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www.rsc.org/materialsChemistryMaterialsJournal of
Injectable scaffolds for tissue regeneration
Qingpu Hou, Paul A. De Bank and Kevin M. Shakesheff*
Tissue Engineering Group, School of Pharmacy, University of Nottingham, University
Park, Nottingham, UK NG7 2RD. E-mail: [email protected]
Received 2nd February 2004, Accepted 6th May 2004 First published as an Advance Article on the web 25th May 2004
Tissue engineering aims to develop functional substitutes
for damaged or diseased tissues through complex constructs
of living cells, bioactive molecules and three-dimensional
porous scaffolds, which support cell attachment, pro-
liferation and differentiation. Such constructs can be
formed either by seeding cells within a pre-formed scaffold
or through injection of a solidifiable precursor and cell
mixture to the defective tissue. As cell and bioactive
molecule carriers, injectable scaffolds are appealing,
particularly from the clinical point of view, because they
offer the possibility of homogeneously distributing cells and
molecular signals throughout the scaffold and can be
injected directly into cavities, even of irregular shape and
size, in a minimally invasive manner. In this paper the
challenges in designing an injectable scaffold from the
viewpoint of materials chemistry and the solidification
mechanisms of injectable precursors are discussed. The
applications of injectable scaffolds in angiogenesis, bone
repair and cartilage regeneration are described.
1. Introduction
The field of tissue engineering holds great promise for the
repair or regeneration of damaged and diseased tissues. 1 Its
underlying objective is to direct a population of cells into
forming a living tissue, structurally and functionally indis-
tinguishable from that found in nature. To guide the transition
from cells to tissue, a three-dimensional scaffold may be
utilized. This lends not only structural form to the cell mass,
but can positively influence cell adhesion, growth and
differentiation by the incorporation of adhesion molecules or
the controlled release of bioactive molecules from the scaffold.
As cells proliferate, deposition of extracellular matrix compo-
nents and biodegradation of the scaffold results in a solid,
three-dimensional tissue construct. The seeding of cells into
such scaffolds can be performed in two distinct ways. Firstly,
cells can be expanded in culture, seeded onto the scaffold
and allowed to mature in vitro before implantation into the
patient. Alternatively, the scaffold can be implanted to fill a
void in damaged tissue and subsequently seeded by the
infiltration of the patient’s own cells. For the latter strategy,
the scaffold can either be a pre-formed, three-dimensional
porous structure or an injectable scaffold; a mixture of bio-
active molecules and solidifiable precursors, which are injected
into the defect and form a three-dimensional structure in situ.
This review will examine the benefits of injectable scaffolds,
the materials chemistry challenges faced in their design, the
Qingpu Hou received his Bachelor (1989)
and Master (1992) degrees from Tianjin
University, China. He attained his Ph.D
from the Institute of Chemistry, the
Chinese Academy of Sciences in 1995.
He is now a research fellow in the School
of Pharmacy at the University of
Nottingham, working on polymer synthe-
sis, modification and processing for tissue
engineering applications.
Paul De Bank is a postdoctoral research
fellow in the School of Pharmacy at the
University of Nottingham. Since complet-
ing his PhD in 2000 on the synthesis and
evaluation of novel potential cannabinoids,
he has worked on the spatially controlled
formation of neuromuscular junctions, non-
viral gene delivery systems for the tissue
engineering of bone and is currently
investigating novel systems for the genera-
tion of functional multicellular organoids.
Kevin Shakesheff is Professor of Tissue
Engineering and Drug Delivery at
the School of Pharmacy, The University
of Nottingham. He became interested
in tissue engineering during his time
as a NATO Postdoctoral Fellow at
MIT in the mid-1990s. His research
focuses on the role of 3D culture
environment on the behaviour of cells
and the engineering of polymers to create
biomimetic environments for tissue
regeneration.
DOI
: 10.1039/b401791a Qingpu Hou
Paul De Bank
Kevin Shakesheff
T h i s j o u r n a l i s ß T h e R o y a l S o c i e t y o f C h e m i s t r y 2 0 0 4 J. M a t e r. C h e m. , 2 0 0 4 , 1 4 , 1 9 1 5 – 1 9 2 3 1 9 1 5
solidification mechanisms employed and their applications in
tissue engineering.
- The clinical need for injectable scaffolds
From a clinical perspective, the use of injectable scaffolds is
very attractive as it minimizes patient discomfort, risk of
infection, scar formation, and the cost of treatment. In the case
of the preformed scaffolds, prior knowledge of the size and
shape of the defect or cavity to be filled is necessary, and defects
with irregular shape and size can prove problematical. In
addition, invasive surgery for implantation of the construct is
required. Moreover, cell seeding methods can be inefficient due
to poor transport of cells through the matrix and cellular
damage.^2 The use of injectable scaffolds can overcome these
limitations. By virtue of the scaffold components being in
suspension or solution before solidification in vivo, a more
homogeneous distribution of bioactive molecules within the
matrices can readily be obtained. What is more, the nature of
these systems makes it possible to co-inject a cell suspension
with the scaffold components, resulting in a cell–scaffold
construct that can fill any size or shape of cavity with minimally
invasive surgery. 3–5^ After injection and solidification an in situ
forming scaffold provides a temporary 3-D matrix on which
the cells can adhere, proliferate and differentiate, forming a
new, functional tissue. 6 Hence, injectable scaffolds are pro-
mising matrices for tissue induction or regeneration, especially
for engineering bone and soft tissues. In addition to serving as
carriers for bioactive molecules, injectable scaffolds can also
act as conduits for the guidance of tissue regeneration, tissue
adhesives for healing and injectable controlled release devices
for local drug delivery. 7,8^ Table 1 summarizes the injectable
materials that have, to date, been utilized for tissue engineering
applications. Prior to injection, they may be in the form of
solution, paste, micro or nanoparticles, beads, or thread-like
material 4 and can be cell-free systems or cells and/or tissue
growth factors suspension systems.
- Materials chemistry challenges in designing injectable scaffolds
Cell viability and function within an injectable scaffold are
closely related to the physical, chemical and biological charac-
teristics of the scaffold used. From the viewpoint of materials
chemistry, several requirements must be met during the design
and fabrication of such a scaffold, including:
. Nontoxic and sterile components,
. Injectability,
. Solidification under mild conditions and cohesivity,
. Mechanical strength and resistance to in situ forces,
. Biodegradation,
. Pore morphology,
. Incorporation of bioactive molecules.
These factors will be considered individually.
Nontoxic and sterile components
Injectable scaffolds should not be deleterious to the health of
both cells and tissue. Each component of the formulation, the
Table 1 Injectable scaffolds reported for tissue regenerationa
Injectable scaffolds Solidification mechanism References
Inorganic materials Calcium phosphate Ceramics setting 9– Natural polymers Chitosan Thermal gelation 17, Methylcellulose Thermal gelation 5 Alginate Photo cross-linking 19 Alginate Ionic gelation 20– Hyaluronic acid Photo cross-linking 19,24, Agarose Thermal gelation 26, Fibrin Thermal gelation 28– Gelatin Thermal gelation 31 Synthetic polymers Poly(aldehyde guluronate) Chemical cross-linking 32 PEG or PEO Photo cross-linking 33– PEO-PPO-PEO Thermal gelation 43, PEO-PLLA-PEO Photo cross-linking 45 PLA-g-PVA Photo cross-linking 45 PEO-PLLA Thermal gelation 46 PLGA-PEG Thermal gelation 47 PLLA-PEG Photo cross-linking 35 PEG-co-Poly(a-Hydroxy Acid) Photo cross-linking 48 PVA Photo cross-linking 49 PLAL-ASP Photo cross-linking 50 P(CL/TMC) Photo cross-linking 51– PLA(Glc-Ser) Photo cross-linking 54 Polyanhydrides Photo cross-linking 55, PPF Photo cross-linking or radical polymerization 57– OPF Photo cross-linking or radical polymerization 66– P(PF-co-EG) Photo cross-linking or radical polymerization 48,62,69– PhosPEG-dMA Photo polymerization 77 PNIPAAm -PEG Thermal gelation 78 PNIPAAm-gelatin Thermal gelation 31 P(NIPAAm-AAc) Thermal gelation 79, PEG based hydrogels Enzymatic cross-linking 81 PEG based hydrogels Michael-type addition reaction 82– PLA-PEG-biotin Self-assembly 2 a (^) Abbreviations: OPF: Oligo(poly(ethylene glycol) fumarate); P(CL/TMC): Poly(-caprolactone-co-trimethylene carbonate); PDLLA:
Poly(D , L-lactide); PEG: Poly(ethylene glycol); PEO: Poly(ethylene oxide); PEO-PPO-PEO: Polyethylene oxide-polypropylene oxide-polyethylene oxide; PhosPEG-dMA: Poly(ethylene glycol) di[ethylphosphatidyl(ethylene glycol)methacrylate]; PLA(Glc-Ser): Poly(L -lactic acid-co-glycolic acid-co- L -serine); PLA-PEG: Poly(lactic acid)-poly(ethylene glycol); PLAL-ASP: Poly(lactic acid-co-lysine)-poly(aspartic acid); PLGA: Poly ( DL- lactic-co-glycolic acid); PLLA: Poly(L -lactic acid); PLLA-PEG: Poly(L -lactide-ethylene glycol); PNIPAAm: Poly(N-isopropylacrylamide); P(NIPAAm-AAc): Poly(N-isopropylacrylamide-acrylic acid); PPF: Poly(propylene fumarate); P(PF-co-EG): Poly(propylene furmarate-co- ethylene glycol); PVA: Poly(vinyl alcohol).
Ceramics setting
Calcium phosphate cements (CPCs) can undergo a self-setting
process within the body after injection, based upon the cement-
ing action of acidic and basic calcium phosphate compounds
on wetting with an aqueous medium. 10 Within a few minutes,
mixing of the cement formulation leads to a solidifying mass
due to crystallization of dahllite. 9,11^ The setting time can be
adjusted by addition of manipulator compounds to the wetting
medium. Recently, a fully injectable calcium phosphate cement
formulation was developed by incorporation of a biocom-
patible gelling agent. The resultant CPC had significantly
improved injectability and cohesive properties.
Thermally or photochemically activated radical polymerization
or cross-linking
Precursors with unsaturated or photosensitive functional
groups can form gels by thermally or photochemically activ-
ated radical polymerization or cross-linking. 19,40,42,73,109,110,
Radicals produced by an initiator or photoinitiator react
with the functional groups of the macromonomers to
cause polymerization or cross-linking, leading to gel forma-
tion. In tissue engineering applications, the most commonly
used macromonomer functional groups are (meth)acry-
loyl,19,35,42,45,48,50,54,107,112^ styryl,^112 coumarin,51,52^ phenyla-
zide, 53 and fumaryl. 66 Fig. 1 illustrates the chemical
structures of some macromonomers that can be polymerized
or cross-linked by this mechanism. 107,113^ The solidification
process is determined by a number of factors including reac-
tivity, functionality, concentration and molecular weight of
the precursors, intensity of visible or UV light, temperature,
reaction time, and the type and concentration of the initiator.
Thermal gelation
Some polymer solutions undergo gelation triggered by a
change in temperature. Typical thermal gelling polymers
include copolymers of N-isopropylacrylamide, poly(ethylene
glycol) (PEG)-based amphiphilic block copolymers, gelatin,
agarose, and cellulose. 5 As one of the most intensively
investigated thermosensitive polymers, N-isopropylacryl-
amide-based copolymers exhibit a sol–gel transition as the
temperature is increased above their lower critical solution
temperature (LCST) due to the drastic solubility difference of
these polymers below and above the LCST. The gelation is
related to the chain entanglement and the gradual chain
collapse as the temperature increases. 114 At room temperature,
the polymer solutions are transparent and remain fluid, while at
37 uC, the matrices become opaque and form gels without
significant gel induction time. The transition temperature
can be further tuned by changing the composition of the
copolymers. Once the gels are formed, they do not change their
water content and the gelation is reversible without appreciable
hysteresis. The factors determining the gelation process include
polymer concentration, molecular weight, and chemical struc-
ture of the copolymer.
Recently block and star copolymers of poly(ethylene glycol)
and poly(N-isopropylacrylamide) of various architectures were
synthesized and their gelation behaviour studied. At low
temperature, they form liquid aqueous solutions with low to
moderate injection viscosities, but form relatively strong elastic
gels upon warming to physiological temperature. It is believed
that the linear copolymer formed a weaker gel by micellar
aggregation, while the star-shaped copolymers formed a strong
network gel via a physical cross-linking mechanism. This
gelation is rapid and shows a low degree of syneresis. 78
Fig. 1 Chemical structures of materials that can be thermally or photochemically activated, polymerized or cross-linked to create hydrogel networks. (A) Poly(ethylene glycol) (PEG) diacrylate ( 1 ), methacrylate ( 2 ), and propylene fumarate derivatives ( 3 );. (B) Cross-linkable poly(vinyl alcohol) derivatives ( 4 and 5 ). (C) Polysaccharide derivatives: methacrylate-modified dextran ( 6 ) and cinnamated hyaluronic acid ( 7 ). (D) Dimethacrylated polyanhydrides: poly(sebacic acid) ( 8 ), poly(1,3-bis(p-carboxyphenoxy)propane) ( 9 ) and poly(1,6-bis(p-carboxyphenoxy)- hexane) ( 10 ). Adapted from refs. 107 and 113.
Besides N-isopropylacrylamide-based copolymers, other
typical examples of thermosensitive polymers are poly(ethylene
oxide) and poly(propylene oxide) copolymers known as
poloxamers or pluronics. 43 Although these two kinds of poly-
mers have been used as injectable scaffolds, their applications
in tissue engineering are limited due to their non-degradability
and toxicity. 46 To overcome these limitations, biodegradable
PEG-based copolymers such as linear or star-shaped poly-
(ethylene glycol- L -lactic acid) and poly(ethylene glycol- DL -lactic
acid-co-glycolic acid) 46,47,99,115^ have been developed. At a high
concentration, diblock or triblock copolymers of ethylene
glycol and lactic acid form a gel at a lower temperature and
become a sol at a higher temperature. The gelation is thought
to be caused by the association of micelles. The sol–gel
transition temperature can be tuned by varying the biodegrad-
able block length and polymer concentration.
As well as synthetic polymers, aqueous solutions of natural
biopolymers, such as gelatin, agarose and cellulose, can also
form gels in response to temperature changes. Similarly several
factors, such as polymer concentration, formulation of the
aqueous solvent, and heating rate, have an effect on the gela-
tion temperature. 5 Recently, a chitosan/glycerophosphate
disodium salt solution was reported to undergo a thermogelling
process by a combination of three molecular forces: hydrogen
bonding, electrostatic interactions and hydrophobic inter-
actions. 17 The formulations remain liquid at physiological
pH below room temperature, but form gels rapidly if heated to
body temperature. Another approach to the formation of
thermogelling chitosan is through the incorporation of thiol
groups into the polymer.^18
Ionic cross-linking
Aqueous solutions of alginate can form gels in the presence of
di- or trivalent cations. Alginate is a linear polysaccharide
composed of 1,4-linked b- D-mannuronic acid (M) and a- L-
guluronic acid (G) residues in varying proportions and
sequential arrangements. The gelation process is controlled
by the cation type and concentration, alginate composition and
concentration, and gelation temperature. 116 The gelation rate
increases with increasing concentration of multivalent cations
such as Ca 21 in the system. Alginate with a higher G content or
longer G segment sequences gels faster. On the contrary, an
increase in alginate concentration leads to a decreased gelation
rate. The gelation rate also has an effect on the uniformity and
the subsequent mechanical properties of the resultant hydro-
gels. A lower gelation rate generally yields hydrogels with a
more homogeneous structure and increased mechanical
strength. In addition, mechanical strength increases with
alginate concentration, total calcium content, molecular
weight and G content of the alginate. Recently, thermally
triggered release of Ca 21 from liposomal compartments was
used to induce rapid gelation of alginate. 117
Michael-type addition reaction
Novel hybrid PEG-peptide gels are formed upon stepwise
copolymerization of multi-arm vinyl sulfone-terminated poly-
(ethylene glycol) macromonomers with bis-cysteine oligopep-
tides via Michael-type addition reactions. 82 The architecture of
the networks can be tailored by variation of the functionality
and molecular weight of the precursor macromonomers.
Besides the macromonomer structure, the preparation condi-
tions including the stoichiometry of the reactive groups,
precursor concentrations and the pH during cross-linking
play a role in the gelation rate of the system and the mechanical
characteristics of the resultant hydrogels. By this approach,
hydrogels containing protease-sensitive sequences were formed
within a few minutes at 37 uC in a humidified incubator at pH
8.5. Recently, based on this mechanism, in situ cross-linked
biomaterials have been developed for hard tissue repair
using water-insoluble precursors in dispersion and reverse
emulsion. 118
Self-assembly mechanism
Recently, a novel injectable scaffold was developed based on a
self-assembly mechanism. 2 Polymer particles composed of
poly(lactic acid)-poly(ethylene glycol)-biotin (PLA-PEG
biotin) were mixed with a suspension of cells in an appropriate
cell culture medium and then co-injected with the cross-linking
protein avidin. By optimizing the concentration of avidin,
cross-linking of the microparticles occurred within seconds of
mixing of the components and a scaffold formed around the
cells. Additionally, biotinylated peptides can also be introduced
to surface engineer the particles and promote integrin-mediated
cell adhesion. Cell culture experiments demonstrated that this
self-assembly cross-linking process did not interfere with cell
function.
In general, chemical cross-linking is highly versatile for the
preparation of injectable scaffolds, and the resulting networks
possess superior mechanical strength. However, toxic chemical
agents are often employed in the formulations, adversely
affecting cells and bioactive molecules during solidification.
Physical cross-linking can overcome these limitations, but the
resultant networks usually possess limited mechanical properties
and stability.^119 Therefore careful selection of precursor
formulation and suitable cross-linking methods are crucial to
the preparation of injectable scaffolds for specific tissue
regeneration. The remainder of this review will highlight specific
applications of injectable scaffolds in tissue engineering.
- Tissue regeneration applications
Angiogenesis
Angiogenesis, the formation of new blood vessels, is a key
process in tissue regeneration. To achieve this, sustained release
of certain growth factors such as vascular endothelial growth
factor (VEGF) and basic fibroblast growth factor (bFGF) can
be employed. However, due to their short half-life and easy
diffusion in vivo, an appropriate delivery system is needed to
enhance the efficacy of the growth factors for highly localized
angiogenesis. Injectable scaffolds have been studied as such a
delivery vehicle due to their easy preparation and handling.
Several matrices have been developed such as alginate/heparin
microparticles, 120 alginate microspheres or beads, 121,
sodium hyaluronate, 123 and PLGA millicylinders. 124 Injectable
gels based on fibrin have also been prepared as angiogenic
growth factor delivery carriers. Using such a system, VEGF
was covalently conjugated to a fibrin network and exhibited
prolonged presentation and delivery within the carrier
matrix. 28 Interestingly, the growth factor release was a result
of matrix degradation induced by cell infiltration. The released
VEGF maintained its bioactivity and was able to trigger the
stimulation of endothelial cell proliferation, enhancing the
angiogenic process. Alternatively, angiogenesis-promoting
fibrin-based matrices can be constructed by covalent modifica-
tion of adhesive domains. For example, avb 3 integrin is
predominantly located on the surface of angiogenic endothelial
cells. When a specific receptor for avb3 integrin, L1Ig6, was
covalently attached to a fibrin matrix, the resultant scaffold
was shown to promote angiogenesis in angiogenic cell types,
but not in control cells. 125 To combine cells in injectable
scaffolds, novel synthetic hydrogels based on a Michael-type
addition reaction have recently been prepared in the presence
of cells. The reaction occurred between a monofunctional cell
adhesion peptide, a difunctional protease substrate peptide
with sensitivity to matrix metalloproteinases (MMPs), and a
tetrafunctional poly(ethylene glycol) under mild conditions.
responded to local cellular stimuli, and degraded to soluble
products. When these gels were used as a delivery carrier for
recombinant human bone morphogenetic protein 2 (rhBMP-2),
the release of the entrapped growth factor was promoted by
matrix degradation induced by cell infiltration, as shown
in Fig. 3. In this way a highly localized BMP-2 release was
achieved, demonstrating that these gels are suitable delivery
matrices to induce bone regeneration.
Cartilage regeneration
A number of injectable scaffolds have been studied for the
purpose of cartilage regeneration, including oligo(poly-
(ethylene glycol) fumarate),^130 poly(N-isopropylacrylamide-
co-acrylic acid), 80 poly(N-isopropylacrylamide)-grafted
gelatin,^31 poly(ethylene oxide), 40,42,131^ alginate, 20,22^ fibrin,^30
PLGA-g-PEG, 47 pluronics, 43,44^ calcium phosphate/hyaluronic
acid composites,
132
hyaluronic acid gel
133–
and chitosan.
17
During cartilage repair, extracellular matrix formation is
strongly affected by the properties of the scaffolds, such as
swelling ratio, compression modulus, degradation rate and cell
seeding density. 30,131^ Therefore, careful control over the cross-
linking density and structure of the macromonomers is
necessary to achieve increased type II collagen synthesis and
homogeneous distribution of glycosaminoglycan (GAG)
within the engineered cartilage. When an injectable scaffold/
chondrocyte construct was further modified with a growth
factor delivery system, a significant increase in GAG produc-
tion was observed. 45 Chitosan/glycerol-phosphate (C/GP),
which undergoes a gelation process when it is heated to body
temperature, was also recently used as an injectable scaffold for
cartilage regeneration. Isolated bovine articular chondrocytes
within a C/GP composite were implanted subcutaneously in
athymic mice. It was demonstrated that the system facilitated
neocartilage formation after 3 weeks implantation, as shown
in Fig. 4.
- Final remarks and future perspectives
While initial studies are encouraging, there remain a number of
crucial challenges for the materials chemist before injectable
scaffolds can be clinically relevant. Key aspects that must be
fully optimized and controlled are the kinetics of scaffold
solidification, mechanical compatibility with the target tissue
and biodegradability of the scaffold. Of particular importance,
however, will be the ability to tailor injectable scaffolds to
control the release and bioavailability of incorporated growth
factors. This is dependent not only on the chemistry of the
matrix, but also the concentration and distribution of the
biomolecules in the scaffold and the nature of the target tissue.
The ability of the scaffold to release growth factors when
mechanically stimulated is one aspect of this that must also be
considered.
In conclusion, the ability to design a biomimetic, injectable
scaffold system with a defined release and degradation profile
has huge potential for the repair and regeneration of damaged
tissues. If the materials chemistry challenges can be overcome,
such tissue engineering devices should soon become a clinical
reality.
Acknowledgements
We thank the EPSRC for funding.
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