The discipline of engineering translates fundamental knowledge in physics,
chemistry, and biology into materials, devices, systems, and strategies to
achieve practical benefits. This discipline also includes the systematic
definition and assessment of each variable that may contribute to the success
or failure of any engineering effort. Tissue engineering applies this
conceptual framework to advance the repair, replacement, or regeneration of
organs and
tissues1-3.
Current tissue-engineering strategies include transplantation of whole organs
or tissues with use of pedicle flaps and microvascular techniques,
transplantation of thin sections of tissues (e.g., split-thickness skin
grafts), transplantation of cell suspensions (e.g., blood transfusions or bone
marrow transplants), and endoprosthetic replacement of tissues. In
orthopaedics, bone4,
cartilage5,6,
tendon7,8,
ligament9, meniscus,
intervertebral
disc10,11,
fat, muscle12, and
nerve are the primary targets.
In recent years, the options for orthopaedic tissue engineering have
increased dramatically. These options include methods for harvest and
transplantation of tissue-forming cells, the use of an expanding array of
bioactive matrix materials as tissue scaffolds, local or systemic delivery of
commercially available peptide hormones and growth factors, and other methods
to control the local chemical and biophysical environment. These new options
highlight a transition from the historically materials-based tissue-level
approach, with which mechanically durable, bioinert, or biocompatible
materials were preferred, to a focus on cell-based or bioactive materials and
stimuli. This evolving approach focuses on the function of cells and the role
of materials, implants, and biophysical stimuli in modulating cell
function.
Cell-based tissue-engineering tools and methods create exciting new
opportunities that might be useful in a broad array of potential clinical
applications. These opportunities also precipitate a critical need for
orthopaedic surgeons to participate actively in the design, development,
critical evaluation, and informed use of these methods. Active participation
requires that orthopaedic surgeons have a solid foundation in the contemporary
concepts and principles of cell-based tissue engineering. This article reviews
the central paradigms of contemporary tissue engineering. Specifically, it
addresses stem cells and progenitor cells in musculoskeletal tissues (the
cells responsible for all new tissue formation), strategies for the clinical
use of these cells, barriers to cell transplantation and cell survival, and
strategies and variables in the design and optimization of cell-based
tissue-engineering scaffolds.
There are four major types of cell-based tissue engineering: (1) local
targeting of connective tissue progenitors where new tissue is needed, (2)
transplanting autogenous connective tissue progenitors to augment the local
population, (3) transplanting culture-expanded or modified connective tissue
progenitors, and (4) transplanting fully formed tissue.
Targeting Connective Tissue Progenitors In Situ
Targeting strategies are designed to promote desired tissue formation by
stimulating the activation, migration, proliferation, and/or differentiation
of local connective tissue progenitors. Implantation of acellular tissue
scaffolds (e.g., allograft bone, ceramics, hyaluronic acid, and synthetic
polymers) is an example of this strategy. The strategy relies on a sufficient
local population of connective tissue progenitors. Tissue scaffolds provide a
surface on which cells and connective tissue progenitors can attach and
migrate as well as a protected void space in which new tissue can form and be
distributed throughout the region where new tissue is desired. When these
properties promote bone-healing, they are referred to as
osteoconduction46,47.
However, the concept of tissue conduction can be applied equally well to any
desired tissue-engineered phenotype.
Locally delivered growth factors (e.g., bone morphogenetic proteins [BMPs],
fibroblast growth factor-2 [FGF-2], and vascular endothelial growth factor
[VEGF]) also target local cells. The capacity of some growth factors to
selectively activate bone-forming connective tissue progenitors and/or enhance
the probability that their progeny will differentiate into bone has been
defined as osteoinduction. Again, the concept of induction can apply
equally well to stimuli that promote activation and differentiation of
connective tissue progenitors toward any desired phenotype.
Biophysical stimulation, such as mechanical
loading48-50,
electromagnetic
stimulation51,52,
or ultrasound53, is
also an example of cell targeting. Systemic pharmacological strategies, such
as the use of parathyroid hormone for the treatment of
osteoporosis54-59
or the use of systemic growth hormones to induce an increase in muscle mass in
the
elderly60,61,
are types of cell-targeted tissue engineering as well.
Transplantation of Autogenous Connective Tissue Progenitors
Transplantation of connective tissue progenitors was designed to compensate
for a deficiency in the number or function of local connective tissue
progenitors, as may occur in regions of previous trauma, infection, previous
irradiation, tissue defects, scar, or compromised vascularity. Transplantation
of connective tissue progenitors can improve the outcome of both conductive
and inductive grafts, even in sites that are surrounded by non-diseased
tissues13,62.
This suggests that many, and perhaps all, settings of normal tissue repair may
be limited by the population of connective tissue progenitors in local
tissues.
Autogenous cancellous bone-grafting has long been the most prevalent and
relatively effective example of cell transplantation, although only a small
fraction of the transplanted cells actually
survive63-65.
In the past decade, several additional transplantation strategies have been
introduced. Several uncontrolled clinical studies have suggested that
transplantation of connective tissue progenitors in aspirated bone marrow has
value in bone-healing
applications22,66,67.
In an uncontrolled, nonrandomized consecutive study, Connolly et
al.68,69
found bone marrow injection to be successful in the treatment of eighteen of
twenty tibial nonunions. Connolly et
al.25 also showed
that concentration of bone marrow cells with use of a centrifuge could
increase osteogenesis further, a strategy that has been supported by other
investigators70,71.
Many surgeons now use bone marrow because of its biological value and low
risk. One of us (G.F.M.) has had clinical experience (albeit without
independent or detailed prospective documentation) with aspiration of bone
marrow in more than 900 patients undergoing elective orthopaedic procedures
over the past fifteen years. The aspiration volumes in the patients ranged
from 16 to 200 mL. Patient cohorts representing a subset of this experience
have been reported on in two
publications24,33.
There were only two reported bruises, no hematomas, no infections, and no
chronic pain at the aspiration site. On direct questioning, most patients
reported no pain at the aspiration site during their hospitalization, and in
no case was the bone marrow aspiration site the reason for the patient taking
pain medication, a factor limiting rehabilitation, or the cause for a delay in
discharge from the hospital.
The aspiration technique is important. One of us (G.F.M.) and
colleagues24 found
that limiting the volume of the aspirate to =2 mL per site reduces dilution
with peripheral blood and significantly increases the concentration of
marrow-derived connective tissue progenitors (p < 0.001). Recent data have
shown that the efficacy of a bone marrow graft can be significantly enhanced
by the use of the surface of some porous implantable materials to selectively
concentrate and select marrow-derived connective tissue progenitors from bone
marrow (p <
0.001)62. Selective
retention of connective tissue progenitors can be used to rapidly enrich the
population of marrow-derived connective tissue progenitors, by removing red
blood cells, serum, and most other cells in marrow and contaminating blood.
Grafts enriched in this way have significantly improved the results of
bone-grafting in a canine spinal fusion model (p <
0.05)62 and have
been approved by the United States Food and Drug Administration for clinical
bone-grafting in spinal fusion and in treatment of bone fractures and defects,
although to date no clinical studies on this strategy have been published, to
our knowledge.
Transplantation of Culture-Expanded Autogenous Cells
Culture-expanded cells can also contribute to new tissue
formation42,72-74.
Preliminary studies have suggested that culture-expanded cells from muscle,
fat, and bone marrow may be useful in regeneration of bone, cartilage, muscle,
and tendon
tissue73,75-79.
In vitro expansion offers the potential to generate a large number of
progenitor cells. However, culture expansion also adds substantial cost and
some risks, such as contamination with bacteria or viruses or depletion of the
proliferative capacity of the connective tissue progenitors prior to
implantation80-82.
In vitro selection of the most rapidly proliferating cells may also select
cells with mutations or epigenetic changes that might confer a tumor-forming
potential. However, we are not aware of any reports of human tumors being
formed by culture-expanded cells, and the risk of tumor formation appears to
be very low. At present, the only Food and Drug Administration-approved
clinical use of culture-expanded cells in orthopaedics is for the repair of
cartilage defects, where expanded autogenous chondrocytes are transplanted
under a periosteal
flap6,83.
Transplantation of Genetically Modified Cells
The intrinsic biological potential and performance of connective tissue
progenitors can be genetically modified with a variety of means that either
transiently or permanently alter the genes that a cell
expresses84. Both
connective tissue progenitors and nonconnective tissue progenitors have been
engineered to secrete factors (e.g., BMP-2) that will influence the behavior
of that cell or the cells nearby. The introduction of new or modified genes is
usually accomplished with use of vectors that are created by
modifying naturally occurring viruses, such as a retrovirus, lentivirus,
adenovirus, or adeno-associated
virus85,86.
Nonconnective tissue progenitors have also been transfected to express genes
that control other genes. LMP-1 (LIM mineralizing protein-1) is one example.
LMP-1 is a nuclear transcription factor, a protein that, when expressed,
functions within the nucleus to activate or inhibit the expression of a number
of other genes. LMP-1 is therefore not secreted but functions only within the
nucleus of a cell that is transfected to express LMP-1. However, LMP-1
activity within the cell induces the secretion of a variety of pro-osteogenic
factors that can target osteogenic connective tissue
progenitors87,88.
The biological risk associated with genetic manipulation is greater than
that associated with the alternatives. As a result, demonstration of safety is
currently as great a challenge as is demonstration of biological
efficacy89. While
transplantation of genetically modified cells may not play a role in elective
clinical tissue engineering in the near future, it has substantial potential
value, particularly in the setting of inherited genetic defects (e.g.,
osteogenesis
imperfecta)90 and
in tissues, such as cartilage, that consist of relatively homogeneous
long-lived cells and in which phenotypic expression may be induced and durably
maintained by expression of a single
gene86,91.
Ex Vivo Tissue Generation and Transplantation
Creation of fully organized and mature tissues outside of the body (ex
vivo) followed by functional transplantation and integration is the common
public vision of tissue engineering. However, this strategy involves three
great challenges: (1) generation of functional tissues, (2) transplantation in
a manner that preserves the viability and function of cells, and (3)
biological and mechanical fixation and integration with surrounding
tissue.
In some areas, transplantation of thin tissue grafts (e.g., cartilage,
corneal, and skin grafts) is possible without an immediate connection to a
developed vascular system. However, in most tissues, cell survival requires a
functioning vascular system. Some reports describe ex vivo generation of
vascular transport
systems92,93,
but these approaches are not currently clinically practical.
In all settings in which cells are transplanted, access to substrate
molecules (oxygen, glucose, and amino acids) and clearance of products of
metabolism (CO2, lactate, and urea) are critical to cell survival.
The movement of these molecules in and out of the graft site is collectively
referred to as mass transport. Mass transport can be mediated by
fluid flow (convection), both in the circulatory system and in the
extracellular space, between the vessel lumen and the cell membrane. The
pressure gradients driving this fluid flow can be induced by tissue
deformation (movement), mechanical loading, muscle contraction, gravitational
pooling, Starling flow, and arterial pulsation. Convection can be particularly
important when cells are embedded in a dense extracellular matrix, such as
bone and cartilage, where it has been shown to play a major role in enhancing
the transport of large molecules (e.g., proteins and growth
factors)94-97.
However, in most tissues, passive diffusion along concentration gradients is
the principal mechanism for mass transport, particularly for small
molecules.
In metabolically active tissues such as trabecular bone and bone marrow,
the distance that oxygen must diffuse between a capillary lumen and a cell
membrane is almost never more than 40 to 200
µm98,99.
This diffusion distance is critical to maintaining the balance between oxygen
delivery to a site and consumption of oxygen by cells, both in native tissues
and in tissue engineering strategies involving cell transplantation. When
cells are transplanted clinically, the vessels that deliver oxygen are
initially confined to the outer surface of the graft site. As one moves deeper
into the graft site, each transplanted cell competes for oxygen and other
nutrients with other transplanted cells. Transplanted cells also compete with
other cells that are recruited as part of the local inflammatory response
following implantation. In most clinical grafts, the diffusion distance for
oxygen and other metabolites from the edge of the graft to the center of the
graft is a minimum of 5 mm, or approximately fifty times the normal diffusion
distance. In this setting, diffusion is able to support only a limited number
of transplanted cells before the balance between metabolic demand and
diffusion creates a zone in the center of the graft where oxygen tension is
too low to support viable cells, resulting in central necrosis. The size of
the necrotic region and the number, distribution, and type of cells that do
survive in the deeper regions of the graft site are a function of many
variables, which can be analyzed with use of basic engineering principles. The
principal variables are the concentration of oxygen at the surface of the
graft site (CO), the concentration and distribution of cells in the
site (including inflammatory cells), the rate of oxygen consumption by cells
within the site, fluid flow within the site, the diffusion constant for
oxygen, and the biological response of cells (survival, proliferation,
migration, and differentiation) to hypoxia. Relatively little is known about
the balance between diffusion and consumption of proteins, peptides, or the
signaling molecules. However, because diffusion of oxygen is relatively slow
and oxygen consumption is high, the transport of other nutrients (e.g.,
glucose and amino acids) is generally more favorable than that of oxygen.
Oxygen is therefore the limiting factor in cell survival in most grafts. As a
result, few cells tolerate diffusion distances of >0.2 mm. For example, rat
osteoblasts seeded on porous scaffolds in vitro form a viable tissue that is
no greater than 0.2 mm
thick100. Islet
cells show necrosis when the diffusion distance exceeds approximately 0.1
mm101-103.
Cartilage is exceptional, maintaining viability in avascular regions >1 mm
thick104, although
oxygen transport in cartilage in vivo may be enhanced by convective
flow95,105.
Theoretical modeling can be used to explore the relationships among cell
density, diffusion distance, and cell viability within a graft.
Figures 5-A and 5-B illustrate
a theoretical cellular implant, where diffusion of oxygen in the x-direction
is balanced by cellular consumption. When the differential equation describing
the balance between diffusion and reaction is made nondimensional, a
parameter, denoted ?2, emerges as the relative rate of reaction
to the relative rate of diffusion. When ?2 = 1, the oxygen
concentration in the center of the graft is 50% of the concentration at the
surface, and when ?2 = 2, the oxygen concentration in the
center of the graft is zero. A rough estimate of the limit of cell density and
diffusion distance can be made by setting ?2 = 1. Assembling
these variables, oxygen consumption averages about 4 × 10-17
mol/cell-sec but varies with cell type (hematopoietic stem cells, 0.47 ~
3.3 × 10-17
mol/cell-sec106;
fibroblasts, 4 ~ 7 × 10-17
mol/cell-sec107;
and granulocytes and monocytes, 0.6 ~ 18 × 10-17
mol/cell-sec98,99,107).
The diffusion coefficient of oxygen in tissue is ~ × 10-5
cm2/sec at
37°C98,99,107,
and oxygen concentration in normal tissues (CO) is about 0.07 mM,
slightly greater than that in venous plasma. Using these values for a graft of
a given thickness, one can estimate the maximum concentration of cells that
can be delivered without central necrosis ([Cell]max). For example,
in a graft that is 2 cm thick, [Cell]max is roughly 70,000
cells/cm3. In contrast, a 1-cm-thick graft could support four times
more cells, or 280,000 cells/cm3. This is about 1000-fold lower
than the concentration of cells in native autogenous cancellous bone (~
× 108 cells/cm3) and 100-fold lower than the mean
concentration of cells in a marrow aspirate (~ × 107
cells/cm3)24.
These estimates predict hypoxia and central necrosis in almost any graft site
with a diffusion distance of more than 500 to 1000 µm. They also define
what is approximately an inverse square relationship between
[Cell]max and diffusion distance. For example, increasing a graft
dimension by a factor of 5 (e.g., transitioning from a rat to a dog) will
decrease [Cell]max by a factor of 25. This is one reason why many
cell transplantation methods work very well in small animals but fail in
larger animals and humans.
Several factors may modify these calculated estimates. First, because cell
delivery systems are generally prepared in room air, the initial implant is
usually saturated with oxygen. This dissolved oxygen will support cell
respiration for at least several hours after implantation, blunting the abrupt
decrease in regional oxygen concentration. Second, not all implanted cells
continue to respire at basal rates. Cells that are very sensitive to the
trauma of transplantation may die, promptly reducing the initial metabolic
demand within the implant. However, cell death also results in the local
release of products of cell lysis and adds debris to the site. This may
increase the intensity of local inflammation and the metabolic demand from
inflammatory cells at a later time.
The survival of transplanted connective tissue progenitors also depends on
the response of these cells to the transplantation environment. Observation
has long supported the concept that at least some connective tissue
progenitors in bone and bone marrow have a high capacity to survive in hypoxic
conditions63,64,108.
Experimental data also have shown that many stem and progenitor cells,
including connective tissue progenitors in bone, exhibit a remarkable
tolerance to, and are even stimulated
by109-113,
hypoxia, not unlike endothelial
cells114,115.
The capacity to convert to glycolysis transiently in response to hypoxia is
one adaptive
mechanism116. The
rate and extent of revascularization are also critical. Prompt
revascularization favors osteoblastic differentiation, whereas prolonged
hypoxia favors formation of cartilage or fibrous
tissue17,117-119.
These concepts can be converted into several practical strategies to
optimize cell survival in clinical grafts. One method is to reduce the
concentration of transplanted cells. Another is to limit the transplanted
cells to only those cells that contribute to the formation of the desired
tissue (i.e., connective tissue progenitors and perhaps endothelial cells),
while excluding red blood cells and the vast majority of other nucleated
cells. Both
concentration25 and
selection
strategies62 have
been shown to enhance graft performance. Finally, in the future,
culture-expanded cells might be preadapted to hypoxic conditions prior to
transplantation and/or selected to enrich for those most likely to survive.
Other options to improve the local matrix environment and to enhance mass
transport are discussed below.
Three-dimensional porous scaffolds play a critical role in both cell
targeting and cell transplantation strategies. Scaffold matrices serve as
space-holders to prevent encroachment of surrounding tissues into the graft
site. They provide surfaces that facilitate the attachment, survival,
migration, proliferation, and differentiation of stem cells and progenitors.
They also provide a void volume in which vascularization, new tissue
formation, and remodeling can occur (Fig.
6). In addition, scaffolds can provide a vehicle for delivery of
cells into a graft site, facilitating their retention and distribution
throughout the region where new tissue is
desired1,46,120-122.
A broad range of scaffolds is already available for clinical use, and many
new scaffolds are under development. Differences between scaffolds can
generally be categorized into one or more of six domains: bulk material,
three-dimensional architecture and porosity, surface chemistry, mechanical
properties, initial scaffold environment (osmolarity and pH), and late
scaffold environment (degradation characteristics). Each domain has important
implications with respect to the biological response to a scaffold and its
utility in transplanting or supporting local stem cells and progenitors.
Bulk Materials
Current clinical scaffolds are made from a broad range of bulk materials.
These include tissue-derived materials (e.g., allograft bone matrix, skin, and
intestinal submucosa), biological polymers (e.g., collagen, hyaluronan,
fibrin, and alginate), ceramics or mineral-based matrices (e.g., tricalcium
phosphate, hydroxyapatite, and calcium sulfate), metals (e.g., titanium,
tantalum, and other alloys), and composites of two or more materials. A
variety of synthetic polymers are also being adapted or developed. These new
materials include water-insoluble polymers (e.g., poly[lactide], polytyrosine
carbonates, poly[caprolactone], varying copolymers, and synthetic gel-like
polymers [polyethylene
oxide-based])123.
Three-Dimensional Architecture and Porosity
Matrix architecture refers to the way in which a bulk material is
distributed in space, at the nanoscale, microscale, and macroscale (i.e.,
molecular, cellular, and tissue-length scales, respectively). Matrix
architecture defines the mechanical structure of the scaffold, but it also
defines the initial void space that is available for connective tissue
progenitors to form new tissue, including new blood vessels, as well as the
pathways for mass transport (convection and diffusion). Most scaffolds are
designed to have an internal porous structure of void spaces that are
interconnected through pores or channels on the scale of 50 to 1000 µm. The
pore size used for most bone ingrowth settings is between 150 and 500 µm,
which is just large enough to support ingrowth of vascular tissues, depending
on the depth of penetration required. Figures
7-A,
7-B, and
7-C illustrate the process of
bone ingrowth into a surface with a pore size of ~ 50 µm. Larger pores
generally support deeper penetration of new tissues, but the optimal pore size
for ingrowth deeper than 3 to 4 mm into tissue scaffolds has not been studied
systematically, to our knowledge. This is relevant to the current clinical
practice of filling large voids with particulate or granular materials, since
the void spaces between packed particles are generally an order of magnitude
larger than the stated microstructure or pore size of most granules
themselves. This may provide the larger macrostructure needed for deeper
revascularization.
Options for the structural design of tissue scaffolds are almost infinite.
The macrostructures include regular geometric shapes (e.g., blocks, pellets,
and dowels), amorphous structures (e.g., randomly packed chips, granules, or
fibers), randomly integrated structures (e.g., foams or freeze-dried
materials), and formally designed regular structures (e.g., machined, printed,
woven, or assembled structures). Gel or putty preparations can also be made
from powders or fibers, by mixing them with plasticizing agents (e.g.,
glycerol, cellulose, and hyaluronan) or by conducting in situ polymerization
with use of chemical, photochemical, or enzymatic
methods120. In
some cases, a desirable structure has been borrowed from nature, such as the
highly interconnected porous structure of cancellous bone or some
corals124.
Selective processing (machining, size and density selection, washing, and
demineralization) now provides a variety of relatively optimized materials for
use in special clinical settings.
Most methods for fabricating porous scaffolds—i.e., particulate
leaching100,
freeze-drying125,
gas infusion126,
and phase
separation120—create
isotropically distributed voids and connecting pores (such as in a sponge) by
using particles or bubbles when the scaffold is solidified. In the past
decade, substantial advances have been made in the methods for producing more
precise hierarchical microstructures from a variety of materials. These are
now being applied to create strategically oriented channels and pores and
defined macroscopic shapes. The most notable innovations involve solid
free-form fabrication methods (e.g., three-dimensional printing process and
stereolithography)1,120,
which provide a feature resolution of approximately 200 µm. Creation of
more defined porous structures offers the potential for greater control over
the distribution of bulk material within a graft site as well as control over
patterns of cell migration, fluid flow, and diffusion throughout the
device127-129.
Nanostructural features (<100 nm) may also play an important role in
scaffold function. Nanopores are too small to influence where cells can or
cannot migrate, but they may still have important effects on cell behavior by
changing surface texture or diffusion of soluble materials. All other features
being equal, the presence of interconnected nanoporosity within the walls of a
porous structure can open up a much wider path for mass transport, thereby
improving cell survival in the scaffold.
Mechanical Properties
Sometimes, graft sites must bear loads at, or close to, physiological
levels very soon after implantation. Internal fixation often provides the
necessary early stability. However, in some bone or soft-tissue settings, the
scaffold must bear or share substantial load immediately, and then
high-strength materials and structures such as cortical bone, metals,
ceramics, or carbon-fiber-based polymers are required.
A scaffold's mechanical properties (strength, modulus, toughness, and
ductility) are determined both by the material properties of the bulk material
and by its structure (macrostructure, microstructure, and nanostructure).
Matching the mechanical properties of a scaffold to the graft environment is
critically important so that progression of tissue healing is not limited by
mechanical failure of the scaffold prior to successful tissue regeneration.
Similarly, because mechanical signals are important mediators of the
differentiation of connective tissue progenitors, a scaffold must create an
appropriate stress environment throughout the site where new tissue is
desired.
One of the greatest challenges in scaffold design is the control of the
mechanical properties of the scaffold over time. Scaffolds that do not degrade
(metals and ceramics) simplify this problem and can provide excellent and
durable function in some settings. However, these materials can also
compromise tissue repair and function. It is obvious that persistence of a
scaffold or implant precludes the formation of new tissue in the space that it
occupies. In addition, following integration of a rigid nondegradable implant,
adjacent tissue is often mechanically protected (stress-shielded), changing
local mechanical signals and resulting in loss of desired local tissue. Stress
concentration at the interface between a high-stiffness implant and native
tissue can increase the risk of mechanical failure (e.g., fracture) and pain.
Finally, if subsequent procedures require removal of the implant scaffold
(e.g., because of infection or migration), all of the new tissue within the
implant may also be lost, eliminating the value of the initial procedure.
Problems arising from retained implants have increased the desire to use
resorbable scaffolds whenever feasible. One example of that strategy is the
use of impaction grafting for reconstruction of contained periprosthetic
defects. Another is the recent shift from the use of very slowly degradable
ceramics (e.g., hydroxyapatite) for bone-void fillers to the use of more
rapidly resorbed materials (e.g., tricalcium
phosphate)130.
This demand for resorbable scaffolds continues to fuel the development of
resorbable inorganic
polymers120.
Almost all new materials being developed for tissue engineering ultimately
resorb, and three key features of the degradation or resorption process appear
to influence performance: the rate at which the matrix loses its mechanical
properties, the rate at which the matrix is removed from the site, and the
nature and concentration of the soluble products that are released into the
site as the material is broken down.
Controlled degradation of mechanical properties alone is a major challenge.
Mechanical properties can be lost as a result of internal degradation of the
bulk material (e.g., hydrolysis) or the accumulation of fatigue damage. In
recent years, it has become more evident that, even in controlled settings,
the mechanical and chemical degradation of the same polymer can vary
substantially between species, individuals, anatomic locations, and clinical
settings. As a result, it has been very difficult to define an optimal
degradation rate for materials to be used in general clinical practice. In
general, most design strategies tend to extend degradation time over months,
in order to minimize the risk of early failure in preference to minimizing the
risks associated with delayed
resorption131.
Surface Chemistry
Interactions between cells and scaffolds occur at the surface and are the
direct result of the unique chemical environment that is created. The surface
chemistry depends on the properties of the bulk material but is not defined by
the bulk material. This is due to the fact that almost all implanted materials
rapidly become coated with proteins and lipids, and these adsorbed
biomolecules are the principal mediators of the cellular response to most
materials. The net effect involves an interaction between a given surface and
available biomolecules that adsorb to the surface. Furthermore, when a protein
adsorbs, it usually undergoes a change in conformation, which may include
denaturation or unfolding. This, in turn, may either hide or expose sites
within the protein that interact with cell surface receptors. For example,
fibronectin is a more active adhesion molecule on hydrophilic surfaces (e.g.,
glass) than on hydrophobic surfaces (e.g., Teflon or
polyethylene)132-135.
Biological fluids contain a vast diversity of proteins, and cells have
hundreds of different types of cell surface receptors. There are twenty-four
distinct cell-matrix receptors in the integrin family
alone136. As a
result, it is not surprising that scaffold materials have been discovered and
selected empirically. However, powerful surface analytical techniques are
being used to illuminate the protein adsorption properties on various surfaces
in an effort to find out why some materials are so favorable for bone cell
adhesion and bone
formation137. For
example, it has been speculated that hydroxyapatite and some other ceramics
may preferentially sequester growth factors, growth factor-binding proteins,
or adhesion molecules that are important for bone regeneration. Indeed,
hydroxyapatite and tricalcium phosphate materials perform successfully as
depot delivery vehicles for BMPs, in both
animals138 and
humans139,140.
Like other implanted materials, allograft bone matrices (both mineralized
and demineralized) rapidly accumulate biomolecules on their surface, which
have biological effects on local cells. However, allograft bone already
contains many embedded adhesion molecules and growth
factors141. These
include the BMPs, although the concentration is far lower than that delivered
with use of purified recombinant BMP products and the release is much slower,
requiring matrix degradation by local cells. Furthermore, the concentration
and presentation of bioactive molecules in allograft bone may vary widely
depending on the age, gender, and genetics of the donor; the tissue site of
origin; and the tissue processing
procedures142.
The attachment, survival, proliferation, and differentiation of stem cells
and progenitor cells can be modulated in vitro if implants are precoated with
selected bioactive
proteins143,144.
Furthermore, proteins can be selectively concentrated and presented on
surfaces with use of nonspecific surface interactions (e.g., dip-coating or
lyophilization). This is the strategy that is used to deliver BMP-2 (Infuse;
Medtronic Sofamor Danek, Memphis, Tennessee) and OP-1 or BMP-7 (OP-1 Device;
Stryker Biotech, Hopkinton, Massachusetts) in two clinical products that are
currently available for improving
bone-healing145-147.
In the case of the BMP-2 product (Infuse), BMP-2 is provided in solution and
is dripped into an absorbable collagen sponge. BMP-2 binds to the collagen and
is then released. In the case of the OP-1 (BMP-7) product (OP-1 Device), 3500
µg of OP-1 is lyophilized onto 1 g of bovine type-I collagen powder. OP-1
is then released by solubilization from the collagen surface.
The pharmacokinetics of delivery of BMPs have been shown to be an important
clinical variable in a variety of materials, including degradable
polymers148-155,
type-I
collagen156-158,
and calcium phosphate
ceramics159,160.
The retention time of implanted BMPs have been shown to correlate with
biological efficacy, presumably because the longer a BMP is retained, the
higher the likelihood that it will act on an appropriate target cell. Protein
residence time can be estimated in vivo by measuring the rate of clearance
following the implantation of radioactive protein, with the assumption that
the protein remains active as long as residual radiation can be measured.
Retention has been shown to be related to
solubility161 and
protein isoelectric
point138. The
release kinetics of BMP-2 from a degradable collagen sponge in rabbits was
described by Bouxsein et
al.162.
Approximately 25% of the delivered BMP-2 was released rapidly from the
implantation site, but as much as 37% remained at the site one week after
implantation.
The advantage of the current clinical strategies for protein delivery is
that they are technically simple. However, they require that the protein be
delivered in a high concentration. This delivery strategy presumably allows
the protein to diffuse into tissues adjacent to the implant site to act on the
local connective tissue progenitor target population. This release pattern may
also establish a concentration gradient around the graft site that may be
important for chemotactive factors. However, the disadvantage of current
strategies is that they provide relatively little control over the rate of
delivery, conformation, presentation, clearance, or degradation of the
delivered protein. While current strategies for delivery of BMPs can be
effective, there is reason to believe, given the very large supraphysiological
doses of BMPs that are needed, that the vast majority of the protein that is
delivered is wasted and that only a small fraction actually comes into contact
with target cells to elicit a receptor-mediated signal that enhances new bone
formation. These methods therefore leave substantial room for improvement in
the delivery kinetics (rate and duration) and distribution of bioactive
proteins to optimize efficiency.
Proteins and small bioactive peptides can also be selectively concentrated
and presented by covalently linking them to a
surface1,120,140.
This provides more control over conformation, a slower rate of release from
the surface, and longer retention. While this method may not be appropriate
for many soluble proteins, particularly signaling molecules that need to be
internalized for function (e.g., steroid hormones), there is growing evidence
suggesting that presenting growth factors in a matrix-bound fashion not only
may be suitable for clinical use, but also may better mimic the native
physiology and improve
outcomes163-166.
This strategy may also be particularly well suited for the design of matrices
with selective affinity for specific cells or sets of cells (e.g., connective
tissue progenitors, endothelial cells, and platelets).
Initial Scaffold Environment: Osmolarity and pH
A scaffold that is used as a delivery system for viable cells must provide
and maintain an environment with physiological pH and osmolarity. For most
scaffolds, simple hydration with normal saline solution prior to exposing them
to cells avoids cell injury. However, some matrices do not allow an isotonic
condition to be created for cell delivery. Examples include many bone matrix
materials that are prepared with use of solutions containing high
concentrations of low-molecular-weight materials to improve handling (e.g.,
glycerol) and materials that dissolve rapidly in water, releasing
hyperos-molar concentrations of local ions (e.g., calcium sulfate). If such
materials are mixed with cells, they can be expected to induce osmotic injury,
reducing or precluding cell viability. Matrices containing
high-molecular-weight carriers (e.g., cellulose, starch, and hyaluronan) may
be acceptable. There is much less osmotic pressure (proportional to the molar
concentration of the solute) with these high-molecular-weight carriers.
Late Scaffold Environment: Degradation Products
All degradable matrices release degradation products into the graft site
environment that must be further degraded or cleared. The effect that these
degradation products have on the cells within the graft site depends on their
concentration, their effect on local pH, and their relative biological
toxicity. Concentration, in turn, is a function of the rate of release of
these products and their rate of clearance from the graft site.
Polyesters, such as polylactides and polyglycolides, are currently the
so-called workhorses of synthetic degradable surgical materials. They have
been used for decades as sutures, surgical meshes, and more recently as
fixation hardware (e.g., suture anchors and
screws)120. The
degradation of these materials can be controlled over a range of weeks to
years. Polymer hydrophobicity and crystallinity both influence the rate at
which water penetrates and hydrolyzes the solid polymer and thus the rate of
device
breakdown140.
However, the degradation products of these materials (lactic acid and glycolic
acid) are not ideal for tissue regeneration. Furthermore, they tend to be
released as a bolus after a long period of residence. This is due to the fact
that they are degraded by hydrolysis in a process that first randomly degrades
the bulk polymer, progressively reducing the molecular weight and the
mechanical properties of the material but leaving the total mass of polymer
essentially the same until the molecular weight of the fragments that are
created is small enough to make them soluble. When this occurs, soluble
material is generated rapidly, liberating the bulk polymer into solution but
creating a profound local decrease in pH. Regression of local bone formation
and sterile cysts in bone and soft tissue have been commonly
observed15,167.
Furthermore, in some settings, the polymer can crystallize as it degrades,
creating particles that persist for years. In highly porous devices, the
effects of degradation products may be less pronounced because the volume
fraction of the bulk polymer in the graft site is smaller and the degradation
products that are released are cleared more readily through a broader surface
of contact with local extra-cellular fluids and vascular perfusion.
Regardless, the use of polylactides and polyglycolides for bone regeneration
remains controversial, and devices based on these materials are not widely
used in clinical practice.
The limitations of polyesters are being addressed by the development of new
classes of degradable materials that possess reasonable mechanical strength
and do not release acidic degradation products. One material consisting of a
copolymer of polyethylene oxide and polybutylene terephthalate showed
promising osteoconductive results in animals but failed to induce bone
formation in the iliac crests of
humans168. A
different class of materials, pseudo-polyamino acids, have been synthesized
with a range of degradation properties and have been shown to offer improved
behavior in bone sites in
animals169,170.
An innovative approach to matching degradation rate to tissue ingrowth is
being developed by Hubbell et al., using polyethylene oxide-based gels that
contain both cell adhesion molecules and other peptides, which are selected to
provide specific biological activity and are released as the gel is
degraded171,172.
The gels can be formed in situ, and they function essentially at a synthetic
extracellular matrix designed as a true surface eroding polymer. The gels are
degraded by specific proteases that are elaborated by cells as they invade the
structure. As a result, the scaffold is removed from the implantation site
gradually, in concert with cell invasion, and it maintains its mechanical
properties until it is degraded. This avoids the late burst release of
material that is characteristic of bulk eroding polymers.
Rapidly advancing knowledge and capabilities in many fields are driving the
next wave of tissue engineering strategies and products. Tissue engineering
strategies that are particularly rich in opportunity include (1) improved
methods for intraoperative selection and concentration of stem cells and
progenitor cells; (2) cell delivery systems that enhance the survival of
transplanted cells by managing the balance of mass transfer and metabolic
demand at the graft site; (3) three-dimensional scaffolds with architectural
and mechanical features that are customized for specific clinical
applications; (4) chemically defined surfaces that present covalently tethered
biologically active molecules (adhesion sites, growth factors, and synthetic
peptides) creating local concentrations and gradients to elicit desired cell
attachment, migration, differentiation, and survival; (5) defined
microtextured surfaces that elicit the desired cell attachment, migration,
differentiation, and survival; (6) scaffold materials that degrade in a manner
that delivers biologically inert or even biologically active molecules, rather
than molecules that may be harmful in the graft site; and (7) delivery systems
for soluble molecules (e.g., BMPs and other protein growth factors) that
ensure both a biologically active conformation and a local concentration or
concentration gradient that is appropriate for the target cell population,
minimizing the total dose of bioactive agent that is required and the risk of
unwanted collateral effects.
Optimizing combinations of cells, matrices, and locally and systemically
active stimuli will remain a complex process characterized by a highly
interdependent set of variables with an almost infinite range of possible
combinations. As a result, these developments must also be informed by a
combination of clinical experience, knowledge of basic biological principles,
medical necessity, and commercial practicality. The responsibility for
rational development is shared by the entire orthopaedic community
(developers, vendors, and physicians). The need for objective and systematic
assessment and reporting is made particularly urgent by the recent rapid
addition of many new options for clinical use. Prospective, randomized
preclinical and clinical trials will play a critical role in the initial
evaluation of new materials for specific indications. Prospective cohort
studies will also be valuable in several settings, including testing whether
the results of controlled studies can be generalized to the broader
orthopaedic community, defining settings where current practice may fall short
of reported or desired outcomes, and assessing settings where randomization is
impractical or unethical because of the absence of equipoise (i.e., the
absence of an alternative method for comparison that is perceived as being
comparable in effectiveness or morbidity).