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nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 457
review article
Published online: 21 may 2009 | doi: 10.1038/nmat
A
human embryo in its first eight weeks of life undergoes an
extraordinary transformation from a single cell to a 3-cm-
long fetus with a beating heart, gut, nervous system, and
limbs with fingers and toes. This progression involves massive
growth, physical folds and twists, and myriad cellular and molecular
events of breathtaking complexity; yet it is the ultimate goal of tissue
engineering (TE) to recreate some of these processes in microcosm,
to replace and regenerate lost tissue. At last the field has entered a
period of fruition, and seems set to realize its potential to treat a
multitude of debilitating and deadly conditions such as myocardial
infarction, spinal injury, osteoarthritis, osteoporosis, diabetes, liver
cirrhosis and retinopathy. The general strategy is usually to seed cells
within a scaffold, a structural device that defines the geometry of the
replacement tissue and provides environmental cues that promote
tissue regeneration. TE skin equivalents have been in clinical use
since 1997 (ref. 1) and a fast-growing arsenal of replacement devices
is in clinical trials or already approved as therapies for tissues includ-
ing cartilage, bone, blood vessel and pancreas (Table 1). In two
recent high-profile studies, seven patients benefited from TE blad-
ders^2 , and a 30-year-old woman became the first person to receive a
TE tracheal segment, a procedure that saved her left lung 3.
Aside from the obvious human benefits, tissue engineering could
bring substantial financial rewards to those who succeed in trans-
lating this new technology to the clinic. Sales of regenerative bio-
materials already exceed US$240 million per annum^4 and the wider
markets that tissue engineering taps into are colossal: costs related
to organ replacement account for 8% of global healthcare spending,
and by 2040 as much as 25% of the US GDP is expected to be related
to healthcare 5. Nevertheless, if the short history of industrial tissue
engineering has taught us anything, it is that the provision of effec-
tive products is not in itself sufficient to ensure commercial success
(Fig. 1). Early TE efforts were plagued by product issues related to
scale-up, shelf-life, quality control and distribution, and suffered
from inappropriate business models and withdrawal of private
finance in the early 2000s 1,6. Since then the field has matured, evi-
denced by the return of large-scale investment and the first regen-
erative medicine companies becoming profitable^4.
Alongside these positive developments, progress in biomaterials
design and engineering are converging to enable a new generation of
instructive materials to emerge as candidates for regenerative medi-
cine. Which of these materials compete successfully in the market
will depend on a combination of clinical performance, marketing
and cost-effectiveness. A central dilemma is that to influence cell
behaviour, scaffolding materials must bear complex information,
Complexity in biomaterials for tissue engineering
elsie s. Place1,2, nicholas d. evans1,2^ and molly m. stevens1,
The molecular and physical information coded within the extracellular milieu is informing the development of a new generation of biomaterials for tissue engineering. Several powerful extracellular influences have already found their way into cell- instructive scaffolds, while others remain largely unexplored. Yet for commercial success tissue engineering products must be not only efficacious but also cost-effective, introducing a potential dichotomy between the need for sophistication and ease of production. This is spurring interest in recreating extracellular influences in simplified forms, from the reduction of biopolymers into short functional domains, to the use of basic chemistries to manipulate cell fate. In the future these exciting developments are likely to help reconcile the clinical and commercial pressures on tissue engineering.
coded in their physical and chemical structures. On the other hand,
financial considerations dictate that complexity must be kept to a
minimum. Clearly there is a danger, by over-engineering devices,
of making their translation to clinical use unlikely. The solutions
to this challenge lie at every phase of product development, begin-
ning with identifying the simplest functional performance required
to resolve a defined clinical problem. The ambitious early aims
of reconstructing entire organs have largely given way to smaller,
more attainable goals: for example, rather than trying to replace an
entire heart, clinical advances in cardiac repair focus on TE coro-
nary arteries, valves and myocardium. Organogenesis Inc. and
Advanced Tissue Sciences Inc. suffered heavily as a result of their
overestimating the number of chronic wounds cases that were best
solved by high-tech, TE skin substitutes (respectively, Apligraf and
Dermagraft; Dermagraft is now produced by Advanced Biohealing)
as opposed to acellular products that aid ongoing repair 6 (Table 1,
Fig. 1). Similarly, an emerging philosophy in tissue engineering
is that rather than attempting to recreate the complexity of living
tissues ex vivo, we should aim to develop synthetic materials that
establish key interactions with cells in ways that unlock the body’s
innate powers of organization and self-repair. In this review we will
consider how this can be achieved, emphasizing how even relatively
simple engineering solutions can deliver considerable functional
benefits. Along the way we will explore how some of these princi-
ples have been applied to specific scientific and commercial tissue-
engineering challenges.
regenerative potential of tissues
Even without any therapeutic intervention, living tissues can have
a staggering capacity for regeneration. For example, the human
liver will regrow to its original size even when more than 50% of
its mass is excised 7. This has been taken to the extreme in rats,
where one group has reported that a single rat’s liver was able to
regenerate fully following each of 12 sequential hepatectomies, a
finding that can be explained by the high replicative potential of
the cell types that make up the liver. Several other tissues — bone
and skin, for example — also have an innate capacity to regener-
ate to fill injuries below a critical size, helped by local or recruited
stem cells. The clinical potential of stem cells has long been rec-
ognized by haematologists, who in the 1960s showed that trans-
planted haematopoietic (literally ‘blood-making’) stem cells from
the bone marrow of a healthy mouse could replace the destroyed
immune system of another mouse, paving the way for a cure for
leukaemia 8,9^. The discovery of other types of cell with multilineage
(^1) Department of Materials; 2 Institute for Biomedical Engineering, Imperial College London, London SW7 2AZ, UK. e‑mail: [email protected]
458 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials
review article NaTure maTerIalS^ doi: 10.1038/nmat
potential has since followed, including neural stem cells from
the brain, and mesenchymal stem cells, which can differentiate
into bone, fat, cartilage and muscle cells 10,11^. Indeed, more recent
evidence suggests that stem cells or progenitor cells can be isolated
from almost every tissue of the body 12,13^. Under the correct condi-
tions, these cells can be stimulated to form new tissue, as we recently
demonstrated using a simple biomaterials-based approach. Here,
either calcium-crosslinked alginate gels or modified hyaluronic
acid gels were injected into an artificial space between the tibia
and the periosteum, the fibrous outer lining of bone. This stimu-
lated bone and cartilage formation from resident progenitor cells
in the inner layer of the periosteum 14 , illustrating that complex
tissues can be generated from relatively simple materials by using
the body as a ‘bioreactor’.
Table 1 | Commercial tissue engineering products and biomaterials at various stages of development.
tissue Product regulatory status
description material Cells use Form
Synthetic resorbable animal derived Plant or bacteriaderived Human derived Growth factor allogenic autologous Skin TransCyte, Advanced Biohealing
1997 Nylon mesh coated with porcine collagen, containing non‑viable human fibroblasts, with upper layer of silicon
✓ ✓ ✓ ✓ Burns Sheet
Apligraf, Organogenesis 1998 Lower layer of human fibroblasts and bovine collagen, upper layer of keratinocytes
✓ ✓ ✓ Leg ulcers Sheet
Dermagraft, Advanced BioHealing
2001 Cryopreserved human fibroblasts on a polyglactin 910 (2‑hydroxy‑propanoic acid polymer with polymerized hydroxyacetic acid) mesh
✓ ✓ ✓ Diabetic foot ulcers
Sheet
ICX‑SKN, Intercytex Phase II Allogenic fibroblasts and human collagen with additional layer of keratinocytes
✓ ✓ ✓ (^) Burns and acute wounds
Sheet
Integra Dermal Regeneration Template, Integra Lifesciences
1996 Porous bovine collagen crosslinked with chondroitin‑6‑sulphate with upper layer of silicon
✓ ✓ ✓ Burns Sheet
Integra Flowable Wound Matrix, Integra Lifesciences
2007 Granulated bovine collagen crosslinked with chondroitin‑6‑sulphate
✓ ✓ Ulcers Gel
Oasis Wound Matrix, Healthpoint
2006 a^ Decellularized porcine small intestinal submucosa
✓ ✓ Burns, ulcers, other wounds
Sheet
PriMatrix, TEI Biosciences 2008 Decellularized fetal bovine skin ✓ ✓ Wounds Sheet Xelma, Molnlycke 2005 EU ECM protein (amelogenins) in propylene glycol alginate carrier
✓ ✓ ✓ Leg ulcers Gel
Bone INFUSE Bone Graft, Medtronic
2002 Bovine type I collagen sponges soaked in rhBMP‑2 in LT‑CAGE Lumbar Tapered Fusion Device
✓ ✓ ✓ ✓ Spinal fusion Solid
OP‑1, Stryker 2001 Bovine type I collagen with rhBMP‑7 ✓^ ✓^ ✓^ Bone injury Paste PuraMatrix, 3DM Preclinic Synthetic 16‑amino‑acid peptide, forming nanofibres
✓ ✓ Dental bone defects
Gel
Vitoss Scaffold FOAM, Orthovita
2004 Porous foam comprising β‑TCP and bovine type I collagen
✓ ✓ ✓ Bone injury Foam
Bioset IC, Pioneer surgical 2008 Human demineralized bone matrix with bovine bone chips in type I collagen carrier
✓ ✓ ✓ Bone injury Paste
FortrOss, Pioneer Surgical 2008 Nanocrystalline hydroxyapatite and E‑matrix (porcine collagen co‑polymerized with dextran)
✓ ✓ ✓ ✓ Bone injury Paste
Regenafil, Regeneration Technologies/Exatech
2005 Human mineralized bone matrix in porcine gelatin carrier
✓ ✓ ✓ Bone injury Paste
GEM 21S, BioMimetic Therapeutics
2005 β‑TCP particles and recombinant human platelet‑derived growth factor‑BB (PDGF‑BB)
✓ ✓ ✓ Dental bone/ gum defects
Paste
BCT001, Bioceramic Therapeutics
Preclinic Strontium releasing bioactive glasses ✓^ ✓^ Bone defects Granules, paste
460 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials
review article NaTure maTerIalS^ doi: 10.1038/nmat
currently in clinical studies in the US, covering a host of therapeutic
applications including following myocardial infarction^15. Cell
therapy alone, however, may have as yet undetermined, unwanted
and poorly controlled consequences. A recent study in mice has
revealed that stem cells injected into heart muscle post-infarction
went on to mineralize, possibly because the stiffer mechanical envi-
ronment of the scar tissue was not conducive to myogenic differen-
tiation16,17. Control of cell fate is perhaps the most limiting factor in
the translation of embryonic stem-cell therapy. When implanted in
an immunocompromised mouse they will form teratomas — benign
tumours made up of a variety of adult tissues, for example teeth, hair
and sections of gut epithelium^18. On the other hand, a bank of only
150 embryonic stem-cell lines could provide a good tissue match
in more than 80% of the UK population^19 and it is now possible to
‘make’ cells resembling embryonic stem cells (‘induced pluripotent
cells’) by artificially introducing up to four genes into adult cells 20–23.
The latest report achieved reprogramming in human cells with no
permanent genetic modification^24 , making translation to clinical
use a tantalizing goal as tissue-specific, transplantable cells could
be host-derived.
Another important challenge is that of how to replace whole
tissues, which are made of many cell types whose organization is
crucial to function. Cells, of course, have natural powers of self-
organization, and under the correct conditions will spontaneously
form complex structures, such as the sprouting tubular networks
formed by the endothelial cells that line blood vessels. Transmission
and receipt of complex molecular information involved in cell
sorting, boundary formation in tissues and cell movement can be
effected through direct cell–cell interaction, largely through cadher-
ins—a family of transmembrane glycoproteins that regulate cell–cell
adhesion^25. When two cell types expressing two different cadherin
molecules are mixed in suspension they will spontaneously sort
themselves on the basis of their cadherin expression 26 , an outcome
fundamental for tissue development and healing. Many embryo-
logical processes rely on cadherin communication. For instance, the
formation of the central nervous system requires the ‘neural tube’
to bud off from epithelial cells, a process that depends on a change
in the expression of cadherins from E-cadherin to N-cadherin 27.
Simpler, artificial cell adhesions have been engineered in a scheme
involving periodate oxidation of cell surfaces followed by biotin
conjugation. The subsequent addition of avidin triggered the assem-
bly of multicellular aggregates through biotin–avidin interaction 28 ,
which was intended to assist the development of more complex
cellular interactions.
extracellular matrix scaffolds
As well as requiring information from each other, cells derive a vast
wealth of information from their environments, including the mate-
rial that surrounds and separates them within tissues, the extra-
cellular matrix (ECM). A TE material scaffold must take on this
instructive role to some degree in order to maintain cell viability and
control cell behaviour. Clues for how to construct bioactive artifi-
cial scaffolds come from naturally bioactive scaffolds. For example,
implantation of demineralized bone matrix (DBM, bone from which
mineral and cells have been removed, leaving only proteinaceous
material) in muscle induces the formation of bone in the surround-
ing muscle tissue^29. This remarkable observation led to the isola-
tion of bone morphogenic protein (BMP) from DBM, and several
companies currently produce DBM commercially from cadavers for
implantation in bone defects^30 (Table 1). As with DBM, many other
cadaver- or animal-derived decellularized ECM products have an
inherent bioactivity sufficient to induce regeneration and have found
clinical use: for instance, products derived from the small intestinal
submucosa of pigs (an example being Oasis Wound Matrix) are used
routinely in reconstructive surgery, and ECM derived from the peri-
cardium of horses can be used as a reconstructive material in the dura
2009 President Obama lifts ban on federal funding of embryonic stem‑cell research
2008 Implantation of tracheal segment engineered from decellularized tissue
2007 • Creation of induced pluripotent stem cells from adult human skin cells
- Osiris named Biotech Company of the year
- ~170 companies offering TE products or services, sales in excess of US$1.3 billion; >1 million patients treated; aggregate economic activity fivefold higher than in 2002
- Organogenesis breaking even, reinvesting profits; Apligraf sales of US$60 milllion per year
2006 • TE bladder appears in the Lancet
- Launch of Proteus Venture Capital Fund, the first dedicated regenerative medicine fund
2005 Carticel becomes profitable
2003 Organogenesis emerges from bankruptcy
2002 • Integra Dermal Regeneration Template by Integra Life Sciences approved for treatment of severe burns
- FDA approval of Medtronic’s INFUSE Bone Graft
- Organogenesis and ATS, both previously valued at US$1 billion, file for bankruptcy
- TE activity halved since 2000, loss of 800 full‑ time employees, capital value of publicly‑traded TE corporations drops from US$2.5 billion to US$300 million
- Circe bioartificial liver completes Phase III clinical trial with statistically significant benefit for a subset of patients; FDA approval not granted as patient group was not identified in advance
- 42% increase in stem‑cell firms, coining of term ‘Regenerative medicine’
2001 • President Bush restricts federal funding for embryonic stem‑cell research
2000 • Time magazine names Tissue Engineer as ‘Hottest Job’ for the future, 3,000 people pursue TE careers
- US$580 million spent annually on TE R&D, public TE companies valued at US$2.5 billion
1998 • FDA approves Apligraf, first allogenic TE product
- Human embryonic stem cells isolated
1997 • TransCyte becomes first FDA‑approved TE product
- Carticel autologous cartilage implant approved by FDA
1996 The Tissue Engineering Society founded (now Tissue Engineering Regenerative Medicine International Society, TERMIS)
1990 US$3.5 billion invested worldwide in TE, 90% from private finance
Late 1980s Early TE work in Massachusetts, term TE appears in literature
Early 1970s Cells combined with biomaterials in research into artificial skin and biohybrid pancreas
1968 First bone‑marrow transplant
1950s Organ transplantation with identical twins
Figure 1 | Tissue engineering timeline. Tissue engineering gained in profile through the nineties, hitting a peak around the turn of the millennium, but several early commercial ventures failed and large‑scale private financing was withdrawn. Improved business planning and a sounder scientific base have since propelled it towards a new era of success 1–4,22,^.
nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 461
NaTure maTerIalS doi: 10.1038/nmat2441 review article
mater layer of the brain meninges following a craniotomy (Table 1).
In a further development, last year’s transplanted TE airway con-
firmed this approach as being at the forefront of whole-organ tissue
engineering 3,31. The scaffold in this case was a decellularized human
donor trachea that was repopulated with the patient’s own cells
expanded from biopsy. In contrast with traditional transplant surgery,
the decellularization protocol solved the problem of tissue rejection
by removing virtually all traces of human leukocyte antigens (the
proteins that to a large extent determine tissue compatibility), with
the consequence that the patient required no immunosuppressive
drugs. As well as immediately restoring airway patency, the device
facilitated the rapid development of an internal cellular lining and
blood vessel network. Although we focus here on scaffolds designed
and assembled in ‘bottom-up’ mode in the laboratory, it is apparent
that both lab-built scaffolds and decellularized tissues offer distinct
and important benefits for tissue engineering, and equally, that nei-
ther approach represents a universal biomaterials solution.
substituting physical aspects of the extracellular matrix
Typically, biomaterials-engineering approaches focus on a few
mechanisms (chemical or physical) by which ECM influences cells,
and attempt to present these influences effectively for a given tissue.
Regardless of the chemistry that we apply within scaffolds, the con-
struct must usually also provide some level of physical support from
the moment of implantation, to assist tissue function while new
matrix is being deposited32–35. For example, the extreme softness
of the lamina propria of the human vocal fold (elastic modulus
E = 100–1,000 Pa) is essential for proper phonation, and its function
is easily impaired by scarring. This has prompted the development of
soft (E ≈ 500 Pa), highly elastic gels of double-crosslinked hyaluronic
acid microparticles for vocal-fold tissue engineering 36. The par-
ticles are synthesized by crosslinking with divinyl sulphone, then
surface-oxidized and lightly crosslinked together using hyaluronic
acid modified with adipic dihydrazide. The gels have controllable
viscoelasticity, and a reduction in dynamic viscosity with frequency
occurs at a similar rate to that of human vocal-fold mucosa.
In many cases, physical demands on scaffolding materials are
complicated by the anisotropy inherent in most living tissues (con-
sider the parallel arrangement of collagen fibres within tendons
and the concentrically layered sheets of the intervertebral disk).
Nevertheless, engineering solutions need not be costly or comp-
licated: substituting a rotating for a static collector yields orien-
tated electrospun fibres 37 , and crosslinking hydrogels under strain
can result in highly biomimetic anisotropic mechanical properties.
For instance, thermal cycling of poly(vinyl alcohol) leads to the
growth of crystallites that function as physical crosslinks, leading
to gelation. Early in the crosslinking process, these crystallites can
be aligned by applying strain, the degree of which dictates the level
of anisotropy. Thermal cycling is recommenced, with the number
of cycles determining the overall amount of crosslinking and thus
stiffness. By optimizing these two parameters, hydrogels with aniso-
tropic stiffnesses closely resembling those of porcine aorta have
been developed^38. Tissues are also heterogeneously organized into
mechanically distinct zones, for example the superficial, transitional,
Achieving a strong bond between mechanically dissimilar materials
is as much a challenge in tissue engineering as in other branches of
engineering. The morphological specialities and mechanical gra-
dients seen at interfaces between musculoskeletal tissues in vivo
reduce impedance mismatch and minimize stress concentrations
as loads are redistributed, yet even with nature’s elegant solutions,
many chronic musculoskeletal injuries occur at tissue boundaries.
Unsurprisingly, rupture at insertion sites is also the most common
cause of failure of ligament and tendon grafts^131. Although aware-
ness of this problem is growing, most orthopaedic TE devices
do not feature distinct transition zones to improve load transfer
between tissues. This includes most osteochondral plugs — bilami-
nar bone and cartilage TE constructs that have been developed to
improve the assimilation of cartilage into joints. Here, the accu-
mulation of matrix can effect good integration between the two
phases132,133, but few examples contain regions of calcified cartilage
reminiscent of the ‘tidemark’ seen adjacent to subchondral bone
in vivo. An interesting exception is an osteochondral graft consist-
ing of a ‘bone’ component of hydroxyapatite populated with BMP-
transduced fibroblasts (connective tissue cells), and a poly(lactic
acid) sponge seeded with chondrocytes (cartilage cells). Pockets
of mineralized cartilage were seen at the boundary between the
two layers of this scaffold^134. Conversely, trilaminar scaffolds by
design possess a middle layer with a distinct composition^135 and/
or seeded with different cell types. Any combination of cells can be
straightforwardly zoned within hydrogels at the point of fabrica-
tion by the layer-by-layer partial photo-polymerization of cell and
macromolecular precursor suspensions^136. Constructs resembling
ligament insertion sites, wherein bone and ligament are united by
means of fibrocartilage (Fig. B1), have been produced by seeding
fibroblasts, chondrocytes and osteoblasts (bone cells) separately
into the three layers of a preformed scaffold^137. Another strategy
uses just one cell type, namely fibroblasts, to produce scaffolds
with an internal gradient of mineralization. Retrovirus encoding
the bone-specific transcription factor Runx2 was immobilized on
a layer of poly(l-lysine). The thickness of the poly(l-lysine) layer
could be graduated by dip-coating collagen scaffolds, leading in
turn to a tapering of retrovirus concentration, osteoblastic gene
expression, mineralization and stiffness^138. Although the ligament
components in these examples were not optimized for immediate
tensile load bearing, TE ligaments with high tensile toughness (such
as braided polymers^139 ) have been developed^34. It will be interesting
to observe how, in the future, these two strands of ligament tissue
engineering will be united.
Box 1 | Challenges in interface engineering for orthopaedics.
F
BV
Fb
Cc
Ob
a
CF
B
L
T
b
T
L
F
CF
B
Figure B1 | Histology of interface between bone and ligament. a , Schematic; b , photomicrograph of tibial insertion of rabbit anterior cruciate ligament. Adapted with permission from ref. 144. © 1996 Wiley. Ligament (L) insertions occur by means of fibrocartilage, which is divided at the tidemark (T, black line in b ) into non‑calcified (F) and calcified (CF) regions. The calcified fibrocartilage interdigitates with the underlying subchondral bone (B). Fb, fibroblast; Cc, chondrocyte; Ob, osteoblast; BV, blood vessel.
nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 463
NaTure maTerIalS doi: 10.1038/nmat2441 review article
of cell types within the developing embryo. They may also have
profound effects in the control of tissue regeneration. For example,
injured muscle tissue secretes several Wnt proteins, stimulating a
resident population of cells to divide and differentiate to form new
muscle tissue^69 , and as we have already seen, BMPs can induce the
formation of bone ectopically in muscle tissue 29. These observa-
tions have found resonance in stem-cell research with the result that
many growth factors are important components in the differentia-
tion regimes for both adult and embryonic stem cells70–72. In tissue
engineering, the application of growth factors within biomaterials
also represents a powerful tool for controlling cell differentiation
and function. For instance, when murine muscle lacerations were
treated by transplantation of muscle precursor cells within RGD-
coupled alginate gels, recovery was greatly improved by the addition
of hepatocyte growth factor (HGF) and fibroblast growth factor-
(FGF-2)^73. Already, growth factors feature in a handful of commer-
cially available TE products (Table 1), one of which — Medtronic’s
INFUSE — represents the field’s biggest financial success yet 4.
INFUSE is supplied with powdered recombinant human BMP-2,
which is reconstituted in water and added to a collagen sponge
immediately before use.
Controlled-release strategies are frequently adopted to overcome
the short half-life and residence of free growth factors in solution.
For example, microspheres fabricated by double emulsion can
release protein payloads from aqueous pockets within the particles,
and can now be made to nanoscale dimensions using a single sur-
factant 74. Furthermore, in developmental pathways, different factors
become active at different times, and growth factor release profiles
that recapitulate these dynamics are likely to provide more lever-
age over cell behaviour than those that apply growth factors indis-
criminately. Materials schemes based on different degradation rates
or diffusive properties of polymers have been designed with this in
mind75,76^ (Fig. 2).
Although most efforts so far have concentrated on evaluating
the effects of freely diffusible forms of growth factors in solution,
most in fact function at interfaces in vivo, bound to ECM com-
ponents or as part of membrane complexes 77. Although concern
undoubtedly arises over the cellular accessibility and activity of
surface-immobilized proteins, even relatively simple tethering of a
growth factor to a biomaterial matrix can elicit desired biological
responses 78,79^. For example TGF-β1 covalently tethered to PEG not
only retained its ability to stimulate matrix production in vascu-
lar smooth muscle cells, but also did so significantly more than
a comparable concentration of the soluble form of the protein 80.
Fixing growth factors covalently in place carries the added benefit
of preventing internalization of growth-factor–receptor complexes
by cells. More precise, site-specific couplings can be engineered
through the use of recombinant proteins into which additional
amino acids are introduced at the termini, for example Cys-tags
or enzyme substrate sequences that lead to proteolytic release 81,^.
Systems for controlling the kinetics of growth factor release and
presentation have shown potential for aiding blood vessel growth
into scaffolds (Box 2, Fig. 3).
More natural mechanisms of growth factor binding and release
are also being pursued. In vivo, glycosaminoglycans (GAGs), mostly
as components of proteoglycans, have key parts in growth factor
activity, including sequestering them within the matrix, prevent-
ing their degradation and presenting them to cell-surface recep-
tors. GAGs are complex molecules with tissue-specific distribution
and multiple physiological functions, but they share characteristic
linear structures of repeating hexosamine-uronic acid disaccha-
ride units 83. Heparin, and heparan-, chondroitin-, keratin- and
dermatan-sulphate GAGs (HS, CS, KS, DS, respectively) also have
tightly regulated regiospecific sulphation patterns, which determine
their specific interactions with proteins 84,85. These interactions can
be essential for the physiological effects of growth factors. FGF-1,
for example, requires HS binding for dimerization and receptor
activation^86. Heparin has been widely incorporated into TE scaffolds
to offer a slow release mechanism87,88^ (Fig. 2), and CS in commer-
cial products (Table 1) may perform a similar function, including
modulating the activity of cell-derived signalling factors.
simplifying biomolecules for use in biomaterials
Few approved products include recombinant growth factors, but
the enormous success of INFUSE shows the potential commercial
viability of these material/growth factor combinations (Table 1): it
attracted nearly US$700 million of sales in 2007 (ref. 4), an order of
magnitude more than any of its competitor products. Furthermore,
the sophisticated use of growth factors is likely to be important in
advanced TE applications. For example, the patterning of growth
factors within prefabricated scaffolds could aid the generation of
heterogeneous tissues^89. As already discussed for integrin-binding
and protease-digestible proteins, growth-factor-mimicking thera-
peutics where some of the growth factor function is condensed
into relatively short peptide fragments, typically of 30–40 amino
acids90,91, hold promise. Some of these peptides bind their respec-
tive receptors with comparable affinities to recombinant growth fac-
tors, and can trigger signal transduction and lead to appropriate cell
responses. Although the concentrations required to elicit biological
effects are variable, and in some cases exceed those of the native
proteins by orders of magnitude, angiogenesis has been induced
by one FGF-2 mimetic peptide at similar concentrations to recom-
binant FGF-2 (ref. 91). This molecule contains two 15-amino-acid
receptor-binding domains and a 9-amino-acid heparin-binding
Rel
eas
eo
f>^1
growth
factor
d oM
o (^) s e
s (^) f
t a p
l a i
rp
s e
ne
at
oit n
Scaffold surface
Different rates of diffusion
Different rates of polymer breakdown
Loaded polymer and microspheres Loaded polymer coatings R ele as es tra te gie s
Protease (^) Cleavable peptide
Cell- demanded release
GAG sequestered
Enhanced binding
Free in solution
Tethered: random orientation Tethered: specific orientation
Use of spacer such as PEG
Figure 2 | Presentation and release of growth factors from Te scaffolds. Anticlockwise, from top: growth factors within TE scaffolds may be loaded into polymers whose rate of degradation or diffusive properties can be modulated to tailor release rate, and which may be combined into systems releasing multiple factors with distinct kinetics75,76. The exposure of cells to different growth factors with time may therefore imitate developmental pathways and healing responses. An alternative to presenting growth factors in soluble form is to bind them to a surface in either random or specific orientations, with the possible use of a spacer molecule78–80. Non‑covalent associations with matrix components, particularly glycosaminoglycans (GAGs), can effect slow release and in some cases may potentiate binding to membrane receptors87,88. Cell‑demanded release is based on the presence of protease‑sensitive peptide sequences within the growth factor protein81,82.
464 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials
review article NaTure maTerIalS^ doi: 10.1038/nmat
sequence. Furthermore, such mimetics can have demonstrable
effects at the whole-organism level: a 15-amino-acid peptide, based
on the neural cell adhesion molecule (NCAM) binding site for FGF-
receptor-1, imparted to rats a long-lasting improvement in memory
upon intracerebroventricular administration^92. A further advantage
to these mimetics is their relatively high stability relative to native
growth factors, such as BMP-2, which are necessarily used in supra-
physiological doses^93.
Even active compounds bearing no relation to the primary
structures of growth factors can be identified from peptide 94 or
small-molecule 95 libraries. As many cytokines and growth factors
and their receptors are arranged in dimers, bi- or oligovalency can
increase the activity of these compounds. The covalent dimeriza-
tion of a 20-amino-acid erythropoietin (EPO) mimetic peptide
increased its affinity for the EPO receptor 100-fold 94. Mice erythro-
poiesis assays, which measure the incorporation of radioactive
59 Fe into the blood, revealed a similar increase in in vivo potency
of this peptide, although its activity still remained orders of mag-
nitude short of native EPO. More widely, an expanding palette of
small molecules and ions such as retinoic acid, dexamethasone
and thyroid hormones are known to influence differentiation 95,^.
Bioactive glasses such as PerioGlas can be made to release various
ions including calcium and silicon, which can effect upregulation
of genetic pathways relevant to bone differentiation 97,98^. The bioac-
tive glass ‘BCT001’ additionally releases strontium to help combat
osteoporosis (Table 1).
The ability to bind growth factors and thus modulate cellular
functions can be recreated in synthetic GAG mimetics84,85. Synthetic,
sulphated di- and tetra-saccharides in side-chain positions along
a polymer backbone can successfully compete with neural CS for
growth factor binding, despite non-native molecular architectures^84.
One of these glycopolymers, a polymerized CS tetrasaccharide, had
similar potency to neural CS. Interestingly, non-specific chemical
sulphation of hyaluronic acid, a GAG occurring naturally in a non-
sulphated form, induces less extensive structural rearrangements of
adsorbed and covalently bound fibronectin, which translates into
a higher level of cell attachment 99. Alginate can also be chemically
sulphated to yield a substance with binding affinities to growth
factors (for example VEGF, PDGF-BB and HGF) comparable to or
higher than heparin and with the ability to enhance FGF-induced
blood-vessel formation^100. A sulphated and carboxylated dextran
derivative potentiated VEGF binding to its receptors, resulting in
angiogenic effects 101. Even incorporating sulphonated monomers
(sulphopropyl acrylate potassium) into poly(acrylamide) gels
increases the uptake of serum proteins 102. These materials carry the
advantages of scalable chemical synthesis and more closely defined
material properties, and are gaining interest as replacements for
heparin and CS as modulators of growth factor release and activity.
In addition to binding a broad spectrum of proteins, small oligo-
saccharide domains present within larger GAG sequences can
also regulate cellular function through their involvement in spe-
cific structural interactions with their binding partners85,86,103,104.
Tetrasaccharides, for example, represent the minimal CS epitope
necessary to stimulate neuronal growth^105 , and the anticoagulant
activity of heparin has been localized to a pentasaccharide motif that
interacts selectively with antithrombin^106. Thus, in the same way that
short peptide sequences have been used to isolate specific protein
functions, oligosaccharide fragments can emulate the function of
Creating a functional vasculature represents one of the most
fundamental challenges in tissue engineering, and most notable
successes so far have been in thin or avascular structures such
as skin, bladder and cartilage. Surgical approaches whereby
implants are sited alongside a rich external blood supply 2 are
likely to complement materials strategies that attempt to induce
or organize vessel formation, either de novo (vasculogenesis)
or by sprouting of existing vessels (angiogenesis). Endothelial
cells have an inherent ability to form tubular structures, but it
is essential that these are stabilized if regression is not to occur.
Permanent vasculature possesses smooth muscle cells and peri-
cytes as well as an endothelial component, and several studies
have shown the potential of co-culture with various cell types
to improve the longevity of vascular networks 127,140,141^. Pericytes
and endothelial cells in co-culture produce tissue inhibitor of
metalloproteinase (TIMP) -3 and -2, respectively, which stabi-
lizes vessels by arresting the matrix breakdown that is associated
initially with vascular invasion and lumen formation, but also
ultimately with regression 127. Several materials-based approaches
have used vascular endothelial growth factor (VEGF), a potent
angiogenic factor involved in the early stages of blood vessel for-
mation. VEGF has a narrow therapeutic concentration range,
above which capillary formation is prolific but aberrant: the
resulting structures are malformed, leaky and unstable, regress-
ing quickly on withdrawal of VEGF 124. On top of this, VEGF has
a half-life of under 90 minutes in the circulation 142 , hence the
need for it to be delivered through biomaterials that can release
it in low concentrations on a timescale of weeks. An interesting
scheme devised to this end involved the production of a recom-
binant VEGF variant, which was enzymatically incorporated
into fibrin matrices, and released upon matrix breakdown by
cell-produced enzymes. The release rate was accelerated by the
selection of VEGF isoforms with a plasmin-cleavable sequence
near the conjugation site. Interestingly, durable vessel formation
and in vivo vascularization was higher with this VEGF molecular
variant than for native VEGF in a mouse model despite a lower
upregulation of VEGF receptor 2 (the receptor through which
VEGF exerts its effects on endothelial cells) 143.
Other groups have taken a combinatorial approach to growth
factor delivery whereby VEGF is released in tandem or sequentially
with other growth factors involved in the orchestration of angio-
genesis, such as platelet-derived growth factor-BB (PDGF-BB),
FGF-2 and angiopoietin-1 and -2 (refs 140, 142). PDGF-BB is
important in recruiting smooth muscle cells to stabilize nascent
vessel walls; when packaged with VEGF-A inside alginate hydro-
gels, the two growth factors show distinct release kinetics because
of their different affinities for alginate. Used therapeutically, this
material stimulated the formation of mature blood vessels with
associated smooth muscle cells, and improved cardiac function
in a rat model of myocardial infarction^125. Yet another possible
avenue is to use gene transfer such that cells constitutively produce
low levels of VEGF or other desired proteins 124. One target is the
transcription factor hypoxia-inducible factor 1α (HIF-1α), which
holds the key to intracellular detection of hypoxia and subsequent
upregulation of VEGF and other proteins involved in the cascade
of angiogenesis 126. A gene was delivered that encoded a stabilized
form of HIF-1α, which lacked the oxygen-sensitive degradation
domain present in the native form and could therefore initiate
angiogenic events under normoxic conditions. The plasmid was
packaged within designer peptides, one of which incorporated
a factor XIIIa substrate sequence as a means to immobilize the
DNA within a fibrin gel (Fig. 3). Clearly, although vascularization
remains an inadequately resolved challenge, encouraging develop-
ments are being made in this area.
Box 2 | Challenges in vascularization.
466 nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials
review article NaTure maTerIalS^ doi: 10.1038/nmat
High Medium Low
Structural synthetic mimics
Anionic and phosphate groups
Multi-domain peptide
Sulphated and carboxylated dextran
Enzyme-sensitive peptide crosslinkers
Cryptic site peptides
Sulphonated synthetic polymers
Poly(glutamic acid) peptide
Domain III
Integrin binding sequence
Fibronectin
Structure and function of ECM molecules
Protease
Glycosylated synthetic polymers
Proteoglycan aggregate
Chondroitin sulphate, Sulphated groups a GAG
Proteoglycan
Enzymatic cleavage sequences
Amino acids
Bone sialoprotein interaction
R
D G
Cell behaviour Functional synthetic mimics Cell–cell adhesion
Global response: including viability, adhesion and differentiation
Biotin–avidin crosslinking
Polymer microarray
Bone sialoprotein- derived peptide (^) P
C O O
O
O
O
O
Synthetic polymer matrix
Dimerized affinity peptides
Receptor binding Heparin binding
Biotin
Avidin
Cell surface
Protein complex Cell membrane E-cadherin Cytoskeletal actin
Cell junctions
Fibroblast growth factor 1 (FGF-1) bound as dimer
FGF receptor 2 Bound heparan sulphate
Collagen triple helices
Hydroxyapatite crystals
S O
O
O
O (^) S O
O-
Oligosaccharides such as heparin oligomer
- Poly(glutamic acid) sequences
Integrin recognition sequences
RGD IKVAV
YIGSR^ PHSRN PDSGR
EEEEEEEE
^
E
^
E
EEEEEE EE
...GPQGIWQG...
...GPQ G
IWQG...
e
f
Active fragment released
Integrin binding, protease sensitivity
Mineralization mediators
Growth factors
Protein binding
a
b
d
Carbohydrates
c
Proteins
Figure 4 | Synthetic mimics of biological structures. Many characteristics of ECM macromolecules have been reproduced in simpler compounds with biologically inspired structures. a , Certain protein functions, including integrin binding for cell attachment and protease degradability, can be isolated to short amino‑acid sequences. These sequences can be combined with synthetic polymers or incorporated into complex peptides to enable cells to attach to or break down the material, respectively55,56,58–60,63,64. b , Some glycoproteins involved in bone mineralization, such as bone sialoprotein, possess runs of negatively charged amino acids. Peptides that incorporate these sequences, and synthetic polymers with negatively charged chemical groups, can display improved mineral‑nucleating activity120,121,128,129. c , Growth factor action has been demonstrated in peptides possessing receptor‑binding domains; heparin‑ binding sequences may also be included to aid growth factor sequestration90,91. Random peptide libraries have allowed the identification of peptides with affinity to particular receptors, and dimerization of these molecules in some cases can improve receptor binding and physiological response^94. Growth factor action is sometimes potentiated through the actions of glycosaminoglycans (GAGs) such as the binding of heparan sulphate to the FGF‑1/FGF receptor 2 complex. This specific interaction can be achieved using short heparin oligosaccharides^86. d , Furthermore, the protein‑binding function of ECM GAGs such as chondroitin sulphate can be mimicked by grouping sulphated oligosaccharides by polymerization, by sulphating natural carbohydrates such as dextran, or by sulphonating synthetic polymers84,99–102,130. Additionally, certain biological functions can even be supported by chemistries with no relation to biological structures. e , Whereas cell–cell adhesion occurs predominantly through the complex binding of cell surface cadherin proteins, biotin–avidin interactions have been used to artificially aggregate cells^28. f , A range of responses, such as cell adhesion, viability and differentiation, can be differentially affected by particular synthetic substrates. These cell–material interactions can be assayed using high‑throughput screening of cells on polymer microarrays117–119. (Depictions of protein and peptides do not represent structures accurately.)
nature materials | VOL 8 | JUNE 2009 | www.nature.com/naturematerials 467
NaTure maTerIalS doi: 10.1038/nmat2441 review article
a non-adhesive polymer such as poly(hydroxyethyl methacrylate),
which prevents cell migration between spots. Following cell culture,
standard immunohistochemical techniques and microarray scan-
ning can be performed. This provides a way of identifying poly-
mers that support desired responses from specified cell types, for
example those that promote differentiation of human embryonic
stem cells117–119.
Another approach has been to select chemical functionalities
based on their resemblance to characteristic chemical features
of particular ECMs. Earlier we described the bio-inspired use of
sulphonate groups within hydrogels, mimicking their presence in
GAG chains, which increased protein uptake^102. This approach is
well established in bone tissue engineering, where there exist many
examples of materials incorporating anionic chemical moieties that
improve mineral deposition, for instance by NaOH treatment of
scaffold surfaces or by the incorporation of functionalized mono-
mers such as methacrylated amino acids (GlyMA, SerMA, AspMA,
GluMA)120,121. This practice stems from the observation that many
glycoproteins involved in bone mineralization display a high pro-
portion of negatively charged amino acids: for example, bone
sialoprotein (BSP-II) possesses two polyglutamic acid sequences,
and osteopontin (BSP-I) contains a run of 10–12 aspartic acid resi-
dues^122. Phosphate groups also nucleate mineral, and although often
delivered in soluble form in vitro (as β-glycerophosphate) can be
incorporated covalently into scaffolds.
Moreover, these chemical groups may also be instructive to
cells. In a recent study, small defined chemical groups were incor-
porated into PEG gels, and encapsulated human mesenchymal stem
cells differentiated towards cells of those tissues that the functional
groups chemically resembled^54. Thus, those cells cultured in gels
with charged phosphate groups increased expression of RUNX
(CBFA1; an early bone transcription factor), produced a collagen-
rich pericellular matrix, and synthesized osteopontin. Hydrophobic
t‑butyl groups pushed cells towards an adipocytic (fat cell) lineage,
demonstrated by upregulation of the transcription factor PPARγ
and the deposition of intracellular lipid deposits. It is unknown (and
for practical purposes, arguably irrelevant) whether the role of the
chemical modifications was to act directly on the cells, or to cause the
preferential accumulation of particular cell-derived molecules, these
molecules in turn providing behavioural signals to cells. An example
of the latter mechanism in action is the ability of mineral deposits to
sequester osteopontin, which improves cell adhesion and viability
within phosphate-containing PEG gels^123. Whatever the modes of
action, the complexity of biomaterials could be massively reduced if
the essential chemical character of ECM influences could be distilled
into simple chemical functionalities. A summary of the various ways
in which relatively simple molecules can mimic the molecular infor-
mation within the ECM is given schematically in Fig. 4.
Concluding remarks and perspectives
The examples discussed herein demonstrate the importance of
the extracellular environment in determining cell behaviour, and
highlight the need for regenerative materials to provide cells with
biological cues. Much is still unknown about the mechanisms by
which tissues form and heal, yet already insights from developmen-
tal biology and other biological disciplines are actively guiding the
development of intelligent materials that work with nature’s own
mechanisms of repair. These expanding possibilities raise the ques-
tion of how much extrinsic physiochemical information is required
to mobilize endogenous or transplanted cells into producing a
complex tissue, and specifically, what minimum level of materials
complexity is required for a given task. Evidently, a careful appraisal
of the job in hand will reveal that the broader cost and treatment
implications for any biomaterials approach vary with several
interrelated factors including the form of the device, the mode of
delivery, the nature of the cellular component and any regulatory
implications. To elaborate briefly, injectable matrices help to tackle
problems of surgical invasiveness whereas tissue engineering prod-
ucts in sheet form (Table 1) confront problems related to nutrient
supply by limiting diffusional distances. Moreover, materials that
can recruit endogenous cells into scaffolds avoid the expense and
difficulties associated with culture, storage and distribution of cells,
not to mention immune considerations. Encouragingly, however,
it is clear that comparatively simple materials in combination with
an appropriate cellular component can support a high level of tis-
sue organization14,60. The optimization of mechanical and struc-
tural features of scaffolds and their potential to direct aspects of cell
behaviour illustrates that functional sophistication is not necessarily
synonymous with high manufacturing costs.
A large number of commercially viable products for connective
tissues are based on purified ECM components such as collagens
and hyaluronic acid (Table 1), representing a relatively generic ECM
backdrop conducive to the activities of differentiated cells. Imposing a
tissue-specific identity on stem cells in many cases is likely to require
more specific influences, within materials if not during cell culture.
These more advanced biomaterial approaches are just beginning to
trickle through product-development pathways, but the runaway suc-
cess of INFUSE^4 demonstrates the potential impact of schemes that
make use of growth factor activity. It is promising that the outcome
of growth factor administration can be improved enormously with
the use of technically simple slow-release schemes, such as delivery
using polymers. Such considerations may prove critical for the resolu-
tion of complex tissue engineering challenges such as that of vascu-
larization (Box 2). However, the generation of thick or heterogeneous
constructs, and even complex organs, will require further innovations
in biomaterials research. Interest is also growing in the exciting pos-
sibility of using simple chemistries to influence cell behaviour, and in
the development of a range of therapeutics with intrinsic or modulat-
ing growth factor activity, including designer carbohydrates. Several
laboratories in their own ways are actively pursuing simple but effec-
tive solutions to tissue engineering problems, such that the ideal of
structurally simple, yet functionally complex, biomaterials is becom-
ing a plausible possibility for the near future. More widely, there is
evidence in the resurgence of tissue engineering since the gloomy
days of the early millennium that companies offering these products
have become wise to the demands and realities of the marketplace.
The industry has benefited from a heavy dose of reality and, lessons
learned, is ready to prosper.
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acknowledgements We thank R. Langer, K. Godula and A. Ratcliffe for feedback on the manuscript. M.M.S. acknowledges an ERC Individual Investigator Grant for funding, and EPSRC for the funding of E.S.P. and N.D.E.
additional information Supplementary information accompanies this paper on www.nature.com/naturematerials.
