points and at end-to-side
anastomoses. This study illustrated the effects of diameter mismatch and emphasized the importance
of diameter and compliance matching between native artery and synthetic grafts during vascular
surgery [152].
Hsu and Kambie, using canine carotid artery and polyurethane grafts, conducted another study
of compliance mismatch. They postulated that the pressure dependent elasticity of arteries resulted
from the changing ratios of elastin to collagen in the native vessel, a feature not yet achievable in
artificial prostheses. The tested grafts were isotropic, in contrast to anisotropic arteries, which offer
more extensibility in the longitudinal direction. These intrinsic differences in moduli make complete
compliance matching very difficult; however, the proper design of synthetic grafts still should
consider pressure dependence, frequency dependence, and anisotropy [153].
5.3. Endothelial cells
Seeding of synthetic vascular grafts with autologous endothelial cells (ECs), the cells which
naturally line the vasculature, prior to implantation significantly improves small-diameter graft
survival. This procedure, however, proves difficult to implement and involves major obstacles such
as the source for endothelial cells, the efficiency of cell seeding, and the detachment of adherent cells
after implantation. Williams provides a good review of these issues and mentions that the primary
sources of ECs for transplantation include vein-derived cells and adipose tissue. He also presents a
summary of pre-clinical animal trials and human studies involving endothelial cell transplantation.
In both types of studies, the seeding of endothelial cells on vascular prostheses improved the healing
and patency of the grafts. For further details, we direct the reader to this paper [154].
Research regarding the efficacy of cell seeding continuously aims to improve the adhesion and
proliferation of endothelial cells. Imbert et al. compared the proliferation of primary human
umbilical vein endothelial cells (HUVECs) and the endothelial cell line EC-RF24 on several
polymer surfaces coated with different concentrations of fibronectin. The vascular graft materials,
Dacron and Teflon were among the surfaces studied. HUVECs and EC-RF24 cells displayed
different growth behavior on all surfaces tested. The fibronectin concentration did not affect cellular
proliferation on tissue culture plate surfaces (tissue culture polystyrene and Permanox). EC-RF24
monolayers contained much greater cell numbers than HUVEC monolayers. This result was
attributed to the incorporation of E6/E7 DNA into the genome of the HUVECs, which resulted in the
disturbance of normal cell cycle control. Differences in cell morphology were attributed to
differences in the deposition of matrix proteins and/or synthesis of growth factors [155]. The effect
of different polymers, surface structure, and surface treatments on endothelial cell behavior remains
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under investigation. Marois et al. studied these differences in behavior and found that material
porosity played an important factor in cell adhesion and migration. An increase in porosity resulted
in increased adhesion and migration, while polyester grafts with a fluoropolymer treatment showed
lower cell adhesion and migration. This study reinforced the important relationship between graft
material selection and the success of endothelial cell coverage [156].
Another important property for endothelial cells is their coagulant function. ECs can exhibit
either anticoagulant or procoagulant functionality depending upon external stimuli. Kottke-Marchant
et al. prepared polymer surfaces of varying hydrophilicity through radiofrequency plasma
polymerization and then cultured human aortic endothelial cells on their surfaces. The amount of
von Willdebrand's factor (vWF) released was measured since a decrease in vWF production
indicates a less thrombogenic state of endothelium. Increased cell growth and spreading correlated
with increased surface hydrophilicity. In comparison, normalized vWF production decreased on the
more hydrophilic surfaces. This study suggested that hydrophilic surfaces allow for improved EC
seeding [157].
The above studies indicate that, as with other cell and tissue types, surface chemistry plays a
role in cell attachment, proliferation, and differentiation. The cellular responses may result from
functional groups at the surface or from surface hydrophobicity. Additionally, the adhesion of serum
proteins to the surface likely plays an intermediate role in controlling cellular response. The studies
also demonstrate that material morphology remains an important parameter. We again see that
increased porosity seems to improve the desired biological response. Regardless, it remains clear that
the surface of a material plays a role in EC seeding.
A different way of seeding vascular grafts is to produce confluent cell layers by culturing ECs
on fibronectin-coated surfaces. The layers are cultured to confluence over 48 h prior to use and
remain intact for up to 24 h when subsequently exposed to a shear stress of 25 dyn/cm2. PooleWarren et al. used this method to coat ePTFE and microporous polyurethane (PU) grafts (4 mm
diameter). They loaded fibronectin (FN) onto the grafts by physical adsorption or covalent binding,
and then seeded and grew ECs to confluence. After implantation in the ovine carotid interposition
model, the grafts were explanted at one, 3 and 6 weeks. Stable confluent monolayers of ECs were
observed on both pre-coated-ePTFE and pre-coated-PU grafts. EC coverage was not different
between the covalently bound and physically adsorbed fibronectin-coated grafts, however, the
amount of associated thrombus appeared lower on the graft containing covalently bound FN.
Although they achieved positive results, the authors concluded that the patency of small-diameter
arterial reconstructions using synthetic materials (even those seeded with ECs) remained lower than
that observed with autologous arteries. This study clearly demonstrates the challenge of translating
in vitro findings into in vivo systems [158].
Another seeding technique uses autologous bone marrow. Bone marrow contains active,
undifferentiated mesenchymal cells that can differentiate into various cell types, including
endothelial cells, and produce many kinds of cytokines within the implant environment. Since it
becomes trapped inside the graft wall, the bone marrow does not affect mechanical properties or
handling. Noishiki et al. considered the benefits of coating ePTFE grafts with autologous bone
marrow and evaluated the grafts in canine abdominal aortas. The treated grafts showed rapid and
uniform neointima formation and remained patent throughout a 3-month implantation. The authors
did not observe intimal hyperplasia at the distal anastomotic sites, but did see a thin coating of
endothelial cells along the entire graft length [159]. Another study using bone marrow was presented
by Bhattacharya et al. and used CD34‡ bone marrow cells to seed poly(ethylene terephthalate)
(PET) grafts. A recent in vitro experiment showed that pluripotent CD34‡ cells, contained within
bone marrow, could differentiate into endothelial cells. Grafts treated with these cells were implanted
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into the descending thoracic aorta of dogs and explanted after 4 weeks. The treated grafts
experienced endothelial-like cell coverage and enhanced microvessel formation. As improved
labeling techniques become available, isolated and labeled CD34‡ cells could provide definitive
evidence that endothelialization can originate from transplanted CD34‡ cells [160].
Another method of applying endothelial cells to graft surfaces involves microvascular
endothelial cell sodding. In this technique, canine falciform ligament or human adipose tissue
microvessel endothelium is applied to the luminal surface of a prosthetic graft immediately before
implantation. As a result, this method uses a large number of cells to coat the graft surface. Phillips
et al. described the successful sodding of microvascular-derived endothelium in a prosthetic graft
applied to the canine coronary circulation. At explant, they observed a confluent cellular layer
resembling endothelium. In this study, sodding did not improve early graft patency (21 days),
however, other studies have shown long-term improvement using these methods [161].
Significant evidence exists that novel polymers, designed to induce monolayer EC coverage of
the lumen of the vascular graft, can improve graft patency. At the same time, the design of new
materials cannot rely exclusively on this procedure, but needs to include other important parameters.
5.4. Coatings for existing vascular grafts
5.4.1. Growth factors
Local release of basic fibroblast growth factor (bFGF) can enhance the growth of endothelial
cells and increase HUVEC proliferation. Wissink et al. found that increased amounts of immobilized
heparin on an N-(3-dimethylaminopropyl)-N0 -ethylcarbodiimide (EDC)/N-hydroxysuccinimide
(NHS) crosslinked collagen substrate allowed greater bFGF binding. By elevating the ratio of
EDC to activated heparin carboxylic acid groups, the surface could immobilize more heparin. These
heparinized surfaces offered a longer sustained release of bFGF compared to the nonheparinized
substrates [162]. This research was further extended to study how the immobilization of heparin to
EDC/NHS-crosslinked collagen affects the proliferation of HUVECs. On these substrates, bFGF
loading enhanced HUVEC proliferation, where the largest amount of surface-bound bFGF resulted
in maximum proliferation. Thus, binding bFGF to heparin did not seem to inhibit the bioactivity of
the growth factor [163].
Another study demonstrated the possibility of directly adsorbing bFGF onto Dacron vascular
grafts. During in vitro studies, 40% of the bFGF was released in the first 24 h. After the first day, the
release slowed down, and the amount of bFGF remaining on the graft decreased over 2 weeks. From
in vivo studies in dogs, bFGF also induced fibroblast migration and proliferation, which may not
represent a desirable response as it can result in occlusion of the graft. These experiments do,
however, illustrate that the growth factor retains its biological activity when adhered directly to a
surface [164].
Greisler also used growth factors impregnated into either Dacron or ePTFE vascular grafts to
stimulate endothelial cell proliferation; however, he concentrated on acidic fibroblast growth factor
(aFGF) due to its strong endothelial cell mitogenic activity and its weak effect on proliferating
smooth muscle cells. In vitro studies, incorporating aFGF and heparin suspension within a
fibrinogen/thrombin matrix, showed a dose dependent response and indicated the capacity of aFGF
to support the growth of HUVECs. Also presented were in vivo release kinetic studies involving
ePTFE grafts impregnated with the matrix within a New Zealand white rabbit infrarenal aorta model.
These experiments showed that variations in wall thickness and/or the diameter of void spaces within
the wall can control release kinetics. After a 24 h infusion of aFGF into rabbit aortas, the in vivo
growth factor distribution was measured. Thyroid tissue contained the highest concentrations of
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aFGF; the lungs, liver, kidneys, spleen, and heart contained intermediate levels; blood, muscle, and
fat revealed the lowest concentrations. The treated grafts healed better and showed a confluent
endothelialized surface and extensive transmural capillary infiltration after 28 days in the canine
aorto-iliac. These experiments were repeated with longer grafts and extended implant times with
similar results. The present challenge undertaken by Greisler aims to minimize sub-endothelial
intimal thickening. In vitro studies showed that in the presence of high aFGF concentrations, the
addition of 500 U/ml heparin completely eliminated smooth muscle cell proliferation while
stimulating endothelial cell proliferation. Although promising, this technique requires in vivo
analysis to develop further conclusions [143,150].
The above studies demonstrate that the controlled release of growth factors from current
artificial graft materials such as Dacron and ePTFE can stimulate the formation of EC monolayers.
This effect provides another design parameter that may prove necessary when designing successful
polymeric vascular grafts.
5.4.2. Protein and peptide coatings
Activated smooth muscle cells (SMCs) change their phenotype from contractile to synthetic and
then migrate and proliferate. The biopolymer elastin helps to maintain the phenotype of SMCs. In an
in vitro environment, type I collagen gels containing a-elastin inhibit the migration of smooth
muscle cells in a dose dependent manner. Ito et al. used a-elastin as a soluble elastin, which exhibits
coacervation under the appropriate conditions. Coacervation is an important step in the conversion of
proelastin to elastin fiber in vivo. Once coacervated, elastin inhibits SMC proliferation. In vitro
studies showed that SMC migration was prevented by a-elastin crosslinked with glycerol
poly(glycidyl ether), whereas migration of ECs did not change. In addition, the experiment
demonstrated the feasibility of coating a Dacron graft with crosslinked elastin [165].
Adhesive peptides, namely YIGSR and RGD, have been shown to improve endothelial cell
adhesion and can influence cellular migration. Kouvroukoglou et al. focused on elucidating the
effect of modifying surfaces with YIGSR and RGD on the migration behavior of ECs. To investigate
this phenomenon, they used an improved migration assay combining spatially enhanced video
microscopy and digital time-lapse recording to track and analyze the movement of a large number of
cells. They found that, rather than affecting migration speed, adhesive peptides increased the
persistence of cell motion and enhanced the random motility coefficient of endothelial cells. Cell
colony growth on glass modified with YIGSR peptides in the presence of soluble basic growth factor
was significantly enhanced [166]. Another study involving adhesive peptides used dialdehyde starch
(DAS)-coated polymer surfaces. The reactive aldehyde groups on the DAS surface coupled with the
RGD-containing synthetic peptide amine groups to form stable carbon±nitrogen bonds. The
synthetic peptide used by Holland et al. to create a biologically active surface was the synthetic
peptide GRGDSPK. Consistent with other findings, the RGD containing peptide had a major and
direct impact on endothelial cell adhesion. Cells on the modified substrate spread at a faster rate than
on polymer or fibronectin-coated polymer control surfaces [167].
Another peptide sequence derived from the 33/66 kDa carboxy-terminal heparin-binding
domains of the fibronectin molecule (referred to as FN-C/H-V) promoted the adhesion and
proliferation of vascular endothelial cells. The structure of this sequence was WQPPRARI. Huebsch
et al. generated a photoreactive analog of FN-C/H-V (ASD-V) and covalently linked it to
polystyrene (PS) and PET films. Both modified surfaces encouraged EC adhesion and spreading,
indicating that the modified FN-C/H-V retained its biological activity [168].
Chinn et al. modified the surface of a poly(ethylene terephthalate) (PET) graft using either a
fluoropolymer (chemical modification) or an RGD-containing peptide and evaluated the grafts in
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different animal models. In animals anticoagulated with heparin, platelet accumulation on modified
surfaces was less than on unmodified grafts. Overall, the fluoropolymer modification exhibited less
platelet accumulation and inflammatory response than surfaces coated with the RGD-containing
peptide [169].
5.5. Biological sealants
The porosity of vascular grafts is important to their long-term patency since materials with
pores allow the ingrowth of tissue. One disadvantage to high porosity, however, is the high
permeability of the graft to blood during implantation, which results in severe blood leakage through
the graft walls. In general, pre-clotting can seal highly porous prostheses. This technique illuminates
another disadvantage of porous grafts since pre-clotting precludes the administration of large doses
of anticoagulants because of the inherent risk of dissolving the fibrin network that seals the graft.
Biological sealants, other than fibrin, eliminate the need for lengthy pre-clotting of vascular grafts
prior to implantation. Until fully degraded, sealants, in general, slow healing and tissue ingrowth;
however, normal healing occurs following degradation. No long-term detrimental effects on healing
have been found. Lee et al. found that an alginate-coated Dacron vascular graft was impermeable to
blood and did not significantly change the mechanical properties of the graft. Three months postimplantation, the alginate had completely resorbed. This study suggested the feasibility of using
biodegradable alginate as a sealant [170]. Another example of a biological sealant is the
carbodiimide crosslinked gelatin-impregnated knitted polyester prosthesis (Uni-graft1). Evaluation
of this prosthesis was performed in the thoraco-abdominal bypass within the canine model for
durations ranging from 4 h to 6 months. Upon implantation, the graft was immediately impervious to
blood flow and exhibited low surface thrombogenicity. The degradation of the crosslinked gelatin
occurred through the lysis of peptide bonds without the release of crosslinking agent. As a result, the
authors did not observe an adverse cellular reaction; rather, they noticed that once the gelatin
resorbed, the healing process proceeded in the usual fashion [171].
Due to its greater availability, adipose tissue also has been explored as a sealant. Matsumoto
et al. clotted a highly porous vascular prosthesis with autologous adipose tissue harvested from the
lower abdominal subcutaneous layer. After several series of animal experiments, this procedure
became documented as both safe and reliable, and in this study, was performed on 36 human patients
randomly selected at the Yokohama City University Hospital. Thirty-five of these patients had
arteriosclerosis and one had an aneurysm of the common iliac artery. Several types of grafts were
used including the Micron, the Sauvage EXS, and the MILLIKNIT. Each surgery began with the
removal of the adipose tissue followed by graft preparation, which took place after exposing the
occluded artery. After surgery, occlusion of the graft occurred in three cases, each of which was not
related to the diameter or type of graft. The overall patency rate for this trial (spanning 274 190
days) reached 91.0%. Postoperative angiography showed neither stenotic changes at the anastomotic
sites nor irregularities of the inner surface of the prosthetic grafts due to exposure of adipose tissue
embedded in the prostheses wall [172].
Noishiki et al. evaluated Dacron vascular prostheses (6 mm i.d.) coated with succinylated,
thermally crosslinked collagen (SC) in the abdominal aorta of dogs. After 3 weeks, the SC grafts
showed neointima formation throughout the graft with endothelial cell lining, abundant fibroblast
infiltration, and capillary blood vessel ingrowth without foreign body reactions. This result showed
that uncontaminated collagen could support natural wound healing [173].
Another graft currently used in Europe and under clinical trials in the US is the DeBakey
albumin-coated graft. The use of albumin reduces blood loss during surgery, but healing does not
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begin until the crosslinked albumin resorbs. Coating knitted polyester grafts with gluteraldehyde
crosslinked human albumin demonstrated short-term anti-thrombogenic and hemocompatible
properties that were superior to those observed for pre-clotted polyester grafts. Marois et al.
evaluated the use of this graft in the thoracoabdominal bypass canine model with a length of 30 cm.
This study confirmed the advantages of albumin since the albumin-coated graft prostheses did not
experience any blood loss. After a few weeks, healing proceeded in a fashion similar to that of preclotted grafts [174].
In all cases, normal healing occurred after resorption of the sealing agent. Sealing then,
improves handling by eliminating the need to preseal a graft in the operating room. As a result, this
technique offers a significant improvement to vascular graft technology by reducing surgical time.
However, since many existing sealants can elicit thrombogenic responses, they fail to improve the
patency of vascular grafts.
5.6. Plasma treatments
Plasma treatments offer a method of altering the surface characteristics of materials without
affecting the material's physical properties. Following treatment, a wide variety of chemicals can
then be used to incorporate specific functional groups into the substrate. Hsu and Chen used plasmainduced graft polymerization of L-lactide onto a polyurethane substrate. To verify the presence of Llactide groups, they presented a complete surface characterization including contact angle
measurements and ESCA survey scans. Enhanced adhesion of fibroblasts and endothelial cells
was observed on the modified surfaces. In contrast, these same surfaces experienced a decrease in
the adhesion of platelets. This finding represents a positive example of a grafting technique utilizing
biodegradable polyesters commonly used in tissue engineering [175].
Chandy et al. described the immobilization of bioactive molecules, including prostaglandin,
heparin and phosphatidyl choline, onto collagen and laminin modified Dacron and PTFE grafts.
After receiving plasma treatments, the grafts were dipped into collagen solutions in phosphatebuffered saline and then coated with a laminin layer. These modified surfaces offered an
environment conducive for binding various biomolecules. Platelet adhesion to the different
biomolecules was reduced significantly compared with untreated surfaces. In addition, the
immobilization of biomolecules onto the grafts substantially modified fibrinogen adsorption. More
detailed in vivo studies are needed, however, to further characterize the improvements [142].
Lin and Cooper used low density polyethylene (LDPE) tubing to show that the incorporation of
acidic sulfur-containing functionalities onto the inner surface of the tubing results in a more
thrombogenic surface. All of the inner surfaces were modified uniformly using SO2 plasma. The
authors used LDPE in this experiment only for its chemical simplicity since LDPE is not really
suitable for use as a vascular graft [176].
Tseng and Edelman applied amide and amine plasma treatments to improve endothelial cell
adherence to ePTFE vascular grafts and observed increased surface tension and enhanced endothelial
cell adhesion. They also noted that the surface functional groups resulting after plasma treatment
contained high levels of oxygen and nitrogen; both oxygen and nitrogen groups are known to
enhance cell adhesion. Arterial endothelial cells experienced shear stress caused by blood flow and
circumferential stretching forces normal to the vessel wall as a result of arterial expansion and
contraction. Under pulsatile flow conditions, significant cell detachment was not observed; however,
the authors did see a confluent monolayer of endothelial cells, which were not oriented with the
direction of flow [177]. Another group using ammonia plasma to modify surface properties is Chu
et al. In their studies, the addition of amine and amide groups on substrates promoted ionic bonding
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with mammalian cells. The substrates also became more hydrophilic after ammonia plasma
treatment. The authors used poly(L-lactic acid) (PLLA) to show improved HUVEC growth on the
plasma modified surfaces, and the addition of a fibronectin coat on the modified PLLA surface
further enhanced this improvement [178]. Ramires et al. also used ammonia plasma treatments
to improve HUVEC adhesion to a polymer substrate. In this example, PET surfaces were
plasma treated with ammonia and oxygen gas mixtures. The treated surfaces were not cytotoxic and
showed good biocompatibility. As in the above studies, plasma treatments improved HUVEC
proliferation [179].
5.7. rHir
Recombinant hirudin (rHir) is the most potent specific inhibitor of thrombin, blocking both the
proteolytic and fibrin binding sites of thrombin. Phaneuf et al. covalently immobilized the specific
antithrombin agent rHir to a sodium hydroxide hydrolysis modified Dacron surface. This method
should inhibit all enzymatic, chemotactic and mitogenic properties of thrombin. The authors
observed significantly greater thrombin inhibition and binding for the modified surfaces [180]. rHir
was then covalently bound to a novel small bore (4 mm i.d.) polyurethane graft with functional
groups as anchor sites for protein attachment (cPU). This method incorporated a polycarbonatebased urethane with the chain extender 2,2-bis(hydroxymethyl)-propionic acid to generate
carboxylic acid groups on the graft surface [181]. Canine serum albumin (CSA) was used as the
basecoat protein via a crosslinker, and the rHir was bound using sulfosuccinimidyl 4-(Nmaleimidomethyl) cyclohexane-1-carboxylate (sulfo-SMCC), resulting in cPU-CSA-SMCC-S-rHir.
This modified surface bound significantly more rHir than nonspecifically or covalently bound CSA
without a crosslinker. Furthermore, more thrombin attached to the rHir on modified surfaces than to
control surfaces [182].
5.8. Photocurable coatings
To enhance antithrombogenicity at the luminal surface and promote tissue regeneration at the
outer surface, Kito and Matsuda engineered two photocurable surfaces: chondroitin sulfate (CS) or
hyaluronic acid (HyA) for the luminal coating and gelatin for the outer coating. In vitro platelet and
endothelial cell adhesion experiments demonstrated good adhesion on photocured gelatin surfaces
and reduced adhesion on the photocured CS or HyA surfaces. This system was then tested in vivo
using 5 mm diameter grafts composed of either Dacron or polyurethane (PU). These grafts were
treated and implanted into the infrarenal abdominal aorta of mongrel dogs. Results indicated that the
photocurable CS and HA surfaces did not provide uniform coatings. Consequently, additional
investigations will attempt to develop further coating techniques. Improving these techniques
remains critical since nonuniform coatings lead to thrombus formation in some parts of the luminal
surface. Although the systems tested in this example have not yet been optimized, the concept of
using different artificial extracellular matrices for the luminal and outer surfaces of vascular grafts
proves interesting and merits further investigation [183].
5.9. Heparin
Heparin has been studied for its anticoagulant effect. When heparin binds to surfaces, such as
traditional vascular grafts, the surfaces become more blood compatible. Heparin can also inhibit both
smooth muscle cell (SMC) proliferation and intimal hyperplasia. In one example, Laemmel et al.
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studied the effect of immobilizing heparin on the surface of gelatin or albumin proteins using a
water-soluble carbodiimide. On heparinized gels, they noted an inversely proportional relationship
between the growth of SMCs and the dose of heparin. Although it decreased with time, this
inhibitory effect was still significant after 8 days [184].
Werkmeister et al. demonstrated another example of how to use heparin to improve a
standard ePTFE graft. They modified the grafts using either collagen or heparin treatments and
then implanted a 100±120 mm long prosthesis as a single aorto-left-external iliac bypass graft.
Both types of modifications elicited positive tissue responses as indicated by good healing and a
lack of fibrosis and tissue encapsulation. These findings suggested that modified prostheses
provided a uniform flat collagen-rich surface that promoted endothelial adhesion and proliferation
[185]. A study by Ramshaw et al. showed that the addition of a collagen or heparin coating to
OmniflowTM II grafts led to better collagen coverage and a decreased occurrence of mesh
breakthrough. OmniflowTM II consists of a mesh-reinforced ovine collagen tube designed for use
as a peripheral vascular replacement [186]. Walpoth et al. coated ePTFE and polyurethane grafts
with heparin and evaluated them in a rat model. This study confirmed that heparin significantly
reduced overall graft thrombosis, but did not indicate that the heparin coating had any effect
on intimal hyperplasia. Throughout the experiments, ePTFE grafts outperformed polyurethane
grafts [187].
Heparin-like materials are prepared by reacting amino acids with grafted chlorosulfonated
polystyrene. Polystyrene, however, does not have suitable mechanical properties for vascular graft
applications. Porte-Durrieu et al. investigated grafting styrene onto poly(vinylidenedifluoride)
(PVDF) films and poly(hexafluoropropylene vinylidenedifluoride) copolymer films via gamma
radiation or swift heavy ion techniques. Through chemical modification, the authors created
functionalized polymers. These fluoropolymers displayed suitable mechanical and chemical
properties for use in vascular graft applications. Heparin-like activity was observed on bound
polymers and was enhanced by derivatization with amino acids containing acidic side chains.
Furthermore, a thrombin time test showed an improvement in the antithrombotic character of the
modified PVDF. For polymers prepared by gamma radiation, a homogeneous distribution of the
styrene was observed regardless of surface modification or functionalization. For the swift heavy ion
radiation, the topographies obtained by SEM and AFM showed circular islands of different sizes
depending on the grafting yield value, a measurement based on the FTIR spectra of the grafted
films [188].
All of the studies in this section used surface coatings or modifications to diminish the
inflammatory response and improve both wound healing and patency of vascular grafts. Some of
these methods will play a significant role in the search for improved artificial vascular grafts.
Currently, the inherent inflammatory response of the two materials approved by the FDA dictates the
design and implementation of coating approaches. It seems that in order to improve the patency of
artificial grafts, especially small-diameter grafts, these coatings/modifications need to be combined
with different polymeric materials when designing new vascular biomaterials. Thus, new or modified
materials that incorporate biologically enhanced coatings may offer the best approach for creating
successful small-diameter vascular grafts.
5.10. Other polymers
5.10.1. Segmented polyurethanes
Segmented polyurethanes (SPUs) offer desirable physical properties as well as good
biocompatibility. Polyurethane-amides (PUAms) form strong intermolecular hydrogen bonds due
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to the presence of amine bonds in the macromolecular chains. The incorporation of chain extenders
such as fumaric or maleic acid provides a way to insert reactive double bonds into the polymer chain.
These bonds can then serve as grafting sites for further derivatization. Petrini et al. synthesized and
characterized PUAms carrying polyether or polycarbonate soft segments for use in vascular graft
applications. As expected, tensile tests showed good mechanical properties, and cytocompatibility
experiments revealed successful interactions between the base polymers and human skin fibroblasts.
The ability of these grafts to bind fibronectin and cell-adhesive peptides, however, remains under
investigation [189]. Another modification to SPUs involves carbon deposition. Kaibara et al.
demonstrated endothelial cell adhesion and proliferation on carbon-deposited SPU. The use of a
synthetic peptide (GRGDSP) further enhanced EC adhesion and proliferation [190].
Preliminary studies have shown that bovine aortic endothelial cells (BAECs) proliferate to
confluence on a segmented polyurethane (SPU) sheet treated with air plasma. Kawamoto et al.
extended this research by using air plasma to modify the inner surface of a SPU-coated glass tube
(i.d. 1.2, 2, or 3 mm). The BAECs proliferated well and reached the middle portion of the inner
surface of the treated tubes. The adhesive strength of the BAECs measured at least 5.5 Pa, well over
the flow shear stress of 2 Pa created by blood flow [191].
Yoneyama et al. used SPU as a coating and evaluated it in the rabbit model. They coated a
small-diameter Dacron graft (2 mm i.d.) with a blend of segmented poly(urethane) (SPU) and 2methacryloyloxyethyl phosphorylcholine (MPC) in the form of poly(MPC-co-2-ethylhexyl
methacrylate) (PMEH). MPC polymers effectively suppressed clot formation, platelet adhesion,
and platelet activation in blood contact studies. The authors found that XPS graphs of coated
polymers contained peaks of nitrogen, phosphorus, carbon, and oxygen, indicating the presence of
the MPC unit at the surface of the prosthesis. After a 5 day implantation, the prosthesis remained
clear and allowed continued blood flow. Several studies have shown that MPC can suppress clot
formation and platelet adhesion and activation. Further evaluation of the performance of SPU
coatings require long-term in vivo assessment that builds off the results of the above study and the
known favorable properties of MPC [192]. In another study, Yoneyama et al. prepared two kinds of
polyurethane vascular prostheses (2 mm i.d.). The two types of grafts contained 7.5 wt.% (SPU/
MPC(7.5)) and 10.0 wt.% (SPU/MPC(10)) MPC polymers, respectively. These grafts were
implanted into rabbit carotid arteries and harvested after 4 weeks. The SPU/MPC(10) prosthesis
proved less thrombogenic than the SPU/MPC(7.5), and after the 4-week implantation, the surface of
the SPU/MPC(10) appeared not only macroscopically clear, but also devoid of either fibrin or a
pseudointima [193].
5.10.2. Polyurethane
Due to their mechanical properties and good hemocompatibility, polyether-type polyurethanes
are widely used as cardiovascular biomaterials. The synthesis of polyurethane (PU) vascular grafts
involves a traditional solvent-casting salt leaching technique. To enhance compatibility, one group
coated the inner surface with a layer of gelatin and then crosslinked the gelatin with the epoxy
Denacol (EX-810 epoxy, Nagase Chemicals, Osaka, Japan). Results from tensile tests showed
Young's moduli very similar to those of carotid artery. Nonporous PU surfaces exhibited less platelet
activation, and the gelatin-coated graft gave the best short-term blood compatibility performance
among the grafts tested. In addition, gelatin-coated PU grafts promoted faster endothelial cell
adherence due to the improved cell spreading on these surfaces [194]. Another group working on
improving the blood compatibility of polyurethanes is Wetzels and Koole. They presented a novel
surface modification for polyurethanes based on the use of poly(N-vinylpyrrolidinone). They
hypothesized that the water-soluble polymer would mask the presence of the graft and improve
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hemocompatibility. Initial results showed that the anchoring of the poly(N-vinylpyrrolidinone)
proved to be an easy photochemical coating procedure and led to decreased surface thrombogenicity
[195].
Surface-induced calcification limits the durability of polymeric materials. It causes a decrease
in the flexibility of biomaterials, which results in mechanical failure. Chandy et al. investigated the
hypothesis that certain antibiotics (ampicillin, neomycin sulfate, streptomycin sulfate, and
gentamycin) modulate surface calcium binding on polyurethane by changing the calcium
mobilization and crystallization. It appeared that these antibiotics interfered with the transport of
calcium and phosphorus ions by blocking the calcium entry channels or masking the surface sites for
calcium nodulation. More studies are needed, however, to develop effective combination therapies
that allow the use of antibiotics while limiting calcium deposition [196].
Eberhart et al. reviewed three polyurethane vascular grafts (Corvita1 (Corvita Corp., Miami,
FL), Thoratec1 (Thoratec Lab. Corp., Berkeley, CA) and the Pulse-Tec1 (Newtec Vascular Products
Ltd., Clwyd, UK)) with novel designs with respect to their physical, chemical, and mechanical
properties. The Corvita graft consists of two layers: a fibrillar PU inner layer and a warp-knitted
polyester PET mesh outer layer, impregnated with a sealant for reinforcement. The Thoratec graft
contains a spirally wound polyester monofilament and a nonporous layer embedded within the PU
wall. The Pulse-Tec graft is a microporous polyetherurethane. The reinforcements provided by the
Corvita and Thoratec grafts have helped increase their resistance to kinking. Both of these grafts
offer a higher degree of integration of the reinforcement into the prosthetic wall. The nonporous
layer under the luminal surface of the Thoratec avoids any transmural tissue infiltration in the
development of endothelium; this graft depends on arterial transport and pannus migration for
endothelialization. The Pulse-Tec graft on the other hand, is highly porous and ongoing in vivo
studies will provide information on the success of endothelialization. The tensile strengths and
compliance of all three grafts were in the range of the native artery [197].
Clinical trials of new vascular graft alternatives test the safety and efficacy of the device.
Specifically, these trials evaluate the biocompatibility and performance of the chemistry and
structure used in prosthesis construction in order to identify potential device defects. For example,
weakening due to degradation could lead to catastrophic failure. Vascugraft1 (B. Braun Melsungen
AG, Melsungen, Germany), a microfibrous polyesterurethane vascular prosthesis showed good
healing characteristics, but was withdrawn from the market due to concerns with haemodynamic
conditions. In a study conducted by Marois et al., chemical modifications of the graft were observed
during the implantation schedule of Vascugraft1. Furthermore, breaks of microfibers near the
anastomotic sites during longer implantation periods also led to safety concerns [198]. This study
emphasized the need for new devices to outperform the current alternatives. Another study using
Vascugraft1, conducted by Zhang et al., also noted some deterioration in the fibrous structure after
about 1-year post-implantation. Moreover, the clinical performance of this prosthesis in below-knee
substitutions remained similar to that of ePTFE currently used. As noted above, the manufacturer has
withdrawn this product from the market in good faith [199].
5.10.3. Poly(glycolic acid) and polydioxanone
The PGA grafts implanted by Greisler into rabbit infrarenal aortas showed better healing
characteristics than Dacron grafts. Since an increase in the ratio of lactide:glycolide rings resulted
in slower resorption and slower kinetics of tissue ingrowth, further studies were performed using
polydioxanone grafts with this altered lactide/glycolide ratio. Mechanical studies of explanted
grafts revealed arterial-like elasticity characteristics and better resistance to fatigue or bursting
[143,150].
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5.11. Microporosity
Doi et al. used a laser technique to create micropores in a vascular graft. They fabricated a
segmented polyurethane (SPU) prosthesis (2 mm i.d.) by dip-coating the SPU solution on a glass
mandril and air-drying the mandril. Applied laser pulses then created micropores of a controlled size
and distribution. The outer and luminal surfaces of the graft were then coated with photoreactive
gelatin and irradiated with UV light. Findings from this study revealed that the elasticity of the
microporous tube increased with the number of pores. As a result, this approach provided a
mechanism to achieve better compliance matching between the graft and the native artery [200].
This research was extended into an in vivo study in the infrarenal aortas of rats. In this experiment,
the microporous grafts were coated with a mixed solution of photoreactive gelatin, heparin, and
cytokines (vascular endothelial growth factor and/or basic fibroblast growth factor). The cytokine
treated group experienced a much higher percentage of endothelial cell (EC) coverage than the
noncytokine containing control group. No significant difference in EC coverage was observed
among the different cytokine groups (VEGF, bFGF and VEGF/bFGF). In all groups, the
morphologic features of the ECs appeared the same, however, the VEGF group showed some
capillary orifices at the midportion of the graft. Conclusions from this study stated that the
impregnation of VEGF enhanced both transanastomotic tissue and transmural ingrowth and in
particular, increased capillary ingrowth [201].
Gravenwoger et al. perforated a biosynthetic vascular prosthesis (Omniflow, BioNova,
Melbourne, Australia) to stimulate transmural capillary ingrowth. These modified grafts were
implanted into the carotid arteries of sheep for 3 months. At explantation, the authors noticed a white
and almost entirely endothelialized neointimal layer extending along the entire length of the
perforated grafts. Moreover, transmural capillary growth was also seen [202]. Okoshi et al. examined
the effect of penetrating micropores using polyurethane-polydimethylsiloxane vascular grafts
implanted in the infrarenal aorta of rats. They found that endothelialization was limited to
anastomoses in grafts with a hydraulic permeability of less than 37 ml/min cm2 per 120 mmHg. For
grafts with hydraulic permeabilities greater than 37, endothelial cells covered most of the graft area,
and the voids of the graft wall contained newly formed capillaries. The study also demonstrated that
graft patency increased with higher porosities [203].
5.12. New materials
Huynh et al. discussed the use of the intestinal collagen layer (ICL) as a biomaterial for the
integration and remodeling of a small-diameter vascular graft. After mechanical and chemical
cleaning, the ICL yielded a sheet of predominantly type I collagen, which retained its inherent
organization and strength. These sheets were used to construct 4 mm diameter tubes impregnated
with a layer of dense fibrillar bovine collagen. These tubes were then crosslinked with a
carbodiimide and coated with a heparin complex before implantation into the carotid artery of
rabbits. After 3 months, the grafts showed extensive remodeling which made it difficult to identify
the original collagen. Although the lumen of the grafts had endothelialized within 3 months,
additional studies need to demonstrate mechanical strength and long-term resistance to both
hyperplasia and aneurysm formation. Overall, this study offered promising results for the future of
vascular graft tissue engineering [204].
Grafts prepared from photo-oxidized bovine internal carotid or internal mammary arteries were
interposed in the femoral artery of dogs by Moazami et al. After 6 weeks, 82% of the photo-oxidized
grafts remained patent. Preliminary data indicated that the grafts did not promote acute thrombosis
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and presented excellent handling characteristics. As with many studies, more extensive experiments
are needed before developing any concrete conclusions [205].
A new polyester gelatin-sealed graft fabricated using a different and more compact knitted
structure has been evaluated in vivo and in vitro. Mary et al. performed a thorough evaluation of this
graft by investigating textile properties, physical properties, and chemical characteristics. They
found that this graft had a higher porosity (75%) than four commercially available polyester knitted
prostheses. The gelatin sealant made the new graft less permeable to water, and both the dimensional
stability and dilation resistance of the sharkskin knitted graft were superior to the control grafts.
Additionally, during canine thoraco-abdominal bypass procedures, the flexible new graft allowed for
easy implantation. With restored blood flow, no bleeding occurred through the graft wall, and grafts
remained patent at 3, 6, and 12 months post-implantation. The luminal surfaces of the prostheses
excised at 3 months were covered with a thin layer of collagen and endothelial-like cells; isolated
areas of endothelium and thrombotic deposits were observed at 12 months [206].
As determined by graft integration and endothelialization, the short-term success of smalldiameter grafts made from biopolymers or more biocompatible traditional polymers appears to be
greater than that of synthetic polymers currently approved by the FDA. One explanation may be that
biopolymers can more easily mimic the natural extracellular matrix in terms of bioactivity,
mechanical properties, and structure. Another explanation may involve the lack of bioactive factors
in synthetic scaffolds. The more biocompatible polymers tend to elicit less of an immune response.
Furthermore, modifications, such as coatings, to these grafts could provide a way to greatly reduce
inflammatory response at the implantation site. As we design novel scaffolds, we must consider all
of the properties affecting graft success, including mechanical, chemical and biological attributes.
Although natural biopolymers have not yet proven entirely successful, the current response of
synthetic polymers for small-diameter vascular graft applications also remains inadequate. As we
design new materials, we have more control with synthetic, biodegradable scaffolds, especially for
mechanical integrity. For this reason, we should strive to use this control to achieve new materials
that integrate with the surrounding tissue.
5.13. Degradable scaffolds
An ongoing quest of several investigators focuses on engineering new tissues by transplanting
cells onto biodegradable polymer scaffolds. After implantation, cell growth would cover the scaffold,
and, after an appropriate amount of time, the scaffold would completely degrade away leaving
behind a new, all-natural tissue construct. Several challenges exist for the successful engineering of
such a scaffold. For vascular graft applications, one of these concerns relates to correctly matching
mechanical properties of scaffolds and natural tissues since vascular grafts must resist large
compressive forces.
Mooney et al. investigated the strength of tubes formed from PGA fiber meshes. They stabilized
these meshes by spraying atomized solutions of poly(L-lactic acid) (PLLA) and a 50:50 copolymer of
poly(D,L-lactic-co-glycolic acid) (PLGA) dissolved in chloroform. Regulation of both the
concentration of polymer in the atomized solution and the total mass of polymer sprayed on the
mesh allowed for control of the pattern and extent of bonding. This method bound the outermost
fibers of the device and conserved porosity throughout the interior sections. As the extent of bonding
increased, the compressive resistance of the devices also increased. A comparison of the different
materials showed that the PLLA bonded tubes exhibited greater resistance than those made with
PLGA. During in vitro experimentation, endothelial cells and smooth muscle cells seeded onto these
devices proliferated and grew into the meshes [207].
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A new poly(glactin)-poly(glycolic acid) and polyhydroxyalkanoate (PHA) composite will
slowly degrade and can withstand systemic pressures. Shum-Tim et al. engineered a tubular conduit
using an inner layer of poly(glactin)-poly(glycolic acid) and a biocompatible outer layer of PHA.
They then seeded a mixed cell population of endothelial cells, smooth muscle cells, and fibroblasts
onto the PHA surface. Following cell seeding, the grafts were implanted in the infrarenal aorta of
lambs. The composite grafts showed superior tensile strength, flexibility, and ease of handling and
suturing relative to acellular tubular conduits. After 5 months, the grafts remained patent, and the
authors detected the presence of matrix metalloproteinases. The presence of metalloproteinases
suggested that the graft could support favorable biological events. Other observations were that
most of the cells from in vitro seeding remained in the tissue-engineered structures and that the
elastic and collagen fibers seemed to organize uniformly according to the direction of blood flow
[208].
5.14. Drug delivery
Heparin is widely used as an anticoagulant and antiplatelet agent. Since systemic administration
of heparin elevates the risk for hemorrhage, site-specific delivery of heparin to injured vessels
provides a favorable alternative to systemic delivery. Teomim et al. investigated one such sitespecific delivery system using flexible sheets of laminated poly(dimer erucic acid-co-sebacic acid)
and poly(lactic acid). By mixing heparin powder in the polymer melt, they were able to test the effect
of heparin loading within rat endothelial injuries. In vitro release studies showed full recovery of
loaded heparin, indicating that heparin did not react during polymerization. Approximately 50% of
the heparin was released in the first 24 h, and the concentration data from in vivo studies matched the
results of in vitro experiments. In the treated injuries, a very thin layer (50% less than untreated) of
neointima was observed. These results show the promise of using heparin within polymeric devices
for local drug delivery [209].
Another local delivery system of heparin involves heparin encapsulation. For example, an
aqueous heparin solution can be spray-dried into biodegradable poly(D,L-lactic-co-glycolic acid)
(PLGA). Yang et al. detailed this system and investigated the effect of the heparin delivery on
smooth muscle cell (SMC) proliferation. An initial bolus release accounted for about 25% of the
total heparin content, and cumulative release times for PLGA 50:50 and PLGA 75:25 were 15 and 30
days, respectively. The heparin release had a notable inhibitory effect on the proliferation of SMCs.
An interesting observation from this study was that heparin inhibited SMC proliferation only from
cells isolated from male patients (not with those from female patients). One reason for this different
interaction suggested a dissimilarity in the phenotypes of the cells; however, the true explanation
remains unclear [210].
Another coating of interest is the commercial platelet inhibitor, Triflusal, 2-acetoxy-4trifluoromethyl benzoic acid. This drug has a chemical structure closely related to aspirin, which is
known for its anti-aggregating effects. Rodriguez et al. found that Triflusal, used as a coating for
PTFE or Gore-Tex grafts, improved anti-aggregation of platelets in static studies with ovine blood.
They also characterized this coating as both biocompatible and bioresorbable. Regarding
degradation, the hydrophilic character of the polymer surface controlled the rate of release of
Triflusal [211]. Another group, Roman et al., examined the benefits of salicylic acid. They coated
Dacron grafts with a copolymer consisting of 2-methacryloyloxybenzoic acid and 2-hydroxyethyl
methacrylate. The resulting coating contained weak and reversible covalent bonds, which could be
hydrolyzed in physiological medium. Furthermore, the system could then release salicylic acid
through hydrolysis of the acrylic ester group. Findings from in vitro studies revealed a constant
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release rate of salicylic acid residues over 10±12 days. Additionally, the coated grafts showed a
noticeable decrease in the number of platelets adhered [212].
Pulfer et al. presented a novel method of using the properties of nitric oxide (NO) to inhibit
platelet aggregation and smooth muscle cell proliferation. A new class of NO releasing compounds,
diazeniumdiolates, were shown to express bioactivity and spontaneously generate NO in aqueous
media at a rate dependent on pH and temperature. NO releasing microspheres (10±50 mm) were
created from crosslinked polyethyleneimine/NO adduct and incorporated into the pores of a Gore-tex
vascular graft. In vitro studies demonstrated a controlled release of NO, with a calculated half-life of
66.2 h [213].
We have learned a great deal regarding the hemocompatibility of materials. Moreover, we have
also succeeding in combining bioactive signals with synthetic polymers via immobilization and
encapsulation for controlled release. Novel polymers, which integrate this knowledge and these
methods and which attempt to mimic the structure and mechanics of native blood vessels, may
greatly improve vascular graft performance.
6. Nerve regeneration
The nervous system contains two major divisions: the central nervous system (CNS) and the
peripheral nervous system (PNS). Constituting the brain and spinal cord, the CNS consists of a vast
number of neurons, astroglia, microgliga, and oligodendrocytes that act to coordinate, recognize,
initiate, propagate, and process signals from external or internal stimuli. The PNS contains all the
elements of the nervous system not fully contained within the brain or spinal cord. Some of these
elements include ganglia, motor neurons, and sensory nerves. Fig. 8 shows the structure of a PNS
neuron. Although many different types of neurons exist, the general morphology of a multipolar
neuron usually includes a cell body or soma, dendrites, and an axon. By means of ionic currents,
neurons act to conduct electrical impulses (action potentials) from the dendrites, which serve as
signal receivers, through the soma and onto the axon terminal, the signal transmitter. Once the signal
reaches the end of an axon, several vesicles containing chemicals called neurotransmitters release
their contents into the synaptic cleft, the region between the axon terminus and its target (another
neuron or muscle fiber). In the case of a motor neuron, for example, this electrical signal causes
muscle activation through the release of acetylcholine. The soma and dendrite system of many
neurons are located within the CNS; however, these neurons can extend axonal processes deep
within the periphery (some axons can measure more than 1 m in length). Surrounding the axon of a
PNS neuron are Schwann cells. These glial cells give structural support to the axon and also aid in
accelerating the propagation velocity of action potentials by insulating the neuron with a myelin
sheath [4,113]. Though the seemingly simple primary focus of nerve regeneration aims to reestablish
electrical and chemical communication across a severed pathway, the complex interaction between
neurons and glia, especially during injury, creates a formidable challenge.
Materials design and testing for peripheral and central nervous system regeneration is an active
area of research. Impaired nervous system function occurs through accident and disease. Stroke leads
to brain damage and loss of function, spinal cord injury can lead to partial or complete paralysis
below the point of injury, while peripheral injury can lead to loss of function of small areas of the
body. In addition, diseases, such as Parkinson's, exist that lead to nervous system impairment. To
date, greater success has been achieved in repair of the peripheral nervous system. One thought as to
why relates to the density of the extracellular matrix of the tissue, with the central nervous system
containing a more dense neural tissue than the peripheral nervous system. Higher extracellular
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Fig. 8. A pen drawing of a myelinated neuron adapted from [264]. The general morphology of a multipolar neuron usually
includes a cell body or soma, dendrites, and an axon. Surrounding the axon of a PNS neuron are Schwann cells. These glial
cells give structural support to the axon and also aid in accelerating the propagation velocity of action potentials by
insulating the neuron with a myelin sheath.
matrix density may make it more difficult for growing axons to extend through the tissue. As with all
tissues and organs, several other factors, including the release of bioactive signals, must also play a
significant role in the regenerative process. This review focuses mainly on efforts to improve the
local environment for peripheral nerve regeneration. A detailed review of central nervous system
regeneration is provided by Stichel and Muler [214].
Excellent reviews exist for polymer based controlled release systems for biologically active
factors [215] as well as for polymers designed for controlled release of neurotrophic factors that
cannot naturally cross the blood brain barrier [216]. As further discussion of these materials would
be repetitive, those interested in controlled release, both in nervous tissue and other areas of the
body, are referred to the cited references.
6.1. Conductive polymers
Researchers have postulated that conductive polymers may interact with nerve electrical
conductance in some manner to improve nerve regeneration in both the central and peripheral
nervous systems. Polypyrrole remains a potential candidate as it has been shown to be
biocompatible. Collier et al. synthesized polypyrrole-hyaluronic acid (PP-HyA) composite films
to study their effect on neurite extension in PC12 cells and their ability to induce angiogenesis.
Neurite extension and angiogenesis have been aptly coupled to promote proper perfusion of the
regenerating tissue. The films were more brittle and had a lower conductivity than polypyrrolepolystyrenesulfonate (PP-PSS) films previously studied by this group. This decrease in properties
likely occurred due to HyA nodule formation during polymerization of the films as a result of the
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
increased solution viscosity. To reduce brittleness, studies were carried out using thin, 0.05 mm, PPHyA films cast on PP-PSS films for a total thickness of 0.15 or 2.0 mm. While much of the current
passed through the PP-PSS layer in this configuration, some current still passed through the PP-HyA
layer. In vitro, PC12 cells, in the presence of nerve growth factor, attached and spread neurites on the
0.15 mm PP-HyA films. In vivo, subcutaneous studies indicated little inflammation as a result of
implantation of the 2.0 mm films. After 2 weeks, a two-fold increase in angiogenesis was seen with
PP-HyA coated PP-PSS films than with PP-PSS films alone. At 6 weeks, however, the two types of
film demonstrated an equivalent angiogenic response [217]. This increased rate of angiogenesis in
the PP-HyA films may be crucial for successful neuronal growth. As a result, one important design
parameter involves simultaneous neurite and capillary growth. Lag time in capillary growth could
result in arrested axon extension. Jakubiec et al. studied the biocompatibility of polypyrrole-coated
polyethylene terephthalate woven fabrics and assessed the optimal conductivity of these materials
with respect to biocompatibility. They found that as the level of conductivity increased, endothelial
cell migration and viability decreased as determined using a direct contact assay with chick
embryonic aortic endothelial cells. Increased conductivity correlated directly with lower levels of
foreign body response. These results point to an optimal, intermediate level of conductivity of
polypyrrole-coated polyester fabrics in the range of 103±104 O/square [218]. This study needs to be
extended to other polypyrrole containing materials to see if any generalized trends regarding nerve
regeneration materials emerge. In addition, the results need to be coupled with neurite extension
studies to optimize extension, endothelial cell growth, or angiogenesis and foreign body response.
6.2. Nerve regeneration matrices in the absence of tubulation
As will be discussed below, implantation of nerve guide tubes are currently the optimal method
to achieve nerve regeneration of transected nerves in the absence of transplanted nerve. It is believed
that the nerve guide tube provides a barrier between the growing axons and the surrounding
environment. This barrier may help limit unwanted scar tissue formation. At the same time, the guide
tube may need to be removed at a later time to prevent tissue build-up within the tube that can cause
pinching of the nerve. Degradable tubes may eliminate the need for a second surgery, but
degradation can add complexity to the design of the regenerative system. In an attempt to design
hydrogel-like materials for use in the absence of tubular nerve growth guides, Suzuki et al. studied a
1% sodium alginate solution crosslinked with ethylenediamine and carbodiimide followed by freeze
drying. The alginate gels were then used to repair 7 mm gaps in the sciatic nerve of male wister rats.
As determined by functional studies of the limbs and histological evaluation 18 weeks after
implantation, alginate gels supported nerve regeneration, and the alginate fully resorbed after 18
weeks. Regenerated nerves tended to have smaller diameters than healthy nerves from the same
animal. As controls, nerves were severed and left untreated. Control animals demonstrated no
functional recovery of macroscopic regeneration in the absence of alginate [219]. Regeneration in
the absence of tubulation demonstrates the possibility of obtaining fully regenerated nerves without
tubulation and partial isolation from the body. As seen below, a lot of effort has focused on
degradable nerve guide tubes. The ability of a material to function in the absence of a degradable
tube may remove one layer of the complexity of designing materials to interact in a favorable way
with the body. However, the possible significance of a barrier between the growing axons and the rest
of the body remains an important consideration. Further studies with freeze-dried alginate wrapped
with a poly(glycolic acid) mesh in a 50 mm cat sciatic nerve gap showed good functional recovery
after 13 weeks and complete degradation of the implanted material. In addition, the nerve fiber
distribution was different from that of normal nerve and the perineurium was less developed [220].
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Although the authors did not discuss the barrier activity of the poly(glycolic acid) mesh, these results
are exciting due to the length of nerve that regenerated. It would be interesting to study the alginate
hydrogels in the presence of a true guide tube to assess whether synergy exists between the two
systems.
6.3. Directional nerve guidance
6.3.1. Tube material and structure
Tubulation can provide directional guidance as well as support to nerve regeneration by acting
as a barrier between the regenerating environment and the rest of the body. Heath and Rutkowski
give an excellent review and discussion of the necessary parameters for tubulation. They discuss the
usefulness and limitations of silicone tubes such as improved nerve regeneration and the need for
secondary surgery to remove the tube in order to prevent constriction of the nerve from foreign body
response. They also discuss the required parameters for degradable guide tubes including flexibility,
stability for the time it takes for a nerve to regenerate, but degradation soon thereafter, mechanical
stability, and porosity to allow diffusion of nutrients and wastes, but small enough pores to limit
cellular invasion. Finally, they discuss the inclusion of Schwann cells for nerve guidance and
neurotrophic factors for nerve extension [221]. Additional information is provided below.
Several material properties likely influence how tubulation affects axonal extension. Buti et al.
studied the effects of the tube length, internal diameter, and wall thickness on peripheral nerve
regeneration using silicone guide tubes. They performed the studies in the sciatic nerve of mice and
determined the degree of regeneration from the following functional parameters: reinnervation of
muscle, skin, and sweat glands. They found that an increased gap length delayed the initiation of
reinnervation and reduced the percentage of recovery. In addition, gap lengths over 2 mm did not
heal in the absence of tubulation. Tubulation allowed repair of 2 and 4 mm gaps, but was less
effective in 6 and 8 mm gaps. In all cases, regeneration resulted in fewer axons than in healthy tissue.
These results suggested that tubes with an inner diameter measuring 2.5 times that of the sciatic
nerve and a wall thickness as thin as possible without collapse provided optimal conditions for nerve
regeneration [222]. Both of these parameters should be optimized for each new material studied
since tubes that showed greater promise than silicone tubes may experience improved success with
the proper diameter and wall thickness. It is likely, however, that a similar diameter will be optimal
for all materials.
Ellis and Yannas reviewed the use of collagen-blend-glycosaminoglycan polymers for use in
tissue engineering. They also discussed some of the physical parameters of these materials that are
useful in nerve regeneration. Uniaxial pore orientation coupled with pore sizes ranging from 5 to
300 mm were optimal for nerve regeneration with functional recovery, which improved with
increasing pore size [223]. Chamberlain et al. used silicone, collagen and porous type I collagen
nerve guide tubes to study nerve regeneration of the sciatic nerve in adult rats. One benefit for using
degradable type I collagen tubes is that they eliminate the need for surgical removal of the guide
following nerve regeneration They either filled tubes with a type I collagen/chondroitin-6-sulfate
copolymer averaging 95% volume fraction with a pore diameter of 5 mm or left tubes empty. They
saw significant improvement of neurite extension in terms of the number of neurites and thickness of
neurites in filled tubes within each group, e.g. porous collagen filled and empty tubes, but were
unable to draw conclusions between different types of filled tubes. The collagen/chondroitin-6sulfate filled matrix significantly enhanced nerve regeneration within each type of tube studied. As
determined by myoblast presence along the walls of silicone tubes and its absence along the walls of
the porous collagen, the porous collagen tubes appeared to provide a better nerve guide substrate
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than silicone [224]. In all cases, axon numbers were well below that of normal nerve after a 6-week
period. In a follow up study, the authors further defined the collagen/GAG matrix to be 98% collagen
and 2% chondroitin-6-sulfate. They again followed healing in 10 mm rat sciatic nerve gaps with and
without tubulation. Collagen/GAG matrix had significant effect on healing with silicone tubes with
all of the filled tubes showing axon regeneration while only two of nine unfilled silicone tubes
showed regeneration. All of the silicone tubes had continuous capsules comprised of a layer of
myoblasts 10±15 cells thick. In contrast, the observation that collagen/GAG matrix had no effect in
the healing process of collagen tubes contradicted the previous study. Both filled and empty collagen
tubes showed axonal growth and 1±2 cell thick noncontinuous lined capsules at 30 and 60 weeks,
which was promising since normal nerves have a 1±2 cell continuous layer. Nontubulated 10 mm
nerve gaps experienced significant wound healing and contraction with thick scar formation over the
distal and proximal stump. The authors postulated that the myoblast ingrowth played a role in
limiting nerve regeneration and the lack of myoblast growth in collagen tubes improved nerve
regeneration in these tubes [225]. These studies clearly illustrate the importance of porosity as a
parameter for guide tube design. Some porosity allows for the exchange of smaller macromolecules
while limiting cellular invasion. The diffusion of biological macromolecules may provide the
environment that limits myoblast infiltration of the system.
As with other tissue systems, a lot of investigation has focused on using degradable polyesters
because of their know biocompatibility. Porous PLLA tubes were made using a salt extraction
procedure with a salt weight fraction of 90% coupled with crystal sizes of 150±300 mm. The tubes
had inside diameters of 1.6 mm, outside diameters of 3.2 mm, and lengths of 12 mm, and the
resulting tensile strength of the tubes measured 81:7 35:1 MPa with a modulus of 1:0 0:4 MPa.
These tubes were implanted into 12 mm gaps in the Sprague±Dawley rat sciatic nerve model. As
controls, isografts from donor animals were used to repair injured sites. Functional and histological
evaluation performed at 6 and 16 weeks post-implantation showed improved functional recovery for
both graft types through 16 weeks. Compared to control grafts, both the number and density of axons
were significantly less for the tubulated implants with the exception of nerve fiber density at 16
weeks. Degradation of the polymer did not seem to inhibit nerve growth, however, no comparison
was made with normal nerves [226].
Gautier et al. constructed tubes of poly(D,L-lactic-co-glycolic acid) 50:50, (PLA25GA50),
poly(D,L-lactic acid) and high MW blend low MW tubes of poly(L-lactic acid) PLA. They studied the
degradation rates of the materials in vitro and the toxicity of the degradation products to Schwann
cells in vitro. They found that tubes of (PLA25GA50) began to degrade at day 7 and poly(D,L-lactic
acid) began to degrade at day 28. Increasing amounts of oligoPLA increased the degradation rate of
the PLA tubes with 30% low molecular weight incorporation causing degradation similar to that of
poly(D,L-lactic acid). In vitro, the break down products showed no adverse effects on Schwann cell
cultures. In vivo, the degradation had no effect on the inflammatory response or degeneration of the
nerves [227]. These results suggested that the poly(hydroxyacid) tubes can be tailored to have specific
degradation rates and mechanical properties without negatively impacting the healing rate of nerves.
Rodriguez et al. have compared the degree of nerve regeneration in a 6 mm gap injury of the rat
sciatic nerve using tubes of varying porosity. Poly(L-lactide-co-e-caprolactone) tubes with 1 mm
inside diameters and wall thickness of 150 mm were constructed with high, low, and no porosity.
High porosity was achieved by incorporating 10 mm glucose particles in equal weight to the polymer
during molding, while low porosity was achieved by including an equal weight of amylose particles
of less than 10 mm in diameter. Controls consisted of nondegradable porous and nonporous
polysulfone tubes. Functional as well as histological evaluation determined the degree of
regeneration for up to 120 days post-implantation. The high porosity polycaprolactone repeatedly
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showed faster and increased regeneration over other tubes and consistently outperformed
polysulfone tubes. The results from these experiments showed that both tube porosity and
degradation are significant factors in nerve regeneration [228].
Degradable polyphosphazenes have been studied as nerve guide tubes. poly{(ethylanato)1.4(imidazolyl)0.6(phosphazene)} (PEIP) was synthesized, and tubes were made by deposition of the
polymer solutions in methylene chloride onto a glass tube followed by solvent evaporation. Tubes
10 mm long, 1.3 mm i.d., and 1.8 mm o.d. were formed and implanted into a 5 mm rat sciatic nerve
gap with silicone tubes implanted as a control. Compared to the silicone implants, significantly less
scar tissue formed in the walls of the PEIP. Histological evaluation 45 days post-implantation
revealed a highly vascular epineurium layer surrounding regenerated nerve bundles with a thicker
epineurium layer seen in the PEIP implants. Furthermore, the degradable guide tubes showed
improved regeneration and decreased foreign body response [229].
The degradable polymers allowed significantly improved nerve regeneration over silicone tubes.
It would be beneficial to compare the different degradable polymers in one study to try to evaluate
which polymer would provide the best regenerative results. Once a polymer or polymer family is
identified, additional studies should attempt to understand why the use of a particular polymer results
in enhanced regeneration.
6.3.2. Fill material and structure
One thought as to how to improve guidance and extension of growing axons is to align the
extracellular matrix fibers that fill the nerve guide tube. This organization would provide a template
along which the growing axons could extend. To this end, Dubey et al. have developed a method to
align collagen fibers within a 4 mm i.d. ePTFE tube 10 mm in length. Pregelled and gelling
collagen-filled tubes (2 mg/ml) were exposed to high strength magnetic fields (4.7 and 9.4 T) for 2 h.
Stronger applied magnetic fields induce a higher degree of collagen fiber orientation along the tube
axis as determined by birefringence. Once aligned, the gels were then removed from the tube and
cultured with chick and rat dorsal root ganglia (DRGs) at 378C for 4 days prior to analysis. Using
vector analysis, it was found that neurites in an environment of increased fiber orientation extended
longitudinally along the axis of the tube to a greater extent than controls. The length of axons also
increased with fiber orientation. Immunohistochemical staining revealed that Schwann cell extension
occurred prior to neurite extension thereby providing a path for the axons [230]. The existence of
Schwann cells appears to be important to the regenerating axon. The Schwann cells likely secrete a
variety of molecules, including extracellular matrix and growth factors, which promote axon
extension. Work will be presented below which tries to manipulate the regenerating potential of
Schwann cells. In vivo analysis of aligned collagen filled type I collagen tubes also was performed.
In these studies, the mouse sciatic nerve was injured, and collagen filled guide tubes were implanted
into 4±6 mm gap injuries. At 30 days post-implantation, the collagen tubes had partially resorbed; by
60 days they had fully resorbed. In agreement with other studies using unfilled tubes in 6 mm gaps in
the mouse model, only one of six controls (implanted with unaligned, filled collagen tubes) showed
regeneration of the sciatic nerve. In contrast, all four of the mice implanted with aligned, collagen
filled collagen tubes exhibited regeneration at 60 days, but contained fewer axons than that seen in
healthy sciatic nerve [231]. The possibility remains that a formulation that slows fiber resorption
would improve the number of regeneration axons. Moreover, it is also possible that other fibers may
enhance nerve regeneration. As a result, it would be interesting to compare aligned fibers of fibrin,
collagen, and laminin to observe their effect on regeneration. Some as of yet unknown combination
of extracellular fibers may also have a profound effect on the rate, number, and distance of neurite
extension.
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It may not be necessary to align fill gels as they are gelling within the tubes since premade fibers
also seem to enhance regeneration. Silicone tubes with an internal diameter of 1.5 mm, an external
diameter of 2.5 mm, and a length of 15 mm were used to study the effect of collagen fibers and
laminin or YIGSR (a short neurotrophic peptide derived from laminin) coated fibers on peripheral
nerve regeneration. Type I collagen fibers were synthesized with 100±150 mm diameters, and eight
of these fibers were inserted into the silicone tubes. Other fibers were coated with either laminin or
YIGSR, and eight of these were inserted into silicone tubes. Following preparation, all of the tubes
were then implanted into the male, rat sciatic nerve model with empty silicone tubes serving as the
control. After 8 weeks of implantation, the authors observed nerve bridging in 7 of 12 collagen fiber
filled tubes, 7 of 9 YIGSR-coated collagen filled tubes, and 6 of 9 laminin-coated collagen filled
tubes. Collagen gel filled and empty tubes showed no nerve bridging at 8 weeks. Furthermore, the
density of axons in the tubes containing YIGSR or laminin-coated collagen fibers was significantly
higher than in tubes with collagen fibers alone [232]. These data support the idea that axially aligned
nerve guide tube matrices improve nerve regeneration and that the neurotrophic factors laminin and
YIGSR also improve nerve regeneration. Due to the difficulty of obtaining purified laminin, the short
laminin peptide YIGSR remains the preferred neurotrophic factor. These studies combined the
positive effects of fiber alignment with those of bioactive signals to improve neurite growth. Further
studies combining bioactive signals and additional ECM fibers may result in better nerve growth/
guidance materials.
Further research has combined several materials in an attempt to restore nerve function across
large gaps. One example includes the work by Matsumoto et al. Recognizing the need for bridging
gaps longer than what has successfully been attempted (50 mm), they designed a composite conduit
and examined its effect in 80 mm gaps. The materials consisted of a cylindrical PGA tube coated
with collagen and filled with laminin-coated collagen fibers (70±80% type I and 20±30% type III)
measuring 50 mm in diameter and 90 mm in length. By using a PGA mesh, the authors hoped to
reinforce the tube and prevent scar tissue infiltration. The actual PGA tube had a 4 mm i.d., 90 mm
length, and 50 mm wall thickness. To evaluate this material, the authors implanted the tubes within an
80 mm gap created by removing a section of the peroneal nerve in adult beagle dogs. Following
surgery, the electrical activity of several muscle groups was recorded every month for 1 year. After 3
months, potentials could be recorded from muscle groups within 10 of the 12 dogs receiving grafts,
however, the levels of these potentials were lower than controls even after 12 months. These results
indicated that although they did not achieve complete functionality, some nerves did manage to span
across the 80 mm gap to reinnervate distal targets. The restoration of the treated dogs' ability to walk
almost normally without load after 10±12 months also showed that the novel polymer device aided
tissue repair. Histological evaluation of sacrificed animals revealed that the regenerated axons in
dogs receiving polymer implants had smaller diameters and thinner myelin sheaths than normal,
untreated axon controls. Only scar tissue filled the defect cavity in animals not receiving PGA tubes.
This study demonstrated the feasibility of attempting to restore nerve injuries spanning many
millimeters [233]. Although complete restoration of functionality was not obtained, the collagencoated PGA tubes filled with laminin-coated collagen fibers showed promise as a scaffold for
guiding nerve regeneration. The addition of other neurotrophic factors or biological molecules to this
material may further improve the quality of repair tissue.
Several unaligned matrices have been studied within tubes for nerve regeneration. Silicone
tubes, with an internal diameter of 1.5 mm and a length of 12 mm, were filled with three different
matrices: collagen gels, Biomatrix gels (a laminin rich extracellular matrix), and 2% methylcellulose
gels. Each of these materials was then evaluated with and without platelet-derived growth factor-BB
(PDGF-BB) and insulin-like growth factor I (IGF-I) by filling the tubes with 25 ml of solution
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resulting in 0.375 mg of PDGF-BB and 0.75 mg of IGF-I. Four weeks after implantation into
Sprague±Dawley rat sciatic nerves with gap injuries of 8 mm, collagen gel and methylcellulose gel
filled tubes containing growth factor had the greatest number of regenerating axons in the center of
the tubes. In fact, the number of axons reached on the order of 3500. Interestingly, at 8 weeks postimplantation, the positive effects of the growth factor had been masked, and both collagen gel and
methylcellulose gel filled tubes with and without growth factor experienced similar numbers of
regenerating axons. Methylcellulose filled tubes resulted in more highly functional nerves at 8 weeks
as determined by nerve conduction velocities. Again, these results were independent of the presence
of growth factors [234]. The degree of regeneration with the various gels may have resulted from
differences in bioactivity; however, this study did not confirm this hypothesis. The differences may
also be attributed to the density of the gels as gel density can act as an important parameter in nerve
regeneration. Unfortunately, the authors did not present polymer concentrations for the Biomatrix or
the collagen gels.
Fibrin, a biological hydrogel, is a natural wound-healing matrix that results from the process of
blood clotting. As such, many groups have used this material as a matrix for tissue engineered
products. Herbert et al. studied the effects of fibrinogen density and calcium ion concentration on the
structure of fibrin gels and neurite extension. Solutions of 5, 10 and 15 mg/ml fibrinogen and 2, 6,
10 mM Ca2‡ were studied. The authors found that as Ca2‡ concentration decreased at a constant
fibrinogen concentration, the number of fiber bundles increased while the average bundle diameter
decreased. As the fibrin concentration increased from 5 to 15 mg/ml, the number of fibrin bundles
increased, and the bundle diameter decreased with Ca2‡ concentration. Decrease in bundle diameter
should result in more branching and smaller pore sizes, which, in turn, should affect axonal
regeneration. Indeed, optimal regeneration was observed with 5 mg/ml fibrinogen and 10 mM Ca2‡
with neurite length decreasing with decreased bundle diameter [235]. Sakiyama et al. evaluated
3.5 mg/ml fibrin gels for neurite extension. They crosslinked into the fibrin gels, via factor XIIIa
transglutaminase sites, peptide sequences containing heparin binding domains with varying affinity
to heparin. Approximately 8.7 mol peptide/mol of fibrin were incorporated into the gels. Chick
dorsal root ganglia (DRGs) were then cultured within the gels for 48 h. Evaluation of the length of
the neurites was carried out and compared to control gels of fibrin without crosslinked peptide. The
heparin binding domain from antithrombin III, FAKLAARLYRKA, exhibited the highest affinity for
heparin. It also showed the greatest potential to induce neurite outgrowth in the fibrin gel/chick DRG
system averaging 73% greater extension than fibrin alone [236]. Schense and Hubbell evaluated the
peptide sequences RGD, RDG and DGEA when crosslinked into a fibrin matrix. They found that
8 mol peptide/mol of fibrinogen could be crosslinked into the gels via factor XIIIa crosslinking.
After 48 culture with chick DRGs they found that incorporation of RGD limited neurite extension to
0.9 of that of fibrin alone, DGEA increased extension to 1.2 that of fibrin and RDG, the control, had
no effect as would be expected [237]. Some combination of adhesive peptides may induce extension
further.
Another natural material investigated as a matrix for coordinating nerve regeneration is agarose.
Agarose, a thermally reversible polysaccharide hydrogel, contains repeating units of a-L-galactose
and b-D-galactose. The melting temperature can be modulated by altering some of the functional
groups located on the sugar residues. Bellamkonda et al. used hydroxyethylated agarose as a threedimensional scaffold for studying the growth of chick DRGs and PC12 cells. Using 1,1carbonyldiimidazole, they successfully derivatized the agarose and were able to crosslink several
peptide sequences to the agarose matrix. The active portions of these peptides included sequences
from laminin such as RGD, YIGSR, and IKVAV; GGGGG was used as a control sequence. The
average pore size for the hydrogels ranged from 310 to 360 nm depending on whether or not the
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agarose was derivatized. When either chick DRGs or PC12 cells were suspended within the gels, the
cells remained viable for up to 6 days. The level of neurite extension, however, depended on the type
of peptide used. For DRGs, the RGD and IKVAV sequences inhibited neurite extension relative to
the naked agarose and GGGGG peptide controls. The YIGSR peptide as well as a peptide cocktail
consisting of the RGD, IKVAV, and YIGSR sequences significantly enhanced neurite outgrowth over
time. PC12 cells in GGGGG-agarose gels did not extend neurites. In contrast, all of the other
peptides induced significantly greater amount of neurite growth relative to plain agarose. With this
cell line, however, the IKVAV peptide had the greatest affect on the neurons followed by RGD and
then YIGSR. No neurite outgrowth was observed in any of the gels when the PC12 cells were not
supplemented with nerve growth factor [238]. The results of this study suggested the importance of
coupling bioactive polymers to three-dimensional matrices in order to provide an environment
conducive to nerve regeneration. Although this experiment did test the effect of a combination of
peptides, it is likely that proper nervous tissue growth needs several additional factors or specific
ratios of signals.
The density of agarose gels can also affect the degree of nerve regeneration. Labrador et al.
created 4 mm gaps in the sciatic nerve in mice and bridged all of the injuries (except controls) with
silicone tubes filled with 0.5, 1, or 2% agarose gels. Functional recovery of muscle, skin, and sweat
gland innervation was then evaluated up to 120 days post-implantation. After about 40 days, the
agarose-filled silicone tubes began to influence functional recovery relative to controls. Although all
of the agarose gels improved muscle, skin, and sweat gland response, the 0.5% agarose gels
significantly outperformed the 2% gels by restoring gland recovery to 71% and skin pinprick
responses to 78%; these values for the 2% gels were 38 and 30%, respectively. The 1% gels tended to
elicit a response greater than 2% gels but not statistical different than 0.5% gels. Consequently, the
ability of neurons to grow depends on the density of the matrix in which it grows as suggested above
by studies with methylcellulose and laminin-rich scaffolding. If a scaffold has too high density, the
pore sizes might be too small for adequate infiltration or the stiffer mechanical properties may
physically inhibit growth. Thus, the addition of factors that enhance neurite growth could impede
regeneration if the size of concentration greatly affects the density of the polymer matrix [239].
6.3.3. Tubes and matrices
Collagen-coated porous PGA tubes, with inside diameters of 4±5 mm, were implanted into
25 mm cat sciatic nerve defects. Control tubes were left unfilled while other tubes were filled with
100 mg 2.5S nerve growth factor, 10 mg basic fibroblast growth factor, and Matrigel1. Five months
after implantation, results showed functional and morphological recovery for both the control group
and the growth factor filled tube group with the diameter and number of regenerated axons being 70
and 130% that of normal nerves, respectively [240]. In this case, the presence of growth factors did
not appear to have an effect on nerve regeneration. This lack of interaction may have resulted from
an overwhelming influence (either negative or positive) of the collagen coating. Overall, more
information on the effect of noncollagen-coated tubes is necessary to evaluate the actual effect of the
collagen coating on nerve regeneration.
Rodriguez et al. studied poly(L-lactide-co-e-caprolactone) tubes with 1 mm i.d. and a 150 mm
wall thickness as an alternative biodegradable tubulation devise. These tubes were filled with
Matrigel1 and Schwann cells and evaluated for peripheral nerve regeneration in 6 mm gap junctions
within mouse sciatic nerve. Functional comparisons between groups having Matrigel1 filled tubes
only, Matrigel1 plus autologous Schwann cells, and Matrigel1 plus allogeneic Schwann cells
indicated that tubes filled with Matrigel1 and autologous cells provided the best healing response.
Functional recovery of autologous Schwann cell containing grafts was comparable to that of
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autologous nerve grafts; however, the group with autografts regenerated faster and achieved higher
levels of mylinated axons. The lower healing response with allogeneic Schwann cells may have
resulted from an immune response to the cells [241]. Although little information on the effectiveness
of poly(L-lactide-co-e-caprolactone) as a tubulation material is given, nerve regeneration did occur
within these tubes. An alternative degradable polyester, poly(L-lactide, D,L-lactide 90:10), was
processed into flat sheets. Since Schwann cells did not adhere to the naked polymer, the polymer
surfaces were rendered hydrophilic through plasma polymerization of O2 for 5 min. Following
plasma polymerization, the materials were coated with a 100 mg/ml poly(D-lysine) solution for 1 h at
378C. These treatments improved Schwann cell adhesion by approximately 7.5 times that of the
untreated polymer. Seven-day embryonic chick DRGs were then cultured on the Schwann cell layers
or on the plasma/poly(D-lysine)-coated polymer to evaluate neurite extension. On Schwann cellcoated polymers, neurite outgrowth typically exceeded 2 mm within 48 h, which was approximately
14.7 times greater than without Schwann cells [242]. The results from the above two studies
demonstrate the usefulness of Schwann cell co-culture for nerve regeneration and indicate that the
Schwann cells likely secrete factors necessary for proper axon extension.
A novel approach to nerve regeneration involves the preliminary growth of a pseudo-nerve in a
silicone tube followed by grafting of the pseudo-nerve in a sciatic nerve defect. To develop this
approach, an allogenic pseudo nerve was generated by grafting a silicone tube in a sciatic nerve gap;
the nerve stumps were sutured 4, 10, or 15 mm apart with care taken to sever the proximal axons far
from the end of the tube. A pseudo-nerve containing longitudinally arranged Schwann cells and an
organized perineurium-like tissue was formed in gaps of 4 and 10 mm. The 10 mm pseudo-nerve
was then grafted into a rat sciatic nerve defect, and controls were grafted with donor nerve.
Morphologically, at 4, 6, and 8 weeks, both the pseudo-nerve and the grafted nerve demonstrated
similar degrees of axon regeneration. Pinch tests at day 6 also showed similar results between the
two experimental groups. This method suggested that appropriate scaffolding can be generated by
nerve stumps without innervation and that this scaffolding can then be used for nerve regeneration.
At this time, however, the inability to grow pseudo-nerves longer than 10 mm remains one limiting
factor. Moreover, there exists also questions regarding potential immunogenicity of the pseudonerve graft material since the grafts would have to be generated in animals for transplantation into
humans [243].
6.4. Neurotrophic factors
Cao and Soichet have developed degradable microspheres for the controlled release of NGF for
peripheral nerve regeneration. Several formulations with varying release times were made from
PLGA and poly(e-caprolactone) (PCL). These formulations included PLGA 50:50, PLGA 85:15,
PCL and PCL:PLGA 50:50 (1:1, w/w). Microspheres were made by a solvent evaporation process,
which resulted in spheres with average diameters of 17±24 mm as determined by SEM at 1 kV and
Coulter counter analysis. Degradation studies revealed varying degradation rates for the different
polymer blends. Over all release was dependent on the degradation rate of the polymer microsphere.
After 84 days, PLGA 50:50 degraded completely at 84 d, PLGA 85:15 degraded 90% PCL by
25% and PCL:PLGA by 30%. The majority of the NGF was released from the spheres within the
first 12 days due to an initial burst release of proteins from the microsphere surface. However,
release of active NGF was sustained for up to 91 days as determined by neurite extension in PC12
cells cultured in vitro. The degree of neurite extension correlated with the rate of NGF release
from the microspheres [244]. NGF has also been studied in the brain as a factor to induce
regeneration [216].
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Houweling looked at the effects of neurotrophin 3 (NT3) on spinal cord regeneration. NT3 was
dissolved in type I collagen at a concentration of 0.1 mg/ml and crosslinked under basic conditions.
No concentration of collagen was given. The gels were then implanted into the corticospinal track
between T8 and T10. The authors found that NT3 stimulated and directed regrowth of axons. They
also found improved functional recovery with NT3 application as opposed to collagen gels alone
[245].
Recently, PLGA foams were filled with inosine, a chemical with potential neurotrophic
properties, and used to investigate nerve growth across a 7 mm defect in the sciatic nerve of rats.
Prior to implantation, an in vitro evaluation of inosine release was performed using cylindrical 85:15
PLGA materials containing 0.76 wt.% inosine. Upon incubation in water, the authors noted an initial
bolus release of inosine equivalent to 3% of the total concentration. The inosine continued to steadily
diffuse from the foam material over a period of 67 days, suggesting a pattern of sustained release.
Furthermore, the release of the neurotrophin correlated with the degradation of the polymer matrix.
In vivo experiments lasted 10.5 weeks and compared the repair tissue within PLGA conduits with
and without inosine. Cross sections of implants after both 6 and 10.5 weeks did not show any
significant difference between controls and grafts containing inosine. The presence of inosine did,
however, induce the growth of neurons with larger diameter fibers. Even though the chemically
loaded materials contained neurons with larger widths than control tubes, functional recovery was
not statistically different [246].
6.5. Micropatterning
Micropattening of surfaces may provide substrates for controlled neurite adhesion and
controlled guidance. To this end, a negative polydimethylsiloxane mold was made with 12, 30, 50
and 70 mm wide lines separated by 250 mm gaps. After O2 plasma treatment made it hydrophilic, the
mold was placed onto a film of PLA-PEG-biotin. Avidin was then bound to the biotin. Avidin
contains four binding sites that allow for binding to the biotin surface and further binding with an
additional biotin bound bioactive substrate. In vitro evaluation of PC12 cells cultured on 50 and
70 mm patterned surfaces coupled to avidin-IKVAV revealed both cell specific adhesion to IKVAVcontaining regions and guided neurite extension along these patterns [247]. PDMS microcontact
printing has also been used to pattern IKVAV-containing peptides and laminin onto silicone-oxide
based chips. Grids consisting of 3, 5, 8, and 10 mm lines were constructed, and resulting surfaces
were used to culture hipocampal neurons. On 5, 8, and 10 mm surfaces, axonal elongation was
observed along the patterned surface in conjunction with the formation of neuronal junctions [248].
The 3 mm patterning was apparently too small for proper cell alignment.
As discussed above, degradable polymers allow for improved nerve regeneration when
compared with silicone tubes. Thus, when evaluating a new system on nerve regeneration, an
optimal degradable polymer should be combined with appropriate fill materials and biological
signals as determined in other experiments. We have learned that matrix-filled tubes are superior to
saline filled tubes, but that very low concentrations of proteins (on the order of 1±2 mg/ml) should be
used [249]. Furthermore, there exists some data suggesting that collagen acts as a better fill material
than laminin [249], while other data suggest the usefulness of fibrin as a fill material [237].
Neurotrophic factors can be incorporated into the fill materials to further improve regeneration.
Nerve growth factor, for example, has been shown to improve regeneration [244] as has neurotrophin
3 [245]. Other growth factors and peptides, however, can also stimulate neurite outgrowth.
As we think about which factors to incorporate into a system to influence regeneration, we also
must remember that factors that slow growth may also play an important role. It is believed that the
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native extracellular matrix of nerves contains signals that both retard and enhance growth. As a
result, the proper balance of these signals may well be necessary for optimal nerve regeneration.
7. Liver regeneration
There exists a great need for partial liver function support and for the development of artificial
livers that provide full function. According to the American Liver Foundation, 25 million Americans
suffer from gallbladder and liver disease. Annually, 27,000 people die of chronic liver disease while
only 3000 receive liver transplants. There are currently few effective treatments for the most sever liver
diseases. One exception includes liver transplantation. Unfortunately, the demand for liver transplants
far exceeds the limited supply. In an effort to bridge the gap between the number of needed liver
transplants and the amount of available organs, tissue engineers attempt to combine the use of polymers with hepatocytes to develop partial function liver support systems and fully functional livers.
The difficulty of liver regeneration results in part from the vast complexity of the tissue. Fig. 9
illustrates a small portion of liver and emphasizes the intricate organization of this organ. The liver is
the largest internal organ and is composed of glandular tissue. The hepatic artery provides
oxygenated blood to the liver to support tissue maintenance and growth. The hepatic portal vein
Fig. 9. A pen drawing of liver by Ashley adapted from [4]. The complexity of the structure of this organ is immediately
apparent. The liver is the largest internal organ and is composed of glandular tissue. The hepatic artery provides
oxygenated blood to the liver to support tissue maintenance and growth. The hepatic portal vein carries blood from the
digestive tube, pancreas and spleen and is the source for blood requiring cleansing. The liver extracts toxic substances from
the blood and degrades them. Some of the degradation biproducts are carried from the liver to the gallbladder by bile ducts.
After concentration in the gallbladder, the bile returns to the liver and the common bile duct then delivers it to the
duodenum. The cells of the liver, hepatocytes, are organized in folded sheets that line the sinusoids. The sinusoids are the
blood filled cavities that perfuse the liver with venous and arterial blood. Contaminant exchange between the blood and
bile via the hepatocytes occurs in the Space of Disse.
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
carries blood from the digestive tract, pancreas, and spleen and is the source of blood requiring
cleansing. The liver extracts toxic substances from the blood and degrades or metabolizes them.
Some of the degradation byproducts from the liver are carried to the gallbladder via bile ducts. After
concentration in the gallbladder, the bile returns to the liver, and the common bile duct then delivers
the bile to the duodenum. The cells of the liver, hepatocytes, are organized in folded sheets that line
the sinusoids. The sinusoids are the blood filled cavities that perfuse the liver with venous and
arterial blood. Contaminant exchange between the blood and the bile via the hepatocytes occurs in
the Space of Disse [4,112]. As can be deduced from the complex structure of the liver, simple
hepatocyte regeneration will not recreate full liver function, however, it does provide a step closer
towards the ultimate goal.
The field of liver regeneration and liver function support remains one of the most complex for
tissue engineers. The liver consists of several tissue types arranged in a complex architecture. Many
researchers have begun the difficult task of replacing hepatocyte function without being concerned
with developing a new liver; while others couple hepatocyte culture with endothelial cell culture to
begin to develop fully functional, vacularized livers. Davis and Vacanti provide a review of the
literature prior to 1996 [250]. They address microcarrier and encapsulated systems as well as
biodegradable scaffolds for hepatocyte regeneration.
8. Surface chemistry
As seen with virtually all cell types requiring adhesion to a substrate, hepatocyte behavior is
influenced by the surface to which they attach. The surface of a polymer affects the ability of the
cells to attach to the material as well as the metabolic activity of hepatocytes. This interaction may
result from both morphology and chemistry of the surface. Polymer membranes, polysulfone,
sulfonated polysulfone, cellulose acetate, and aminated cellulose acetate membranes were cast from
dimethyl formamide. Hepatocytes were then seeded from either serum free medium or medium
containing serum. The results from the serum free medium indicated that significantly more cells
attached to the polysulfone membranes. The mechanism for this interaction has not yet been defined.
The authors also found increased metabolic activity in cells attached to the polysulfone membranes,
in medium containing serum, as indicated by urea synthesis and ammonia levels in solution [251].
Often the presentation of bioactive sequences within proteins depends on the scaffolding
material used. Past research has documented well that peptide and protein conformational changes
occur when the molecules adhere to a surface. This alteration occurs as a direct result of
thermodynamic interactions between the protein molecules, the surface, and the solvent.
Conformational changes can cause the exposure of protein domains that are usually cryptic or a
variation in the conformation of a bioactive site, which in turn increases or decreases biological
activity. Bhadriraju and Hansen addressed some of these issues by using three different RGD
containing molecules on one surface and one RGD containing molecule on two different polymeric
surfaces. They allowed fibronectin, PronectinFTM, and a 23 amino acid sequence to adhere to
Immulon II plastic dishes. All sequences supported well-spread morphology of hepatocytes with
DNA synthesis. Interestingly, the short peptide supported rounded morphology when bacteriological
plastic dishes were used. The water contact angle for the bacteriological plastic was 918 and 708 for
the Immulon [252]. The difference in hydrophobicity of the surfaces may play a role in the adhesion
of the peptides and how the peptides are presented to the cell surface. Most likely, a combination of
base materials and proteins will need coordinated optimization to achieve the proper cellular
response.
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8.1. Morphology
Pore size also plays a role in hepatocyte function. Flat sheets and foams with controlled porosity
exhibit varying hepatocyte adhesion and function. Ranucci and Moghe addressed the question of
polymer morphology on cell attachment and metabolism using 50:50 PLGA flat sheets and porous
sponges. Hepatocytes were seeded on scaffolding of varying pore sizes: 3 mm pores promoted a twodimensional cellular arrangement; 17 mm pores promoted both two and three-dimensional
arrangements; 67 mm pores promoted both two- and three-dimensional arrangements with improved
albumin secretion. Microporosity (1±100 mm) significantly improved cell attachment and albumin
secretion. As a result, when designing biomaterials for liver regeneration, one must clearly consider
pore size as an important parameter [253].
Karamuk et al. confirmed the importance of pore size. They studied hepatocyte morphology on
PLGA and poly(N-p-vinylbenzyl-D-lactoamine)-coated PET fabrics with different pore sizes. The
mesh size of each fabric was between 50 and 800 mm. After seeding these scaffolds with
hepatocytes, they examined spheroid aggregate formation, which suggested favorable biological
interactions. Large mesh sizes (780 mm) supported the formation of aggregates measuring
approximately 150 mm. Overall, the authors found improved aggregation on materials with larger
pore size, which may be attributed to the initial number of cells seeded within the pore. They did not
present any functional studies [254].
Just as pore size plays a significant role in cell morphology and function, bioactive signals also
influence cellular behavior. Meyer et al. fabricated two-dimensional scaffolding with defined
porosity and studied hepatocyte morphology as a result of pore size, seeding density, and cytokine
presence. Poly(ethylene terapthalate) was coated with either PLGA or PDLLA. The pore sizes in
the PET fabric ranged from 50 to 700 mm, and the resulting coated scaffolds had similar pore sizes.
The authors found that the presence of epithelial growth factor and poly(N-p-vinylbenzyl-Dlactoamide), a b-galactose carrying polymer, improved hepatocyte adhesion, while pore size
influenced the morphology and kinetics of cell aggregation. Hepatocytes cultured on scaffolding
with pore sizes in the range of 700 mm showed larger, spherical aggregates. Those cultured on
scaffolds with 50±200 mm pore sizes revealed smaller, rough aggregates. Pore size does indeed play
a large role in overall cell morphology, which affects overall cell function [255]. This study does not
clearly indicate how the presence of degradable scaffolding, e.g. PLGA and PDLLA, influences
hepatocyte behavior. However, it does demonstrate that the presence of bioactive factors is crucial to
cell attachment.
8.2. Regeneration scaffolds
Mooney et al. fabricated highly porous PLGA sponges using particulate leaching technology.
Hepatocytes were then seeded onto the scaffolding followed by implantation into the mesentery of
Lewis Rats. Control groups for the experiment were simply seeded and implanted into rat mesentery,
and two experimental groups were prepared. One group was shunted to the inferior vena cava for
access to neurotrophic factors, while the other group was implanted into the mesentery of rats
subjected to 70% liver transection. In all cases, 95±99% of the hepatocytes died within 24 h. Of the
seeded cells, only 1:3 1:1% of the control group and approximately 6% of the experimental groups
remained after 1 week [256]. One crucial finding of this research showed that hepatocytes do not
survive when seeded within the interior scaffolding. This effect likely occurred due to the absence of
nutrient and waste product transport. The liver is a highly vascularized organ in which no cell is
located more that one cell length away from a blood supply. Consequently, successful liver
B.L. Seal et al. / Materials Science and Engineering R 34 (2001) 147±230
engineering likely will not occur unless materials incorporate designs that can support the
development and vascularization of the entire organ structure.
The presentation, conformation, and density of any ligand is crucial. Work by Griffith and
Lopina suggested that proper conformation of a ligand may lead to full effectiveness at lower
concentrations. Similarly, high densities of a ligand that are not fully accessible to the cell may not
prove sufficient to induce the proper response. In their experiments, galactose was tethered to PEG to
induce hepatocyte adhesion. By altering the tether length and density of galactose, they induced
hepatocyte spreading via the asialoglycoprotein receptor [257]. Gutsche et al. coupled lactose and
heparin to polystyrene foams (with pore sizes up to 100 mm). The lactose and heparin bound to
hepatocytes through the asialoglycoprotein receptor. In contrast to Lopina and Cima, they found that
hepatocytes tended to assume rounded morphologies rather than spreading on the surface. In
addition, the hepatocytes formed aggregates in and around the pores. For up to 1 week, the cells
showed albumin production and testosterone metabolism. Both of these biochemical indicators are
indicative of differentiated hepatocytes [258]. Due to its short-term support of hepatocyte function,
this type of nondegradable porous scaffolding may be useful in vitro during acute liver failure or as a
temporary support system until a patient receives a transplant.
Due to the high water content of soft tissues, hydrogels are frequently used for soft tissue
regeneration. Kneser et al. used highly porous, crosslinked poly(vinyl alcohol) (PVA) as a
scaffolding for hepatocytes. These matrices were implanted into the small intestine mesenteric
leaves of Syngenic male Lewis rats and infused with 10% of the recipient's liver mass. Implants were
removed at 1 week, 1 month, and 1 year post-implantation. Control groups did not contain any
hepatocytes and were devoid of hepatocytes at all times points. In addition to the controls, two
groups were implanted. One group had access to hepatotrophic factors from the portal vein, whereas
the other group did not. After 1 week and 1 year, both groups expressed albumin gene transcript
levels equivalent to normal liver. However, at 1 year, only those hepatocytes exposed to
hepatotrophic factors expressed normal levels of albumin [259]. This research suggests that
continued hepatocyte function depends upon access to appropriate factors. This work was also the
first to demonstrate long-term hepatocyte survival in vivo at levels sufficient to replace normal liver
function. Future work by this group includes designs that incorporate controlled release of necessary
hepatotrophic factors and work with degradable matrices, which may help to reduce the foreign body
response.
Another important technology involves the two-dimensional culture of hepatocytes. Due to the
complexity of the structure and function of the liver, successful organ development will require many
years of research. Until that time, we need technologies that can support and improve the life of
patients today. Two-dimensional culture of hepatocytes on membranes and on hollow fibers
demonstrates one method to bridge the technology gap. These culture systems can support
hepatocytes for short periods of time. During this time, blood is transported along the opposite side
of the membrane, and the hepatocytes act to remove and process toxins or molecules from the blood.
This approach offers some hope to patients undergoing acute liver failure or liver transplant surgery.
Several polymers have been studied for these systems and include both biopolymers and synthetic
polymers.
Traditionally, type I collagen has been the standard material for hepatocyte culture. For this
reason, the results of hepatocytes cultured on many polymers are compared to those of hepatocytes
cultured on collagen. Chitosan appears to provide better scaffolding for two-dimensional hepatocyte
culture than collagen. Since they themselves are too fragile for cell culture, chitosan gels need to be
crosslinked with gluteraldehyde. When the metabolic activity of hepatocytes cultured on these
biological materials was evaluated, lactate dihydrogenase (LDH) activity decreased, indicating less
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cell damage. Relative to collagen, chitosan materials promoted both a higher urea synthesis and a
more rounded cellular morphology typical of that seen in vivo [260]. Packed beds of polyvinyl
formal resin also have been studied for hepatocyte immobilization. Results indicated that
immobilized cells retained a rounded morphology. Urea synthesis and nitrogen metabolism dropped
rapidly over the first week and then stabilized at about 10% activity, a level slightly higher than cells
cultured on collagen-coated plates. In a similar manner, albumin secretion dropped to 60% over the
first 3 days before stabilizing. This value, however, was much higher than the amount of albumin
secreted by hepatocytes cultured on collagen [261].
Other types of materials used to restore liver function include microcarriers. These materials are
conducive for in vitro perfusion systems as well as in vivo systems. In vitro systems offer the
advantage of decreasing immune and foreign body responses. Kino et al. used cellulose
microcarriers (nylon mesh) with 100 mm pores for the immobilization of hepatocytes. They found
that the cells could be immobilized deep within the pores and that the cellulose acted as an artificial
extracellular matrix that supported rounded hepatocytes. Ammonia metabolism and glucose
synthesis were similar to those of hepatocytes grown in monolayers. When incorporated into an in
vitro perfusion system, the hepatocytes within the microcarriers maintained functionality for 9 h, a
sufficient duration for a patient with acute hepatic failure [262].
9. Conclusion
In recent years, tissue engineering has emerged as a revolutionary research thrust. Often
combining the techniques of material science and engineering, medicine and molecular and cellular
engineering, tissue engineering aims to restore, repair, or replace diseased or damaged tissue. As
previously presented, several researches have investigated the potential for regenerating tissues such
as skin, cartilage, bone, blood vessels, nerve, and liver using polymeric devices. Currently, most
tissue engineering research involves the use of biocompatible scaffolds that allow cells of a specific
phenotype to grow in a more natural, three-dimensional environment. One limitation of this
technique, however, remains the lack of control over precise biochemical signals. As a result,
engineered tissues contain regions resembling native tissue, but lack complete biochemical and
mechanical similarity. To overcome this obstacle, research has attempted to redesign scaffolds to
mimic a more natural extracellular environment that will provide specific biological signals for
tissue growth and reorganization. In order to achieve this goal, the marriage between biology and
tissue engineering needs to grow stronger.
Acknowledgements
We would like to thank Sue Ellen Panitch for drawing all of the detailed tissue figures.
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