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Exploiting Advanced Hydrogel Technologies to Address Key Challenges in Regenerative Medicine


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Regenerative medicine aims to tackle a panoply of challenges from repairing focal damage to articular cartilage to preventing pathological tissue remodeling after myocardial infarction. Hydrogels are water-swollen networks formed from synthetic or naturally derived polymers and are emerging as important tools to address these challenges. Recent advances in hydrogel chemistries are enabling researchers to create hydrogels that can act as 3D ex vivo tissue models, allowing them to explore fundamental questions in cell biology by replicating tissues' dynamic and nonlinear physical properties. Enabled by cutting edge techniques such as 3D bioprinting, cell-laden hydrogels are also being developed with highly controlled tissue-specific architectures, vasculature, and biological functions that together can direct tissue repair. Moreover, advanced in situ forming and acellular hydrogels are increasingly finding use as delivery vehicles for bioactive compounds and in mediating host cell response. Here, advances in the design and fabrication of hydrogels for regenerative medicine are reviewed. It is also addressed how controlled chemistries are allowing for precise engineering of spatial and time-dependent properties in hydrogels with a look to how these materials will eventually translate to clinical applications.
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1700939 (1 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
Exploiting Advanced Hydrogel Technologies to Address Key
Challenges in Regenerative Medicine
Daniel A. Foyt, Michael D. A. Norman, Tracy T. L. Yu, and Eileen Gentleman*
DOI: 10.1002/adhm.201700939
1. Introduction
Hydrogels are water-swollen polymer networks formed by cross-
linked polymer chains and are ubiquitous in nature. The slime-
producing, eel-like hagfish releases a mucin-based hydrogel
to choke its predators,[1] and humans rely on a hydrogel
network composed of collagen and hyaluronic acid (HA)
to form the vitreous humor of the eye.[2] However, hydrogels
can also be formed from synthetic polymers, opening a myriad
of synthesis strategies to create materials with widely dif-
ferent physical, chemical, and biological properties. Because of
their hydrophilicity and chemical amenability, hydrogels have
long been an exciting and promising tool in biomaterials and
Regenerative medicine aims to tackle a panoply of challenges from repairing
focal damage to articular cartilage to preventing pathological tissue remod-
eling after myocardial infarction. Hydrogels are water-swollen networks
formed from synthetic or naturally derived polymers and are emerging as
important tools to address these challenges. Recent advances in hydrogel
chemistries are enabling researchers to create hydrogels that can act as
3D ex vivo tissue models, allowing them to explore fundamental questions
in cell biology by replicating tissues’ dynamic and nonlinear physical proper-
ties. Enabled by cutting edge techniques such as 3D bioprinting, cell-laden
hydrogels are also being developed with highly controlled tissue-specific
architectures, vasculature, and biological functions that together can direct
tissue repair. Moreover, advanced in situ forming and acellular hydrogels are
increasingly finding use as delivery vehicles for bioactive compounds and in
mediating host cell response. Here, advances in the design and fabrication
of hydrogels for regenerative medicine are reviewed. It is also addressed how
controlled chemistries are allowing for precise engineering of spatial and
time-dependent properties in hydrogels with a look to how these materials
will eventually translate to clinical applications.
D. A. Foyt, M. D. A. Norman, T. T. L. Yu, Dr. E. Gentleman
Centre for Craniofacial and Regenerative Biology
King’s College London
London SE1 9RT, UK
The ORCID identification number(s) for the author(s) of this article
can be found under
Regenerative Medicine
biomedical research. Indeed, their use
in applications including as soft contact
lenses for correcting vision was first pro-
posed by Czech chemists Wichterle and
Lím more than 50 years ago.[3] However,
the rise of the fields of drug delivery, cell
therapies, and tissue engineering (TE) over
the past decades has widened the scope of
their potential applications and hydrogels
are now being developed to do everything
from repair articular cartilage damaged by
osteoarthritis,[4,5] to regenerate heart tissue
after myocardial infarction.[6]
The native extracellular matrix (ECM)
can be thought of as a cross-linked, hydro-
philic polymer network. Therefore, at a
fundamental level, hydrogels, whether
synthetic or naturally derived, are in many
ways akin to the native ECM. Many hydro-
gels can also be cross-linked under mild
conditions, allowing for the encapsulation
of live cells. Because of these features,
hydrogels have been proposed as 3D ex
vivo tissue models. The 3D network of
hydrogels, which enable encapsulated cells to interact with their
environment in all directions, often better replicates the environ-
ment cells experience within tissues compared to 2D cultures,
which can force cells to adopt unnatural polarities. Hydrogels’
chemical amenability also allows them to be formed with widely
different physical properties, including stiffness, and biological
functionalizations mediated by the incorporation of adhesive
and degradable peptide sequences, which can mimic many bio-
logical and physical properties of the native ECM. Hydrogels
are also being explored as therapeutic delivery vehicles. Acel-
lular hydrogels can be designed for site-specific slow release of
drugs or other bioactive molecules, such as growth factors. And
hydrogels with encapsulated cells are being developed for TE
and other regenerative strategies. By modulating their physical
and biological properties, hydrogels can coax encapsulated cells
to form new tissues. They can also retain therapeutic cells at
specific tissue sites, allowing them to mediate repair either indi-
rectly via paracrine signaling, or directly, by differentiating and
producing tissue.
Despite these exciting developments, hydrogels have also
been subject to criticism. Although their hydrophilic proper-
ties are akin to that of many native tissue ECMs, early gen-
erations of hydrogels used for many biomedical applications
lacked important properties of native tissues that are known to
be key in directing cell behavior. Native tissues, for example,
are heterogeneous in structure, respond dynamically to their
© 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA,
Weinheim. This is an open access article under the terms of the Creative
Commons Attribution License, which permits use, distribution and
reproduction in any medium, provided the original work is properly cited.
Adv. Healthcare Mater. 2018, 1700939
1700939 (2 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
surrounding environment,[7] can often self-heal in response to
injury,[8] and their mechanical properties tend to be nonlinear
and often viscoelastic.[9] Conversely, many standard hydrogels,
particularly those formed from synthetic polymers, are static,
only sparingly adaptable to cell-mediated changes, and their
mechanical properties are often linearly elastic. Moreover,
many hydrogels are structurally homogeneous and cannot
mimic the architectural and mechanical complexity of native
tissues at multiple length scales. In general, many hydrogels
also have relatively weak mechanical properties for TE applica-
tions. For example, hydrogels have been widely proposed for
cartilage TE, however, their tensile and compressive properties
often do not match those of the native tissue.[10]
The last decade, however, has witnessed an explosion of
new chemistries, designs, and fabrication methods that have
returned hydrogels to the forefront of cutting-edge bioma-
terials research (Table 1). Researchers are exploring a new
generation of hydrogel-based biomaterials that better act as
tissue models by mimicking the time-dependent and non-
linear properties that govern the behavior of native tissues.
They are also designing materials for regeneration that
interact with cells as never before. Not just delivering cells
locally, but doing so in a controlled way, or designing chem-
istries that recruit cells to the material. 3D printing methods
have also been developed to precisely control tissue architec-
ture and cell localization within tissue-like constructs, and
for the first time, have allowed for the creation of complex
tissue-like structures with vasculature. These next genera-
tion materials require delivery methods to match. Therefore,
researchers have also been exploring exciting means for in
situ cross-linking and injectable delivery. And indeed, hydro-
gels that aim to repair defects in articular cartilage and
restore damaged heart tissue after myocardial infarction are
now in preclinical and clinical trials.
Here, we review some of the state-of-the-art advances in the
design and fabrication of hydrogels for regenerative medicine
(Figure 1). We specifically focus on where exciting new chemi-
stries and manufacturing techniques are allowing researchers
to make materials that better mimic the native tissue. We also
address how controlled chemistries are allowing for more
precise engineering of spatial and time-dependent proper-
ties in these hydrogels with a look to how these materials will
eventually translate to clinical applications.
2. Hydrogels as Ex Vivo Tissue Models
It was not long after George Otto Gey managed to cul-
ture Henrietta Lacks’s cervical cancer cells in a dish[11] that
researchers realized that cells behave differently in the body
than they do on tissue culture plastic.[12] Many cells on 2D sur-
faces adopt unnatural polarities and create large focal adhesion
plaques, behaviors (among a myriad others), which fundamen-
tally differ when cells are within native tissues.[12] Therefore,
to fundamentally understand how cells respond in health and
disease to a variety of stimuli, it is important to develop culture
systems which better mimic cells’ normal 3D environments.
The 3D structure and ECM-like properties of hydrogels make
them one of the best tools biologists have for doing exactly this.
Over the past 15 years, the fields of cell and stem cell biology
have uncovered an increasing role for physical properties of the
ECM in directing cell behaviors. Indeed, materials that control
cell morphology, elastic, and viscoelastic properties of cells’ sub-
strates, and micro- and nanoscale topographies, among other
factors, have been shown to play important roles in directing
stem cell differentiation and driving other fundamental cell
behaviors. As our understanding of how physical properties
of the ECM direct stem cell fate and tissue formation have
grown, we have witnessed a concomitant development of bio-
materials that mimic such properties. However, translating
behaviors on 2D surfaces to 3D tissue-like platforms have
revealed additional complexities. In 3D, not only do substrate
stiffness and topo graphy play important roles, but also cell-
mediated matrix degradability, cell migration, and physical
constraint.[13,14] The field is currently developing new hydro-
gels that allow us to understand the interplay between these
factors, and how they independently and synergistically direct
cell behavior.
2.1. Incorporating Adhesive Motifs
With the notable exception of blood cells, most cells in the
body are anchorage dependent. That is, they must adhere to a
substrate to survive. Anchorage-dependent cells deprived of a
substrate on which to attach will undergo a specialized form
of apoptosis called anoikis.[15] Engaging integrin receptors—
cells’ transmembrane structures that mediate attachment to
extracellular substrates—will not rescue viability[16] because
most cells have to be physically tethered to a surface to sur-
vive. In vitro, this substrate is normally tissue culture plastic,
a plasma-treated polystyrene surface that adsorbs proteins
onto which cells attach.[17] In the body, this substrate is often
the ECM, a network of insoluble protein biopolymers, which
can contain binding sites that mediate interactions with cells.
Hydrogels formed from many ECM-derived biopolymers such
as collagen contain abundant binding sites that mediate inter-
actions between the hydrogel and encapsulated cells. However,
synthetic polymers such as poly(ethylene glycol) (PEG)
Eileen Gentleman is a
Wellcome Trust Research
Career Development Fellow
in the Centre for Craniofacial
and Regenerative Biology
at King’s College London.
Eileen completed her
Ph.D. in Biomedical
Engineering at Tulane
University (USA) in 2005
before moving to Imperial
College London for
postdoctoral work. Her research at King’s focuses on
engineering the 3D cell niche to control stem cell differen-
tiation for tissue engineering.
Adv. Healthcare Mater. 2018, 1700939
1700939 (3 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
and even biologically derived matrices such as alginate do not.
Therefore, adhesive motifs need to be incorporated into hydro-
gels to allow attachment to their surroundings and mediate
cell viability.
Hydrogels intended for cell encapsulation that are lacking in
their own binding motifs are often chemically modified to dis-
play them. Standard strategies include tethering proteins or pep-
tides with integrin-binding sequences to the polymer backbone
of the hydrogel.[18,19] For example, multiarm PEG molecules
are often conjugated to enable some arms to engage in cross-
linking reactions while others present pendant adhesive motifs.
Typical natural ECM components incorporated into hydrogels
include collagen, fibronectin, and laminin. While common
adhesive peptide sequences include fibronectin-derived RGD
and LDV and laminin-derived IKVAV and YIGSR.[20] Cell–ECM
binding interactions can trigger specific signaling cascades
within the cell, including those that control differentiation. This
behavior is mediated through integrins, which only recognize
and interact with specific peptide sequences within the ECM.
For example, although human marrow stromal cell (often
referred to as mesenchymal stem cells, MSC) express a wide
range of integrins, including
2, and
it is their engagement with RGD sequence-containing peptides
that supports their long-term viability. However, as human
MSC (hMSC) differentiate, their integrin expression patterns
change, and the binding of specific motifs drives lineage speci-
fication. For example, the IKVAV binding peptide sequence,
which engages
1 and
1 integrins, promotes osteogenesis
to a greater extent than that induced by YIGSR and RRETAWA
binding sequences. Similarly, both the IKVAV and RRETAWA
sequences, which bind to
1 and
1 integrins,
respectively, are more conducive for adipogenesis than the RGD
binding sequence alone.[19,21] Although beyond the scope of this
Adv. Healthcare Mater. 2018, 1700939
Table 1. Table highlighting important design criteria for biological hydrogels.
Design criteria Design variables Factors to consider
Material Type of material
• Natural:collagen,fibrin,andalginate
• Synthetic:polyacrylamide,polyethyleneglycol(PEG)
• Hybridmaterials:hyaluronicacid(HA),polypeptides
• Biocompatible
• Templatefortissuegrowthin3D
• Allowforvascularization(microporous)
• Appropriatemechanicalproperties
• Biodegradable—reabsorbsatsamerateasregeneration
Cross-linking strategies Physical (noncovalent) cross-linking:
• Ionicinteractions—chargeinteractions
• Peptidebasedself-assembly—Supramolecularstructures,e.g.,
Chemical (covalent) cross-linking:
• Chaingrowthpolymerization—viaredoxorphotoinitiation(UV)
• Clickchemistry—e.g.,Michael-typeaddition
• Enzyme-mediatedcross-linking—e.g.,transglutaminaseandtyrosinase
• Speedofcross-linking
• Complexityofreagents
• Appropriateforcellencapsulation—cytocompatible
• Precisecontrolofmicrostructure(supramolecular)
• Cross-linkingdensity(networkporosity)
• Controlofmechanicalpropertiessuchas:
o Elasticity/viscoelasticity
o Viscosity
Delivery methods • Injectable—insituforminghydrogels
• Hydrogelpatch—Suchastransdermalorepicardial
• Implant—preformedhydrogelscaffold—e.g.,bioprinted
• Invasivenessoftheprocedure
• Targetorgan
• Aimoftreatment—drugdelivery,tissueregeneration,etc.
Biological agents Cell encapsulation:
• Mesenchymalstemcells(MSC)orinducedpluripotentstemcells(iPSC)
• Autologousorallogenic
• Immunogenicity
• Safety(followinggeneticmanipulation)
• Costandavailability
Biological molecules:
• Growthfactors,immunomodulatory,genetherapy
• Dosage
• Choiceofbiologicalagent(s)
• Timing(sustainedreleasevstemporallyspecific)
Figure 1. Summary of state-of-the-art strategies for hydrogel design and fab-
rication and their applications in regenerative medicine. Advancements in
both biology and material science have allowed for the development of com-
plex regenerative strategies. Green: Researchers are designing hydrogels
with various delivery strategies tailored for each biological application. Red:
Our increased understanding of mechanobiology is driving the development
of hydrogels that can aid biologists in understanding these fundamental
processes,andallowresearcherstoexploit themto drivecellresponsefor
regeneration. Yellow: Advanced manufacturing technologies are allowing
for the development of hydrogels with tissue-specific architectures and bio-
logical functionalities. Orange: Acellular hydrogels are being developed to
both deliver relevant biological molecules and direct host tissue response.
1700939 (4 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
review, the reader is referred to excellent summaries of the role
of integrin binding in controlling cell behavior.[22]
Moreover, while the composition of the ECM and ECM-
mimicking peptides incorporated within hydrogels is clearly
important in modulating cell attachment and downstream cell
signaling, the spacing of integrin binding motifs may also be
important. On 2D surfaces, cells are known to be highly sensitive to
interligand spacing. For example, increasing interligand distances
from 58 to 73 nm has been shown to alter cell morphology.[23]
Moreover, smaller interligand spacings appear to induce a bias in
MSC differentiation toward osteogenesis.[24] Most 3D hydrogels
allow for modulation of overall ligand density, and there is evi-
dence that PEG hydrogels modified with dangling RGD-binding
sequences do not support cellular interactions below certain
concentrations;[25] however, the role of precise ligand spacing,
akin to those that have been examined on 2D surfaces, has only
recently been explored in 3D. Pashuck et al. synthesized pep-
tides that utilized a
-sheet motif to self-assemble into nanofiber
hydrogels.[26] As the
-sheet peptide’s backbone spacing was
known, adhesive epitopes (RGDS and PHSRN) were control-
lably spaced at distances between 0.7 and 6.2 nm. Endothelial
cells encapsulated within hydrogels with spacings of 3.2 nm,
which mimicked the spacing found in fibronectin, showed signi-
ficant upregulation of the
5 integrin subunit and adopted
more spread morpho logies when compared to cells encapsu-
lated within hydrogels with 6.2 nm spacing. Ideal networks with
orthogonal geometries are similarly being explored to further
unravel the role, if any, of 3D spacing cues in directing other cell
responses, such as differentiation.[27]
2.2. Controlling Cell Morphology
2D cell shape has long been known to regulate cell behaviors.
Watt et al.[28] determined that cell morphology regulated epi-
dermal cell lineage commitment in the 1980s, and McBeath
et al.[29] reported that hMSC morphology regulated a fate switch
between adipogenesis and osteogenesis through Ras homolog
gene family member A/Rho-associated, coiled-coil containing
protein kinase 1 (RhoA/ROCK) signaling. Many of these find-
ings were made possible by a technique called microcontact
printing, in which 2D patterns of adhesive motifs are stamped
onto an otherwise nonadhesive surface, allowing precise control
of cell size and shape. Using this technique, they and others
established that in MSC, spread cell morphologies promote
cytoskeletal tension, which upregulates osteogenesis, and round
morphologies promote adipogenesis.
Mechanoregulation of MSC fate has recently been asso-
ciated with the downstream hippo pathway effectors Yes-
associated protein (Yap) and transcriptional co-activator with
PDZ-binding motif (Taz).[30] Spread MSC morphologies pro-
mote nuclear localization of Yap/Taz, while restricting cell
spreading, inhibits cytoskeletal tension, excludes Yap/Taz from
cell nuclei, and prompts MSC to adopt adipogenic phenotypes.
However, while Yap/Taz signaling appears to be a key medi-
ator of the conversion of physical to chemical signals, other
candidates such as lamin-A and the retinoic acid pathway are
also thought to be important in the mechanoregulation of a
plethora of cell responses.[31] Stem cells other than MSC
have similarly been shown to be mechanoresponsive. Changes
in actin cytoskeleton organization, for example, have been
shown to induce epidermal stem cell differentiation.[32] Indeed,
when spreading is restricted, keratinocytes differentiate by
regulating serum response factor (SRF). SRF then targets FOS
and JUNB, members of the AP-1 family of transcription factors,
which are known to play important roles in epidermal terminal
differentiation.[32] Although a more detailed discussion is out-
side the scope of this review, putative signaling mechanisms by
which cells sense and respond to their physical environment
have been previously explored.[33]
In addition to showing that cell size influences differentiation,
researchers have also used microcontact printing to demonstrate
that pattern shape, and specifically shape perimeter, also directs
lineage specification. MSC on more rounded patterns that have
smaller perimeters adopt adipogenic phenotypes, while MSC on
patterns with steeper angles that have larger perimeters, such
as star shapes, promote osteogenesis, even when total cell area
is kept constant.[34] For a comprehensive review, the reader is
referred to an excellent review by Vogel and Sheetz.[35]
While control of cell morphology in 2D is relatively straightfor-
ward and has revealed clear effects on MSC lineage specification,
controlling cell morphology in 3D hydrogels is more complex. On
2D surfaces, cells face no external barriers when assuming their
morphologies. However, in both native tissues and within 3D cul-
ture, cell morphology is limited by the presence of the ECM or
encapsulating material. In 3D hydrogels, morphology is generally
controlled through degradation either by incorporating enzyme-
mediated degradation into the hydrogel, or directly 3D patterning
degradative motifs. To change their morphology in this context,
cells often degrade their surroundings by secreting enzymes,
which target specific proteins in their ECM. Within hydrogels
formed from naturally derived polymers such as gelatine, col-
lagen, and fibrin, as well as polysaccharides such as HA,[36] cells
can often degrade their surroundings as they do in native tissues.
However, within synthetic 3D hydrogels, control of cell shape is
often achieved by cross-linking the network with peptides con-
taining sequences that can be cleaved by cell-secreted enzymes.
In nondegradable hydrogels, cells will often adopt round mor-
phologies, even when presented with adhesive motifs. However,
the incorporation of matrix metalloproteinase (MMP)-degradable
peptide sequences allows cells to adopt spread morphologies.[37]
Controlling cell morphology through degradation can have a
profound impact on differentiation. Indeed, hMSC encapsulated
within RGD-modified HA-based hydrogels that are degradable
adopt spread morpho logies, generate cytoskeletal tension and
undergo osteogenic differentiation,[38] while limiting degradation
prompts cells to adopt adipogenic phenotypes.
The alternative strategy to control cell morphologies and
migration within 3D hydrogels is to spatially control the avail-
ability of degradable and adhesive sites. Recently, this has
been achieved with the use of light-sensitive chemistries in
combination with patterned UV irradiation and laser exposure
(Figure 2).[39] Researchers accomplished this by incorporating
nitrobenzyl ether derivatives into hydrogels, which cleave
upon exposure to 365 nm light. The light-sensitive groups
are then linked to adhesion and/or cross-linking components
of the hydrogel. By carefully focusing a laser at specific loca-
tions within the 3D hydrogel, degradation can be induced by
Adv. Healthcare Mater. 2018, 1700939
1700939 (5 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
breaking the cross-links that hold the hydrogel together, or cell
adhesion can be inhibited by cleaving cell-adhesive peptides
from the hydrogel network. Encapsulated cells detect the light-
mediated degradation by altering their migratory behavior; and
sense changes in adhesive ligands by modifying their mor-
phologies. This technology is particularly promising because it
allows for patterning of cell adhesion and degradability in a 3D
spatiotemporal manner with micrometer-scale resolution.
2.3. Modulating Substrate Stiffness
In 2006, Engler et al.[40] reported that 2D substrate stiffness had
a profound effect on the lineage commitment of hMSC in the
absence of chemical induction. Soft substrates that mimicked
the stiffness of brain tissue promoted neurogenesis, stiffer sub-
strates more akin to skeletal muscle prompted cells to become
myogenic, and stiff substrates that matched the mechanical
properties of the developing osteon drove cells to adopt osteo-
genic phenotypes. The implication was that cells physically
“felt” the stiffness of their underlying matrix and used this
information as an important driver of lineage specification.
Although only shown on 2D substrates, this finding opened the
possibility that stiffness itself could be exploited to direct stem
cell differentiation. As a result, there has been tremendous
interest in developing hydrogel systems that allow researchers
to understand the role of 3D stiffness in directing the differen-
tiation of MSC in ex vivo tissue models.
Although the role of 2D stiffness in regulating stem cell fate
has been clear for more than a decade,[40–42] the role of stiffness
in directing MSC lineage specification in 3D is more complex.
Using a nondegradable, RGD-modified, ionically cross-linked
alginate-based hydrogel, Huebsch et al. showed that like on
2D surfaces, MSC responded to the stiffness of their surrounds
in 3D.[43] That is, stiffer hydrogels promoted osteogenesis
and softer adipogenesis. However, unlike in 2D, stiffness did
not impact cell morphology, and all cells maintained round
morphologies independent of hydrogel stiffness. They instead
showed that differentiation depended on the clustering of inte-
grin-binding motifs within the hydrogels, which was dependent
on substrate stiffness.
However, in degradable hydrogels formed from RGD-modified,
methacrylated HA, degradation-mediated cellular traction was
found to direct MSC differentiation, independent of substrate
stiffness.[38] That is, hMSC adopted adipogenic phenotypes
over a wide range of stiffnesses (4.91–91.64 kPa) in nonde-
gradable RGD-modified methacrylated HA hydrogels, but
when hydrogels were modified to allow cell-mediated degrada-
tion, hMSC generated cytoskeletal tension (hydrogel stiffness
3–5 kPa) and differentiated down the osteogenic lineage.
The authors attributed these findings to the relative differ-
ences between ionically and covalently cross-linked hydrogels.
However, the role of stiffness in the differentiation of MSC in
3D hydrogel cultures remains controversial and other factors
could indeed also be playing important roles. For instance,
time-dependent effects in hydrogel systems could confound
results. Cells modify their surrounding matrix by secreting
proteins extracellularly over time when encapsulated in 3D.[44]
Moreover, 3D hydrogel matrices that allow for cell-mediated
degradability will undergo time-dependent changes in local and
perhaps bulk stiffness as the hydrogel degrades, which may
influence cell response.
In summary, utilizing modifiable hydrogels to understand
the contribution of stiffness to lineage specification can be
fraught. While nondegradable systems will likely provide
insight into the differentiation of cells that reside in tissues
in which matrix turnover is slow, degradable systems may
better mimic native tissue niches that experience quicker ECM
turnover. Matrix degradation is known to play central roles in
development, stem cell differentiation, and tissue formation.
For example, in the developing embryo, cells migrate through
3D matrices, undergo cell shape changes concomitant with
differentiation, and remodel their ECM. These behaviors are,
Adv. Healthcare Mater. 2018, 1700939
Figure 2. Microscale photoreversible patterning of proteins within
3D hydrogels. A,B) Fluorescence confocal microscopy images of dual-
protein patterning within hydrogels. Hydrogels were patterned with
covalently immobilized interlocking chains of red protein while sur-
rounding areas were labelled with a green protein. Scale bars = 50 µm.
C) Hydrogel with 3D patterned protein in a staircase pattern. Patterning
was achieved in 3D using focused laser pulses and by varying the
multiphoton laser-scanning conditions, resulting in highly ordered posi-
tioning of proteins. Scale bar = 150 µm. Adapted with permission.[186]
Copyright 2015, Nature Publishing Group.
1700939 (6 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
in many cases, dependent on the degradability of the native
ECM.[45] When ECM degradation in the developing mouse
embryo is repressed by MMP inhibitors, morphogenesis
and development of oil-induced deciduomas are slowed, and
changes take place in precursor stromal cell differentiation and
expansion.[46] Cell-mediated matrix degradation is also impor-
tant in wound healing. When matrix degradation was inhibited
by the broad-spectrum MMP inhibitor BB-94 in full-thickness
skin excisional wounds in rats, myofibroblast formation, stromal
cell proliferation, blood vessel formation, and epithelial wound
coverage, key components of healthy skin wound healing, were
all delayed.[47] In short, many tissues are highly dynamic and
degradation and ECM remodeling are key tools they utilize to
maintain homeostasis. Therefore, incorporation of this feature
into ex vivo models may be important to fully understand and
exploit tissue regeneration.
2.4. Integrating Time-Dependent and Self-Healing Properties
Many tissues in the body, particularly soft tissues, exhibit
viscoelastic mechanical properties. That is, they behave simul-
taneously as both an elastic material and a viscous fluid.
Viscoelastic materials will continue to deform, or creep, if left
under an applied load, or undergo stress relaxation, whereby
they exert less stress over time, when placed under a constant
deformation.[42] Such properties are well described in muscu-
loskeletal tissues such as ligament and tendon where these
behaviors play important roles in normal joint function.[48]
Until recently, most synthetic hydrogels examined for
directing cell behavior in response to mechanical properties
such as stiffness were linearly elastic. That is, they possessed
a linear relationship between stress and strain, and they
returned to their original shape upon unloading, without loss
of energy. However, it is now becoming apparent that cells,
both on 2D surfaces and when encapsulated within 3D hydro-
gels, respond not only to the elastic but also viscoelastic prop-
erties of their underlying or surrounding matrix.[49,50] Chaud-
huri et al. recently showed that a human osteosarcoma cell
line (U2OS) responded differently when cultured on the sur-
face of viscoelastic ionically cross-linking alginate hydrogels
compared to when they were on purely elastic hydrogels of the
same initial stiffness that were formed from covalently cross-
linking the alginate.[49] Indeed, counter to the intuitive expec-
tation that cells would integrate the modulus of a relaxing
substrate over time and behave as if they were on a softer
substrate, U2OS cells actually spread more on substrates that
underwent stress relaxation. The authors suggested that cells’
ability to respond to viscoelasticity is likely a fundamental bio-
logical property.
Similarly intriguing observations were reported when cells
were encapsulated within 3D viscoelastic hydrogels.[51] In RGD-
modified alginate hydrogels, the proliferation and morphology
of encapsulated 3T3 fibroblasts were highly dependent on the
time scale of stress relaxation. That is, in matrices that under-
went fast relaxation, cells tended to spread more and prolif-
erate, while slowly relaxing materials prompted cells to adopt
round morphologies and inhibited proliferation. Similarly,
when murine MSC were encapsulated within fast-relaxing
hydrogels, they differentiated to osteoblasts that formed a
mineralized collagen type I-rich matrix, but in slowly relaxing
hydrogels, MSC adopted adipogenic phenotypes. The authors
attributed these observations to the ability of encapsulated
cells to cluster RGD ligands in the faster-relaxing hydrogels.
They also argued that in fast-relaxing hydrogels, MSC were
more able to mechanically remodel their surrounding matrix.
Interestingly, the authors correlated increased Yap nuclear
localization with faster-relaxing hydrogels, independent of
hydrogel stiffness. However, unlike in 2D systems where
nuclear localization has been correlated with osteogenesis,[30]
substrate viscoelasticity-mediated nuclear translocation of Yap,
did not itself direct lineage specification.
In addition to viscoelastic hydrogels, much interest has
also been focused on creating viscoelastic hydrogels that
are also self-healing, and so can mimic repair processes that
take place in native tissues. These hydrogels are typically
composed of macromolecules that are noncovalently bonded
together via molecular recognition motifs such as hydrophobic
interactions, hydrogen bonding, metal chela-
tion, or van der Waals interactions.[52] Various mechanisms
have been exploited to create self-healing hydrogels, but the
basic premise is that interactions between individual molecules
are locally dynamic, but the bulk hydrogel is globally stable.
The noncovalent and dynamic nature of these interactions not
only make these hydrogels self-healing, but also easily inject-
able and shear thinning. Shear thinning materials can protect
encapsulated cells against fluid shear forces generated during
injection by “un-cross-linking” under shear and re-establishing
cross-links when shear is removed.[53] A key advantage of shear
thinning hydrogels formed through physical interactions is that
they can be formed ex vivo, prior to injection. In this context,
the influence of the surrounding tissues on hydrogel gelation
is negligible, in contrast to hydrogel systems with liquid pre-
cursors, whose gelation may be affected by various constituents
of the native milieu. Self-healing hydrogels are yet to be fully
exploited to understand cell-material interactions relevant to
TE and regenerative medicine. For a comprehensive review of
self-healing biomaterials, the reader is referred to an excellent
review by Webber et al.[52]
3. Hydrogels with Tissue-Specific Mimicry
and Functionality
In addition to developing materials that can be used as ex
vivo tissue models, there is also tremendous interest in using
hydrogels directly in TE. TE aims to treat a myriad of diseases
by replacing lost/damaged tissue with living constructs created
in the laboratory. The applications of TE range from restoring
tissue lost to myocardial infarction, to filling bone defects with
cell-laden scaffolds that respond to load and remodel over
time. However, the repair/regeneration of complex tissues and
organs requires approaches that are specific to each tissue.
When designing a hydrogel that can mediate tissue repair, sev-
eral parameters are key to consider, including the tissue archi-
tecture, mechanical and biological cues, and cell type (Figure 3).
Indeed, advances in all three areas are likely key in ensuring
effective regeneration.
Adv. Healthcare Mater. 2018, 1700939
1700939 (7 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
3.1. Incorporating Tissue-Specific Architectures
into Hydrogel Design
Whereas many early generations of hydrogels were homoge-
neous, isotropic structures that bore little resemblance to com-
plex native tissues, the design of tissue-specific architects for TE
has taken a leap over the past decade with the advent and growth
of 3D bioprinting. Bioprinting is a computer-assisted technology,
which can assemble tissues-like constructs through precise
spatial localization of biological materials in 3D. Bioprinting
technologies utilize “bioinks” often composed of different hydro-
gels, with or without encapsulated cells.[54]
Within the field of bioprinting several different methods are
available, many of which are amenable to hydrogel technolo-
gies. The most common include microextrusion, inkjet, and
light-induced methods, which include stereolithography and
laser-assisted bioprinting (LAB) (Figure 4). For a detailed over-
view, the reader is referred to excellent reviews in this area.[54–56]
Of these technologies, inkjet bioprinting was the first to be
developed.[55] Advantages of this system include its high
printing speed and low cost; however, thermal and mechan-
ical stresses can damage encapsulated cells.[54,57] Another
limitation of inkjet systems is their relatively poor capacity
to handle materials of high viscosities, which can limit their
ability to print hydrogels with high cell densities.[54,55] Inkjet
printing has been shown to be effective in the regeneration of
both skin[58] and cartilage.[59,60] For example, Markstedt et al.
used a nanocellulose bioink to produce precise anatomically
shaped cartilage structures such as the human ear and sheep
Adv. Healthcare Mater. 2018, 1700939
Figure 3. Tissue-specific hydrogel design considerations. When developing tissue-specific hydrogels, a number of factors should be considered.
suchas thosethatallow forthecreationofpatient-specific stemcells,for exampleiPSC,may allowforadditionalopportunities.Hydrogelmaterial
choice is not always straightforward and may be dictated by a range of factors, such as amenability to bioprinting, the need to form a tissue interface
or the necessity of tissue-specific functionality.
1700939 (8 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
Microextrusion printing works on a similar basis to inkjet,
however, it allows materials of higher viscosities to be depos-
ited. This is important as it allows for the deposition of high-
density cell solutions, which are key in replicating many
highly cellular native tissues.[54] This strategy has been used to
print aortic valve conduits,[61] vascular grafts,[62] and cartilage
constructs,[63] among other tissues. LAB systems are lauded
for their high resolution as they can produce resolutions on
the microscale (10 µm), which is important for mimicking
many structures of native tissues at or near the scale of the
single cell.[64] An advantage is that these systems are nozzle
free, which precludes issues with clogging that have been
problematic with the other bioprinting technologies.[54] Ste-
reolithography, like LAB, uses light to print and works by
selectively solidifying bioinks in a layer-by-layer process.[55]
While lacking the high resolution of LAB methodologies
(25 µm), this layer-by-layer process can often increase
printing speeds.[56]
In situ printing, in which materials combined with cells are
printed directly into a tissue defect (as opposed to requiring
later surgical implantation) is also emerging as a method to
create materials in situ with precise architectures to drive
tissue-specific repair. In comparison to classic bioprinting strat-
egies, in situ bioprinting has several advantages, most notably
of which is that it potentially allows for the fast, direct delivery
of cells. This style of printing is often envisioned for future
clinical applications, in which fully automated robotic printers
controlled by surgeons directly print architecturally controlled,
cell-laden constructs at the site of injury. In situ printing of
amniotic fluid-derived stem cells and MSC has proven effective
in mediating wound closure in a severe skin wound model.[58]
Work by Keriquel et al. have similarly demonstrated a LAB-
based bioprinting system in which in situ printing of MSC
collagen/hydroxyapatite bioink was able to stimulate bone
regeneration in a calvarial defect model in a mouse.[65]
The authors were able to print precise scaffold patterns into
the site of injury, and cells within the printed scaffold main-
tained good cell viability.[65]
While bioprinting allows for specific design of tissue
constructs, one drawback of the technology has been its
scalability. That is, there has been a limit thus far in printing
large, structurally sound, biological constructs that are needed
to repair/replace many tissues.[66] However, one pioneering
study demonstrated an integrated tissue-organ printer in which
human-scale tissue constructs could be printed (Figure 5).[66]
This bioprinting system simultaneously dispensed composite
cell-laden hydrogels consisting of fibrinogen, HA, gelatin and
glycerol, alongside a synthetic biodegradable polymer and an
outer sacrificial acellular hydrogel mold. This mold provided
the tissue constructs with enough rigidity to ensure the con-
struct retained its shape, but could then be easily removed.
A lattice of microchannels was also incorporated into the
design to allow for nutrient and oxygen diffusion within the
printed construct.[66] The authors also implemented CT and
MRI imaging to help design the scaffold. Raw imaging data
was processed using computer-aided manufacturing tools and
mathematical modeling to produce 3D rendered models.[66]
This allowed them to capture architectural intricacies, ena-
bling them to print calvarial bone, ear, cartilage, and skeletal
Adv. Healthcare Mater. 2018, 1700939
Figure 4. Examplesof differentbioprinting methods.A)Inkjetbioprintersdeposit smalldroplets ofhydrogelandcellsto buildtissue layer-by-layer.
heat a donor layer (green), which forms a bubble propelling the bioink onto the substrate. D) Stereolithography bioprinters use UV or visible light to
selectively cross-link bioinks layer by layer to build a 3D construct.
1700939 (9 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
Beyond bioprinting, other methods have also shown promise
in integrating tissue-specific architecture into hydrogels.
One particularly promising area is that of self-healing/
self-integrating hydrogels. These hydrogels are based on
dynamic chemistries and thus are potentially exciting for a
variety of TE strategies. Hsieh et al. investigated self-healing
hydrogels to induce blood capillary formation. An inject-
able composite hydrogel was synthesized from chitosan
and fibrin, which through Schiff-base linkages, lent the
hydrogel self-healing properties.[67] When seeded with vas-
cular endothelial cells, the hydrogel allowed for the forma-
tion of capillary-like structures. These self-healing hydro-
gels have several advantages, the most obvious being that
they can repair themselves following damage. In this case,
the composite material was also stronger and more stable
than a fibrin only material.[67] When targeting tissue inter-
faces, self-healing hydrogels can be particularly advanta-
geous as separate hydrogel components specific to each
tissue type can be placed adjacent to one another, which will
then then self-heal to form a complete construct. Indeed,
Hou et al. developed an injectable self-healing-integrating
hydrogel for regenerating the bone–cartilage interface that
allowed for the relatively simple spatial localization of MSC
and chondrocytes[68] (Figure 6).
Approaches that combine 3D printing with self-healing
hydrogels are also being developed. Highley et al., for example,
developed a self-healing shear thinning HA-based bioink modi-
fied with adamantane and
-cyclodextrin that could then be
printed into a “support” HA hydrogel, creating a “guest–host
system. The printed HA-based bioink could then be covalently
cross-linked to form a stable structure. Due to the self-healing
nature of this system, direct assembly (writing) of precise struc-
tures in 3D could be achieved that otherwise would not be pos-
sible with standard “layer by layer” 3D printing strategies.[69]
Adv. Healthcare Mater. 2018, 1700939
Figure 5. Human-scalebioprinting. To printconstructsofsufficient sizeforeventualtranslation intohumans,Kanget al.developedan integrated
tissue-organprinter(ITOP) system in which large, tissue-specific constructs could be printed. a)TheITOP system comprised three major units:
chamber with a humidifier and temperature regulator. b) Using this bioprinter, 3D scaffold architectures could be printed with both multiple cell
types and a PCL polymer to ensure structural rigidity. c) 3D CAD models were generated from medical image data; and using CAD/CAM processing,
theprintercouldbeusedtocreatecomplex3Dstructures,includinga3Dhuman-sizedear.Reproducedwithpermission.[66] Copyright 2016, Nature
Publishing Group.
1700939 (10 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
3.2. Integrating Biomimicry: the Importance
of Mechanical and Biological Cues
In addition to mimicking the architecture of the native tissue,
hydrogels used for tissue regeneration may also need to mimic
other components of the tissue environment, such as mechan-
ical and biological cues. In nature, soluble cues are transported
to cells either via the vasculature or reach cells via diffusion
after being secreted by their neighbors. However, cells also
receive signals by interacting with their ECM. Such interactions
are key in providing tissue-specific signals to cells, instructing
them to proliferate, differentiate, and secrete/degrade ECM
proteins, among other cell behaviors.[70] Indeed, within the
neural stem cell niche of the subventricular zone of the brain,
proteoglycans such as heparan sulfate bind many factors key
for adult neurogenesis, and have been implicated in the con-
trol of proliferation and migration.[71] Similarly, perlecan within
the kidney glomerular basement membrane has been shown to
play important roles in filtration.[72]
ECM-based signals can be similarly incorporated into hydro-
gels, most simply by forming hydrogels from tissue-specific
ECM. This is often achieved through a two-step process in
which ECM material is solubilized into protein monomeric
components, which are then formed into a hydrogel via temper-
ature or pH-controlled neutralization.[73] For many complex tis-
sues, including the heart and brain, ECM-based hydrogels can
provide important biological cues.[73,74] For example, in a stroke-
damaged rat brain model, an ECM-derived hydrogel provided
support for human neural stem cells and allowed for generation
of de novo tissue.[75] Pati et al. similarly bioprinted decellularized
ECM hydrogel bioinks from heart, cartilage, and adipose
tissues.[76] When MSC or adipose-derived stem cells were
encapsulated within these decellularized ECM hydrogels, they
enhanced differentiation down either the chondrogenic or adi-
pogenic lineages, in comparison to culture within collagen
hydrogels.[76] The authors also reported enhanced myoblast mat-
uration when cells were encapsulated in heart-derived ECM.[76]
In short, this bioprinting strategy allowed for both the precise
architectural control of the scaffold while also incorporating
appropriate biological signals. Another area in which ECM hydro-
gels have been successful is in the formation of human breast
tissue.[77] 3D hydrogel scaffolds were formed from breast tissue-
specific ECM including both protein (collagen and fibronectin)
and carbohydrate components (hyaluronan).[77] Primary human
breast epithelial cells encapsulated within the hydrogels were
then shown to rapidly reorganize and form mature mammary
tissue-like structures.[77] Importantly, tissue organization was
observed in the absence of stromal cells, which are often consid-
ered key for ECM production and morphogenesis in vivo.
One concern with ECM-derived bioinks is that they lack
structural integrity. One way this has been addressed is by
printing multiple materials simultaneously: i.e., coprinting a
strong synthetic hydrogel with a weaker ECM-derived hydrogel.
Indeed, with the development of multidispensing bioprinting
systems, strategies such as this may be key in designing
complex scaffolds with both the correct biological cues
and structural integrity. One way this could be imagined is
through further development of the bioprinting system devel-
oped by Kang et al. in which biological cues, such as ECM, are
incorporated into the bioink, which could then be coprinted
Adv. Healthcare Mater. 2018, 1700939
Figure 6. Injectableself-integratinghydrogelsforosteochondralrepair.Schematicillustrationofaself-healinghydrogelthatcouldpotentiallybeused
torepairanosteochondral defect. Twotissue-specific hydrogel formulations made fromthesamebulkmaterialare injected into the damaged
chemistries would then foster the formation of a seamless transition between the chondrogenic and osteogenic hydrogel formulations. The chondro-
genichydrogelcould, for example,containosteochondralprogenitor cells orchondrocytesandchondrogenic factors suchastransforminggrowth
), BMP and/or insulin-like growth factor (IGF) to promote cartilage-like ECM production. Whereas the osteogenic hydrogel might include
1700939 (11 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
with a synthetic polymer and an outer sacrificial hydrogel mold
to ensure structural rigidity (Figure 7).[66,76]
In addition to ECM, other biological factors can be incor-
porated into hydrogels to mimic biological cues of the native
tissue, including growth factors and cytokines. Spatial control
of these signals may be key to replicate tissue-specific pat-
terning during differentiation and recapitulate the heteroge-
neous microenvironments of the target tissue.[78] Work by Lee
and Park addressed this issue by using a thermoresponsive
casting process to create precise PEG hydrogel structures with
spatial distributions of different biological cues. They showed
that by spatially localizing signaling molecules, MSC differ-
entiation could be driven toward osteogenic, chondrogenic,
or adipogenic lineages.[78] Moreover, by utilizing drug-releasing
PLGA microparticles, they were also able to demonstrate sus-
tained release of biological cues.
In addition to biological cues, mechanical cues may also be
important to direct tissue formation. As discussed above, cells
respond to mechanical cues in their environment, such as
stiffness. Lutolf and co-workers showed that within intestinal
organoid cultures, a mechanically dynamic PEG-based hydrogel
was key in regulating appropriate cell responses. An initially
stiff matrix allowed for intestinal stem cell expansion, but when
switched to soft, stimulated intestinal cell differentiation.[13]
Likewise, modulation of hydrogel stiffness (7–33 kPa) has been
shown to promote neocartilage formation from encapsulated
cocultures of adipose-derived stem cells and neonatal chondro-
cytes. In comparison to soft hydrogels, stiffer hydrogels accel-
erated the deposition of sulphated glycosaminoglycans, a key
component of cartilage ECM.[79]
Hydrogels can also be modulated to mimic other important
biological stimuli. Electrical signaling, for example, is central to
normal heart tissue function and hydrogels can be designed
to support this. Work by Yang et al. created a homogeneous
electronically conductive hydrogel by controlling levels of
conductive (Poly(thiophene-3-acetic acid) and flexible (meth-
acrylated aminated gelatin) polymers. By modulating the
ratio of the two components, they could precisely control the
mechanical and conductive properties of
the construct, achieving a Young’s modulus
and electrical conductivity that were both
similar to that of the native heart.[80] This
advance may be crucial, as both mechanical
and electrical mimicry are likely important in
restoring synchronous contractile activity.[80]
Moreover, when brown adipose-derived stem
cells were seeded on the hydrogels, electrical
stimulation appeared to drive cardiomyo-
genic differentiation.[80] Work by Nunes et al.
similarly demonstrated that electrical stimu-
lation of 3D induced pluripotent stem cell
(iPSC) cultures within a collagen gel resulted
in the formation of 3D aligned cardiac tissue
(“biowires”) with ultrastructural organiza-
tion and electrophysiological properties that
better matched those of native tissue cells
compared to controls.[81]
3.3. Incorporating Multiple Cell Types into
Complex Constructs
Besides utilizing a scaffold with appropriate
mechanical and biological cues, the choice of
cells in regenerative applications is essential.
In addition to MSC, iPSC are also a prom-
ising option in hydrogel-based regenerative
strategies as iPSC technology allows for the
development of patient-specific cells.[82] First
described by Takahashi and Yamanaka[83]
in 2006, efforts are being made to develop
robust, chemically driven protocols to differ-
entiate iPSC into nearly every cell type in the
body.[82] As iPSC derivation and differentia-
tion technologies have developed, so too has
their use in hydrogel-based TE strategies. For
example, in a composite bioprinting strategy,
a nanofibrillated cellulous alginate bioink
Adv. Healthcare Mater. 2018, 1700939
Figure 7. Potential strategy for in situ bioprinting. Schematic illustration showing a potential
strategy for in situ bioprinting of tissue-specific hydrogels with different cell types to treat a
nonunion in the radius. Multiple hydrogels could potentially be employed to reconstruct com-
plex,multicellular tissueslikebone. A)Anacellularhydrogel isprintedto actasa structural
support while the bone heals. B) A cell-laden hydrogel with tissue-specific architecture and
biological factors is printed to promote ossification. C) Other cell-laden, biologically targeted
in tissue remodeling.
1700939 (12 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
could successfully drive chondrocyte-derived iPSC to form car-
tilage-like tissue.[84]
Cartilage TE only requires incorporating chondrocytes into a
scaffold; however, for more complex tissues, multiple cell types
are often necessary. Inkjet printing has been used to incorporate
multiple cells types into hydrogels with precise spatial control.[85]
This was achieved in an alginate–collagen composite hydrogel
system by loading different cell types (human amniotic fluid-
derived stem cells, smooth muscle cells, and endothelial cells)
into separate bioink cartilages. Using a drop-by-drop method,
cells could then be directly printed to form a complex hetero-
geneous construct.[85] When implanted in a mouse model, con-
structs were shown to have adequate vascularization, which had
previously been a hurdle for generating functional bone tissue.[85]
Another strategy to address the issues of vascularization is
coprinting. This strategy has been pursued whereby different
cell types, vasculature, and ECM bioinks have been coprinted to
form complex heterogeneous structures[86] and vascularized tis-
sues as thick as 1 cm.[87] In this technique, silicon ink is initially
printed to create a customized perfusion chip. Cell-laden inks are
then printed alongside a temporary fugitive ink, encapsulated in
a castable ECM. Following casting, the fugitive ink is removed
leaving behind a vascular network of interconnecting chan-
nels.[87] Tissue constructs printed in this manner demonstrated
greater perfusion and survived longer culture times (6 weeks vs
14 d) than previous efforts.[86,87]
3.4. Incorporating Cell-Responsive and Dynamic Properties
Mechanical properties of tissues and biomaterials clearly
play a role in directing cellular responses such as differen-
tiation. However, in native tissues, many properties, including
mechanical properties are dynamic. For example, during chick
development, the tissue that arises from the mesoderm that
is destined to become the adult heart stiffens from an elastic
modulus of 0.9 to 8.2 kPa between 36 and 408 h postfertiliza-
tion.[88] Dynamic properties such as this can be incorporated
into hydrogels using a number of strategies. Young and Engler,
for example, matched the stiffening of their thiol-modified HA
hydrogels to that of the dynamic properties of the chick heart.
They found that dynamic hydrogels upregulated the expression
of mature cardiac markers when compared to cells grown on
soft, but static polyacrylamide hydrogels.[88]
Other methods to incorporate dynamic properties into hydro-
gels include exploiting supramolecular chemistry and self-
assembly strategies. For example, by incorporating specific pep-
tides, hydrogels can be designed to respond to physical stimuli,
such as light and temperature, as well as chemical and biolog-
ical stimuli, such as pH or enzyme cleavage.[52,89] Promising
work by Stupp and co-workers has shown that supramolecular
nanofibers are able to effectively delivery the growth factor
BMP-2 for bone regeneration.[90] Peptide sequences can also
be used to mimic complex biological factors such as vascular
endothelial growth factor (VEGF).[91] Moreover, work by Silva et
al. showed that high density presentation of a laminin-derived
epitope on supramolecular nanofibers could selectively differen-
tiate neural progenitors into neurons.[92] Indeed, incorporation
of such supramolecular materials into hydrogels for controlled
delivery may enable the development of dynamic systems in
which the materials themselves are capable of adapting to their
surrounding environment.
In addition to supramolecular materials, researchers are also
exploring dynamic materials based on 4D biomimetic printing.
For example, researchers have used cellulose fibrils to print
anisotropic swelling behavior into hydrogels that change their
shape as a function of time following immersion in water.[93]
The authors used this system to create nature-inspired 3D
shapes, such as that of the Dendrobium helix, a stunning orchid
native to the lowlands of New Guinea. However, such tech-
nologies could also potentially be applied to bioprint tissues
with mesoscale structures that form upon immersion in bio-
logical fluids.[93] While 4D biomimetic systems for regenerative
applications have not yet been reported, we envision a future
that would utilize such technologies for clinical applications.
For example, there has been much effort devoted to creating
TE heart valves to replace stenotic or damaged values resulting
from disease or congenital conditions. However, such advanced
therapies would likely still require traumatic and dangerous
open-heart procedures to surgically implant them. A 4D bio-
printed heart valve, on the other hand, may allow for the future
development of minimally invasive heart valve repair whereby
a new valve could be inserted via a catheter and formed in situ
in the heart. Similarly, maxillofacial surgeries to repair orbital
floor fractures, common in patients in motorbike accidents,
are often necessary to prevent enophthalmos, but can suffer
from complications sometimes associated with the surgical
placement of alloplastic implants.[94] Here, a future in which
4D-bioprinted constructs, fabricated to match a patient’s own
facial structures based on high-resolution imaging, and placed
in a minimally invasive procedure prior to assuming their full
shape, are feasible.
4. Acellular Hydrogel Approaches
As cell-based therapies such as TE offer numerous possibilities
for tissue regeneration, there is good reason for excitement sur-
rounding their development. However, cell-based therapies face
significant challenges in terms of cost, regulatory hurdles, and
scalability.[95,96] Cell free therapies, on the other hand, are often not
as limited in their path to clinical translation because they tend
to be less complex. Traditionally, acellular biomaterials have been
used simply as fillers and for structural support, however, a new
generation of acellular hydrogels are being designed to interact
with endogenous factors, including local tissue and cells, to aid
with healing and promote tissue regeneration (Figure 8).[95,97]
4.1. Hydrogels as Controlled Delivery
Vehicles for Bioactive Molecules
The fundamental approach of many acellular biomaterial
strategies is to trigger tissue regeneration in situ by uti-
lizing the body’s own regenerative capabilities. This is often
achieved by mobilizing host tissue progenitors to the site of
injury or prompting nearby cells to mediate repair. The con-
cept of biomaterial-driven tissue regeneration is not new and
Adv. Healthcare Mater. 2018, 1700939
1700939 (13 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
many biomaterials have been designed to deliver biological
molecules such as growth factors and cytokines to stimulate
matrix production and cell growth. There are many excellent
reviews available on this subject; therefore, we will only com-
ment on a limited number of strategies here.[98]
One of the most well-known drug delivery platforms based on
a biomaterial is Medtronic’s INFUSE system, which comprises
a collagen type I biomaterial that releases recombinant human
BMP2, and is placed surgically to treat long bone fractures and
in spinal fusion procedures.[99] Although the INFUSE system has
been highly criticized for causing side effects such as inflamma-
tion when used in off-label procedures,[100] the fundamental idea
of using a biomaterial to deliver an active biological molecule is
widespread. Indeed, many issues associated with the off-label use
of INFUSE arose from the high dosage of growth factors rather
than delivery strategy itself.[100] Cardiac stents have been similarly
designed to elute drugs that can prevent re-stenosis of coronary
arteries,[101] and wound dressings which release antibiotics and
growth factors to aid healing are widely described.[102]
Hydrogels provide an attractive system for the delivery of
biological molecules, as in addition to good biocompatibility
and amenability for minimally invasive delivery, they are also
highly physically and chemically modifi-
able,[98] which can allow for precise control
of release of biological molecules over min-
utes, hours, days, and perhaps even years.
Most commonly, this release rate is mediated
via diffusion, which can be easily controlled
by varying either the hydrogel’s network size
or degradation rate.[98] One strategy to con-
trol hydrogel degradation involves a similar
strategy to that used to create degradable
ex vivo tissue models: cross-linking hydro-
gels with peptides that are susceptible to
cleavage by MMPs.[103] For example, MMPs
are important in heart tissue homeostasis,
but in specific cardiovascular diseases exces-
sive MMP production can lead to inflam-
mation and tissue damage. Purcell et al.
exploited this tissue response with a polysac-
charide-based hydrogel that electrostatically
sequestered a recombinant MMP inhibitor.
The hydrogel was then cross-linked with
MMP-cleavable peptides such that cleavage
of the peptides mediated the release of the
MMP inhibitor.[104] When placed in a porcine
model of myocardial infarction, the hydrogel
was effective in reducing adverse left ventric-
ular remodeling.[104]
As our understanding of the biological
mechanisms of tissue regeneration con-
tinue to grow, so too has the development of
hydrogels that aim to exploit them. One way
this has been pioneered is through under-
standing the paracrine mechanisms by which
stem cells can stimulate tissue regenera-
tion. For example, MSC appears to mediate
regeneration via release of paracrine fac-
tors rather than de novo regeneration.[105]
Hydrogels provide an exciting opportunity in this context as
they can allow for the controlled delivery of cell-derived regener-
ative factors, such as exosomes, without the regulatory hurdles
of delivering cells themselves. This was demonstrated recently
by Tao et al. who employed a chitosan hydrogel that medi-
ated sustained release of MSC exosomes. When tested in a rat
model, the material accelerated skin wound healing.[106]
4.2. Exploiting Hydrogel Physical Properties
to Direct Host Cell Response
In addition to releasing bioactive factors that can regulate
regeneration, researchers are also developing hydrogels with
physical properties, such as stiffness, degradability, and topog-
raphy, that are matched to those of the native tissue or can
recruit and/or direct cells in regeneration.[107,108] One area in
which this has been particularly successful is in CNS regenera-
tion, as hydrogels can be designed to have similar mechanical
properties to those of the brain.[108] HA-based hydrogels with
elastic moduli in the range of 3–10 kPa have been shown to
be optimal for neural progenitor cell (NPC) differentiation.[109]
Adv. Healthcare Mater. 2018, 1700939
Figure 8. Using acellular hydrogels to promote tissue regeneration. For regenerative strategies
based on acellular hydrogels, various factors can be incorporated into the hydrogel. Such fac-
over time as the hydrogel degrades. Acellular hydrogels can also be used to directly instruct
hostcells.Forexample,tissue-specificECMortopographycanalso beincorporatedintothe
hydrogel to direct host stem cell differentiation as cells come in contact with it. More advanced
acellular hydrogel strategies might incorporate controlled spatial and temporal release of spe-
cific factors. Using cleavable systems such as those mediated by MMP activity, factors will
only be released in response to cell- or tissue-specific stimuli. This could be particularly useful
for tissue interfaces such as the Achilles tendon-bone insertion (enthesis), where the spatial
distribution of biological factors is key in modulating cell behavior.
1700939 (14 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
NPC encapsulated in soft hydrogels that mimicked the stiff-
ness of a neonatal brain (2.6 kPa) differentiated into neurons.
However, when encapsulated in slightly stiffer hydrogels, more
akin to that of adult brain (5.7 kPa), NPC differentiated into
astrocytes.[109] For CNS regeneration, hydrogel degradation is
also key. Silk-based hydrogels have been shown to be advanta-
geous in this context as they have slow degradation rates that
may foster neural regeneration.[110]
Other physical features of hydrogels can also be exploited
to direct host cells. In addition to degradability, cell infiltration
into an acellular hydrogel will also depend on hydrogel porosity/
mesh size, the availability of cell adhesive motifs—often adhe-
sive peptides, syndecans,[111] or cadherins[112]—and the hydrogel’s
amenability to promote the formation of a vasculature. Cells need
to be within 100 to 200 µm of a blood vessel in order to receive
enough nutrients and oxygen for normal function.[113] Hydrogel
architecture appears to play an important role in modu lating this.
Chiu et al.[114] showed that pore sizes within PEG hydrogels
between 50 and 150 µm in diameter allowed for the formation
of vasculature, while pore sizes of 25–50 µm limited vessel infil-
tration.[114] Micropatterning techniques have also been exploited
to guide cell migration in acellular hydrogels. Lee et al. demon-
strated a photo laser scanning photolithography technique in
which RGD sites could be patterned into a collagenase-sensitive
poly(ethylene glycol-co-peptide) diacrylate hydrogel to direct cell
migration.[115] They showed that cell migration along patterns
promoted successful wound healing.
4.3. Hydrogels with Immunomodulatory
and Gene Therapy Functionality
The immune system plays an important role in tissue repair and
is emerging as a key target of biomaterial-based regenerative
strategies. Some biomaterials are known to interact directly with
particular aspects of the immune system.[116] HA, for example,
has anti-inflammatory properties, while natural materials such
as chitosan scavenge free radicals, which can then reduce/sup-
press inflammation.[116,117] Indeed, active control of the immune
system may be key in modulating tissue repair and regenera-
tion. Inflammation is an essential response to injury, and plays
a central role in initiating healing.[116] During normal healing,
acute inflammation is followed by a resolving anti-inflammatory
response, and the restoration of tissue integrity.[116] Issues arise
if inflammation becomes chronic and inhibits healing.[118] This
scenario provides an opportunity for hydrogels with dual func-
tionalities that can help mediate these responses. For example, an
initial hydrogel-based delivery of a proinflammatory signals such
as SDF-1 (an inflammatory and angiogenic cytokine) might aid
in the mobilization of progenitor cells.[119] This response could
then be followed by the release of anti-inflammatory media-
tors such as IL-4 and IL-10, which are key for tissue repair as
they trigger macrophages to switch from a proinflammatory to
a reparative phenotype.[120] Similarly to the MMP-cleavable sys-
tems that have been exploited to treat the effects of myocardial
infarction,[104] such systems could be designed to release a pro-
inflammatory molecule until an appropriate response from the
native tissue triggers release of an antiinflammatory molecule to
resolve tissue healing.
In addition to releasing inflammatory molecules, hydro-
gels can also be used to modulate the host’s immune system.
For example, biomaterials can be used induce an antigen-
specific tolerogenic response to treat a variety of autoimmune
diseases.[121,122] This was demonstrated by Verbeke et al. who
used an alginate hydrogel to deliver a peptide antigen mimotope
to treat type 1 diabetes in a mouse model. This peptide hydrogel
delivery system resulted in expansion of antigen specific T cells
in the lymph nodes.[122] Beyond treating autoimmune diseases,
similar therapies could also be used to modulate transplant
Gene therapy technologies similarly hold great promise for
regeneration.[123] With such therapies, successful delivery is
vital, making hydrogels an attractive platform to locally and sus-
tainably deliver appropriate molecules. For example, Yang et al.
used a PECE thermoresponsive hydrogel to deliver an antionco-
gene. This system was capable of sustainably and locally deliv-
ering the gene, which was important for maximizing its antitu-
mour effects, but minimizing systemic side effects.[124] Hydro-
gels have also proven useful as lentiviral delivery systems, and
there is evidence that they both increase the stability of the
virus as well as help protect it from the immune system.[125]
Indeed, for lentiviral gene therapy, host immune response
can affect efficacy. Hydrogels with small pore sizes can limit
complement and antibody diffusion, potentially protecting the
virus.[125] Hydrogels have been used for the delivery of nucleic
acids to treat spinal cord injury.[126] Aligned nanofiber hydrogel
scaffolds were shown to sustainably release microRNAs that
enhanced axon regeneration.[126] Scaffold design here medi-
ated alignment of the fibers by providing topographical cues to
direct neurite extension.[126] With the development of CRISPR/
Cas9 systems, hydrogels may also have the potential to play an
important role in genome editing technologies in the future.
5. Hydrogel Delivery Strategies
While researchers exploit hydrogel technologies to under-
stand how biochemical and biophysical cues influence cell and
tissue function to create 3D tissue models and TE platforms,
being able to exploit these cues with hydrogels depends upon
our ability to deliver them for appropriate therapeutic applica-
tions (Figure 9). Simple implantation of biomaterials through
traditional surgical means is a standard delivery route. How-
ever, surgery can lead to morbidity at the implantation site, is
associated with surgical and recovery costs, and will almost
always cause patient discomfort.[127] For example, to treat osteo-
chondral defects on the articulating surfaces of the knee, early
hydrogel implants (such as Cartipatch, an alginate-based con-
struct) required surgery, which could cause inflammation and
increased the chance of infection, which could result in treat-
ment failure.[128] To overcome these drawbacks there have been
significant efforts to design injectable, in situ forming, and
nano/microhydrogel systems that can be delivered in a mini-
mally invasive manner.
For any hydrogel-based therapeutic to be a viable clinical
option, the cross-linking/gelation mechanism must be bio-
logically mild as to not damage the surrounding tissue or
biologics, including cells, in the hydrogel. Injectable hydrogels
Adv. Healthcare Mater. 2018, 1700939
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that have been developed for tissue repair employ various gela-
tion methods that are mediated through chemical or physical
interactions. Deciding which gelation mechanism to use
depends on the application and the desired effect. Chemical
interactions are covalent bonds formed between precursors
that cross-link the hydrogel. Covalently formed hydrogels are
often robust and mechanically strong. For example, Michael
addition reactions between thiol groups and acrylates or meth-
acrylates have been employed to form covalently cross-linked
hydrogels and are often aided by UV light.[129] Physical interac-
tions, on the other hand, lead to hydrogels that have reversible
bonds between hydrogel components, and are achieved using
molecular interactions like hydrophobic interactions,
interactions, hydrogen bonding, metal chelation, and van der
Waals interactions.[52] Covalent cross-linking typically requires
physical initiators while noncovalent cross-linking often
arises spontaneously under specific conditions.[108] A combi-
nation of both covalent and physical interactions in the same
hydrogel has also been explored[130,131] to create hydrogels for
cartilage[132] and cornea repair.[133] For example, a network that
contains both rigid and ductile components can be mechani-
cally advantageous because the rigid network can sustain
load, while the ductile network dissipates energy, preventing
failure.[130] Indeed, Shin et al.[130] created a hydrogel that
was made of a covalently linked gellan gum network, which
had a second covalently linked gelatine network incorporated
within it. This biocompatible, double network could be loaded
with cells, and showed excellent load-bearing properties,
approaching those of native articular cartilage.
5.1. Injectable Hydrogels Cross-Linked by Light
UV irradiation-induced cross-linking is a common approach to
covalently cross-link hydrogels or activate biologics in hydrogels
in situ.[134–136] UV cross-linking allows quick gelation within
seconds[137] to minutes.[138] This characteristic is important as
it provides a means to quickly solidify a liquid precursor, which
otherwise may be diluted by blood or other body fluids. As well
as possessing quick gelation times, UV-cross-linkable hydrogels
can be designed with low molecular weight, and thus low-
viscosity, precursors. This is advantageous because it allows
the hydrogel to be injected and then take the shape of a defect
or cavity prior to cross-linking. This space filling property may
be crucial for proper integration with surrounding tissues and
Many types of biological and synthetic molecules/
polymers can be delivered and cross-linked via light, including
polyethylene glycol,[139] poly(2-hydroxyethyl methacrylate),[140]
modified HA,[141] and modified chitosan.[142] The ability to cova-
lently incorporate various moieties to mediate cross-linking and/
or incorporate biological components is commonly achieved by
Adv. Healthcare Mater. 2018, 1700939
Figure 9. Various strategies for hydrogel delivery. Schematic diagram showing a range of strategies for hydrogel delivery and gelation. A) Single cells
canbeencapsulatedinthinhydrogels,whichcanthenbeinjectedintravenouslytobedistributedthroughoutthe body.Thehydrogelcoatingcanbe
designedtoprotectthecellsfromthe immunesystem.B)Hydrogelscanbecovalentlycross-linkedbyenzyme-mediatedreactions betweenreactive
groups on the hydrogel monomers, forming a network. C) Noncovalent cross-linking of hydrogels can be achieved via hydrophobic, interactions,
interactions, hydrogen bonding, metal chelation, or van der Waals interactions. These hydrogels can sometimes be shear thinning, which makes them
spatiotemporal control of their biological activity. F) Surgical implantation of a hydrogel to treat a nonunion fracture. G) Micro- and nanohydrogels,
myocardial infarction.
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attaching acrylate, vinyl, methacrylate, or thiol groups to these
molecules or taking advantage of similar side groups. UV irradia-
tion-inducible cross-linking is often facilitated by non-toxic photo
initiators like Irgacure 2959,[143] which is part of the free radical-
dependent Michael addition that polymerizes the hydrogel via
the incorporated acrylate, vinyl, methacrylate, or thiol groups. UV
irradiation can also be used to spatially control specific features
like stiffness[144] and the presentation of bioactive molecules.[145]
The versatility and adaptability of light-mediated cross-linking
make it particularly attractive when developing hydrogel-based
therapies for in vivo use. Lin et al.[146] engineered an inject-
able, cell-laden hydrogel based on methacrylated gelatine. After
subcutaneous injection, the hydrogel was rapidly cross-linked
via UV irradiation and Irgacure 2959 through the skin. After
7 days, the hydrogels supported the formation of a stable vas-
culature if endothelial colony-forming cells or MSC were incor-
porated. UV-curing hydrogels have also been explored to cross-
link an injectable hydrogel transdermally to deliver model pro-
teins in a mouse model.[134]
While UV irradiation can be used to cross-link hydrogels in situ,
Lee et al. have shown how a similar technique can be used to acti-
vate bioactive components in an implantable hydrogel.[147] This was
achieved with a PEG-based hydrogel modified with RGD sequence-
containing peptides that were protected by a 3-(4,5-dimethoxy-2-
nitrophenyl)-2-butyl ester photolabile caging group, which could
be released upon exposure to UV light. Using this chemistry, they
spatially and temporally activated the RGD peptides using precise
and timed exposure to UV light that released the molecular cage.
In a transdermal mouse model, the application of UV light was
able to regulate cell adhesion, inflammation, fibrous encapsula-
tion, and vascularization.[147] This spatiotemporal control may
have important implications for researchers’ ability to direct tissue
repair because it has the potential to allow for the controlled
formation of vasculature and tissue heterogeneity.
Nevertheless, while UV-mediated cross-linking has numerous
advantages, including that it allows for fast, spatial, and temporal
control of gelation, UV light curing methods may impact cell
viability and negatively affect surrounding host tissue due to the
formation of radicals, and to a lesser extent, from the UV light
itself.[134,143,148] This method is also limited by the depth UV irra-
diation can travel through tissue and the amount of UV exposure
and radicals a tissue can withstand without damage. Indeed, UV-
induced gelation is unlikely to see applications deep in tissues.
To overcome potential toxicity associated with UV light curing
methods, systems which employ visible light to initiate cross-
linking have also been developed.[149] For example, Fu et al. devel-
oped a heparin/PEG-based visible light (525 nm) cross-linkable
hydrogel that uses eosin Y and triethanolamine as photoinitia-
tors. In this system, thiolated-heparin was cross-linked with PEG
diacrylate through a Michael-type addition to form a covalently
cross-linked network. Encapsulated fibroblasts remained viable
in these hydrogels and growth factor loading was comparable to
that achieved in other heparin-based systems.[150]
5.2. Stimulus-Driven In Situ Forming Hydrogels
Another approach to deliver therapeutic hydrogels is by
harnessing their potential for in situ cross-linking. Unlike
UV irradiation-induced cross-linking systems, which use
external stimuli, in situ forming hydrogels rely on specific phys-
iological conditions (pH/temperature),[151] delayed/secondary
cross-linking mechanisms,[136] or enzyme-mediated cross-
linking[152] to trigger hydrogel formation. This type of delivery
approach has many of the benefits of UV cross-linked systems,
like injectability, and fast gelation, but often lacks amenability
for spatiotemporal control of gelation. Some key advantages of
the in situ forming hydrogels are their ease of use in surgery
as well as their adaptability to various applications, such as in
deep tissues where light-mediated methods are not feasible.
Enzymatically cross-linked hydrogels are one promising form
of in situ cross-linked hydrogels. These hydrogels are polymer-
ized through the catalytic activity of enzymes added exogenously,
which facilitate the covalent bonding between two substrates that
are linked to larger molecules that comprise the bulk hydrogel.
Enzymes that have been explored to induce gelation include
transglutaminase, tyrosinase, phosphopantetheinyl transferase,
lysyl oxidase, plasma amine oxidase, phosphatases, thermolysin,
-lactamase, phosphatase/kinase, and peroxidases.[153] Enzymes
of particular interest are transglutaminases, which cross-link via
the formation of ε-(
-glutamyl)lysine bonds or the incorporation
of primary amines with glutamine residues.[154] Transglutami-
nases are advantageous because they can facilitate a tight inte-
gration of the hydrogel with the surrounding host tissue. This
tight integration between the native tissue and hydrogel is due
to the readily available naturally occurring substrate for transglu-
taminases, which exists in native tissues.[155]
A novel use of enzyme-mediated cross-linked hydrogels was
recently reported by Griffin et al.,[156] who engineered PEG-
based microhydrogel beads that were modified by covalently
attaching glutamine and lysine. These molecules can act as
substrates for factor XIII, a transglutaminase enzyme involved
in blood coagulation. Upon addition of exogenous factor XIII,
the hydrogel beads were cross-linked via reactions between the
primary amine on the lysine with the glutamine to produce
porous hydrogels that human dermal fibroblasts, adipose-
derived MSC, and bone marrow-derived MSC could infiltrate.
The authors also injected these materials into a mouse skin
wound model and cross-linked them with exogenous factor
XIII in situ to form a porous scaffold. Wounds treated with the
cross-linked microhydrogels showed faster wound closure and
decreased inflammation compared to those treated with a non-
porous hydrogel made of the same material. Other groups have
also used enzyme cross-linked hydrogels for cartilage TE. For
example, gelatin-hydroxyphenylpropionic acid has been cross-
linked with horseradish peroxidase catalyzed by H2O2 to form
a viscoelastic hydrogel. When these enzyme cross-linked hydro-
gels were used to encapsulate chondrocytes and then injected
into an osteochondral defect in a rabbit, improved cartilage
regeneration was reported in comparison to that in no hydrogel
Nevertheless, while enzymatically cross-linked hydrogels
show promise, the inclusion of exogenous enzymes may have
unforeseen deleterious biological effects. For example, exog-
enous application of transglutaminases in vivo most likely
also cross-links ECM proteins in the tissue surrounding the
hydrogel, which can lead to tissue stiffening. Indeed, while
transglutaminases are essential for many biological processes,
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they can also contribute to the pathophysiology of various
inflammatory, autoimmune, and degenerative conditions.[158]
5.3. Nonstimulus-Driven In Situ Cross-Linking Hydrogels
Hydrogels covalently cross-linked in situ via nonstimulus driven
mechanisms provide another option to deliver therapeutic
hydrogels. These hydrogels are typically composed of precursor
components that are mixed immediately before or during injec-
tion, often rely on simple gelation mechanisms, and are easy to
use. One in situ cross-linking method that has shown promise
is click chemistry. Click chemistry refers to a number of dif-
ferent types of chemical reactions that are highly efficient, even
at low concentration. This is a particular advantage for biological
hydrogels because of the low solid content of most systems.
Click reactions are also advantageous because of their quick
reaction rates and biocompatible reaction conditions. Unlike
other mechanisms, such as enzyme-mediated cross-linking,
click reactions are highly selective chemically, meaning that they
do not readily produce unwanted side reactions. They have also
been shown to be compatible with cells, drugs, and proteins.[159]
Hermann et al.,[160] for example, developed an injectable click-
based hydrogel to delay bone growth in a mouse calvarial model
in which rapid regrowth of bone can lead to craniofacial deformi-
ties, restricted brain growth, and an increase in intracranial pres-
sure. They utilized multivalent PEG precursors that formed a
network upon mixing via ring-strain promoted Cu-free reactions
between dibenzylcyclooctynes and azides attached to the PEG.
Importantly, by incorporating the BMP antagonist Gremlin into
their hydrogel, they were able to delay bone re-growth, dem-
onstrating a potential therapy to control bone over-growth and
reduce the risk of life threatening complications.[160]
As well as being able to form hydrogels to deliver biologics,
click chemistry has also been exploited to construct micropo-
rous hydrogels that when formed in the presence of cells,
create complex polymer–cell composite systems with variable
pore sizes.[161] To accomplish this, microgels were formed via
an inverse suspension polymerization method with either
dibenzocyclooctyne or azide groups. When these two microgels
were combined, they formed a covalently connected microgel
network that could entrap cells. This microgel system provided
tunable properties that controlled cell–material interactions and
cell morphology.
5.4. Intravenous Delivery of Cellular Nano-/Microgels
Although there have been a plethora of clinical trials with MSC,
it is clear that simply injecting cells intravenously limits their
therapeutic value.[162] Injected cells often can only briefly par-
ticipate as immunomodulatory or signaling cells before they
are cleared from the body.[163] While many of the hydrogels
discussed so far have focused on bulk materials that provide
physical support to cells, techniques have also been developed
to encapsulate cells within a thin layer of hydrogel to facilitate
cell-based therapies.[164] The goal of this approach is often to
deliver signaling cells that can modulate immune or regen-
erative responses via paracrine signaling, or signal to other
therapeutically relevant cells/recruit them to a specific site
to participate in repair. For example, single MSC have been
encapsulated with a 5.8 µm thick sodium ion cross-linked algi-
nate hydrogel coatings to form cell-containing microgel beads.
When delivered intravenously in a mouse model, encapsulated
cells remained in the mouse for longer and were shown to pro-
vide a longer sustained release of cell-secreted factors than cells
injected without coating.[163]
One drawback of intravenous injection is that radially
injected cells are often trapped in the lungs,[165] likely due to
their size.[166] Cells coated with a thin hydrogel, however, may
fair a better chance of evading immune system clearance as
hydrogels can be designed to preclude cell detection or simply
provide a physical barrier that discourages engulfment. Hydro-
gels may also provide other interesting solutions to the problem
of pulmonary passage. By incorporating small molecules or
biologics on the surface of cell containing microgels, hydrogels
could passively direct encapsulated cells to an area of damage
or disease before they reach the lungs. Hydrogels could also
act as homing vehicles to direct therapeutically relevant cells to
specific locations in the body.
6. Clinical Translation of
Hydrogel-Based Therapies
The majority of the drugs and therapeutics developed for
human use fail during their discovery and development stages.
Indeed, only 10.4% of all candidates put through Phase I clinical
trials eventually receive approval for use in humans.[167] For more
than 20 years now, TE-based therapies have been proposed and
pushed through the development stages toward clinical trans-
lation in the hope of creating functional tissues to replace
those lost to disease or injury.[168] As of July 2017, there were
371 clinical trials registered worldwide that related to hydrogels
and 69 that focused on TE ( Although
only a handful of these studies aimed to apply recent advances
in hydrogels to TE, there have been some limited successes,
particularly for planar and hollow organs, such as skin, cornea,
urethra, urinary bladder, and blood vessels.[169] However, the
development of TE-based regenerative strategies for more
complex tissues still faces a number of key challenges.[168,170]
These include: (1) provision of adequate oxygen and nutrients
to large tissues, which likely require the formation of a com-
plex vasculature; (2) incorporation of multiple cell types with
precise spatial arrangements; (3) achievement of appropriate,
tissue-specific mechanical properties (such as stiffness, shear
strength and hardness); and (4) integration of TE constructs
with surrounding tissue. Hydrogel-based TE scaffolds offer the
possibility of addressing many of these challenges and perhaps
can be successfully translated into viable therapies. We high-
light some promising preclinical and clinical studies that have
exploited hydrogels for TE-based therapies below.
6.1. Cartilage Repair
Articular cartilage is an obvious target for clinical repair using
hydrogel technologies. The market for cartilage repair is
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tremendous given that 13.8% of the adult population over
the age of 60 suffers from osteoarthritis, a number that will
increase as the global population ages.[171] Articular cartilage
is amenable to TE strategies because it lacks some of the chal-
lenges of more complex tissues because it is, for the most part,
avascular, aneural, and only contains a single type of cell. Much
research has been undertaken with hydrogels to regenerate car-
tilage,[172] including using naturally derived ECM-based hydro-
gels such as type I collagen, and thermoresponsive hydrogels
formed from chitosan/PVA composites. There are also reports
of efforts to engineer cartilage with synthetic hydrogels,[173]
including one based on a PLGA–PEG–PLGA hydrogel.[174]
One particularly promising clinical application of a synthetic
hydrogel for cartilage repair was reported by Sharma et al. in
2013.[4] They developed a photoreactive, adhesive, PEGDA-
based hydrogel, which they injected into focal cartilage defects
in the medial femoral condyle of 15 patients in conjunction
with standard microfracture surgery. Although the injected
hydrogel itself was acellular, the concomitant microfracture sur-
gery allowed autologous cells to invade into the hydrogel. When
compared to outcomes in patients treated with microfracture
alone, those who received the hydrogel-based therapy showed
greater tissue fill and increased tissue organization (both by
MRI). Patients also reported less pain, an important clinical
outcome in this patient group. Although long-term follow-up
is required to determine if the benefits of the treatment persist,
and importantly, if the treatment ultimately prevents patients
with focal lesions from going on to develop osteoarthritis,
results appear promising.
A number of preclinical trials in cartilage regenera-
tion using hydrogels have also been reported. For example,
researchers incorporated kartogenin and MSC into a syn-
thetic PLGA–PEG–PLGA hydrogel and showed good cartilage
repair in a rabbit model.[174] Similarly, a hydrogel formed from
oligo[poly(ethylene glycol)] fumarate combined with encapsu-
lated MSC, was reported to mediate a more hyaline-like repair
than implantation of the scaffold alone in a porcine model.[175]
6.2. Cardiovascular Regeneration
The other application of hydrogels for TE that dominates pre-
clinical and clinical trials is for cardiovascular applications. Car-
diovascular diseases place a tremendous burden on society and
are the leading cause of mortality worldwide.[176] Particularly in
the United States and Europe, around 5% of the acute hospital
admissions are due to cardiac events, and 10% of hospitalized
patients suffer from some form of cardiovascular disease.[176]
In a completed Phase I study (NCT00981006), a gelatin
hydrogel sheet that allowed for the controlled release of bFGF
(200 µg) was combined with autologous cardiac-derived stem
cells (5 × 105 cell kg1) to treat six patients with ischemic cardio-
myopathy following acute myocardial infarction.[177] In this
study, Takehara et al. showed that the therapy was safe and
feasible. However, while some benefits were noted 6 months
posttreatment, the sample size was too small to conclude
efficacy.[177] Similarly, a phase I clinical trial (NCT02057900)
led by Assistance Publique–Hôpitaux de Paris investigated the
feasibility of human embryonic stem cell (ESC)-derived
CD15+ Isl-1+ progenitors embedded within a fibrin hydrogel
patch to treat patients with ischemic heart failure.[178] In this
study, a fibrin hydrogel patch was used, as it had been shown
previously in animal models to play an essential role in
improving cell retention and survival.[179] While no conclusions
could be drawn on efficacy, overall patient functional outcomes
were promising.[178]
Despite some promising results, the overall efficacy of car-
diovascular cell therapies is inconclusive. A meta-analysis that
examined treatments for heart failure showed that among 31
randomized cell therapy trials comprising 1521 patients, exer-
cise capacity, left ventricular ejection fraction, and quality of
life were all improved in the treated patients.[180] However,
a second meta-analysis that examined patient data from 12
trials could find no benefit of intracoronary cell therapies in
treating acute myocardial infarction.[181] Such conflicting data
raise concerns about cell-based therapies for cardiovascular
diseases and highlight the need for improved strategies. For
future treatments, multidisciplinary approaches may provide
an answer. Cell retention in cardiac tissue, for example, is
thought to be important, and hydrogels have the capability to
improve this.[182] Successful differentiation into mature car-
diomyocytes is likely also key, and again, could be aided by
careful scaffold design. Indeed, many cell types are known to
be mechanoresponsive, including cardiomyocytes. Work by
Morez et al. showed that modifying material surface topog-
raphy by creating microfabricated grooves could promote car-
diac progenitor elongation and alignment, driving appropriate
differentiation.[183] Cell-containing cardiac patches that rely
on bioprinting technologies are similarly promising. Using a
bioprinting strategy, precise patterns of human umbilical vein
endothelial cells and hMSC were created; and when implanted
in a rat model of myocardial infarction, led to increased
angiogenesis and improved cardiac function in comparison
to patches in which cells were seeded randomly.[184] Another
consideration for successful treatment of myocardial infarc-
tion is delivery time. A disadvantage of patch-based delivery
systems is that they often require surgical procedures. Inject-
able hydrogels, on the other hand, can be delivered quickly
to the site of injury.[185] Indeed, injectable hydrogels could be
delivered minimally invasively as transcoronary infusions or
transendocardial injections.[185] Such fast delivery of cells/scaf-
folds may help prevent damage to the heart tissue postinfarc-
tion. While no such injectable hydrogel systems have been
used clinically, future developments in dynamic, injectable
hydrogels may make this achievable.
7. Conclusions
Over the last decade, hydrogel technologies have improved
dramatically allowing researchers to create ex vivo tissue
models that replicate that native tissue better than ever before.
Researchers are also developing hydrogel-based biomaterials
with controlled architectures and biological and physical prop-
erties that can be used for TE. Moreover, acellular hydrogels
that can deliver bioactive molecules are increasingly finding
use in drug delivery applications, often relying on their inject-
ability and chemistries that allow for in situ gelation. Taken
Adv. Healthcare Mater. 2018, 1700939
1700939 (19 of 22) © 2018 The Authors. Published by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
together, the result is a toolbox of hydrogel-based materials
that can be used by both bench-based researchers to answer
fundamental questions in cell biology, and physicians to either
replace damaged tissues or deliver cells/molecules to mediate
repair. Indeed, the advancements in regenerative medicine that
these new hydrogel technologies are likely to foster are surely
something to feel “swell” about.
Despite such exciting developments, however, developments
in hydrogel research only seem to trickle into a few limited clin-
ical applications. This may be partly attributable to regulatory
and funding limitations, but may also be because of risk. Tissue
regeneration is clearly a complex process, which will require com-
plex biological and material-based strategies to tackle. Indeed, it is
likely that only through careful consideration of the both the target
tissue’s biological, physical and mechanical properties can hydro-
gels be developed that can truly mediate or participate in regen-
eration. However, as fundamental research in cell biology reveals
how cells respond to their ECM and their niche, and new insights
into endogenous tissue regenerative mechanisms are identified,
materials scientists and chemists will inevitably develop hydro-
gels that can exploit and target them for regeneration.
D.A.F. was supported by a fellowship from the Whitaker International
Program. M.D.A.N. acknowledges support from the BBSRC London
Interdisciplinary Doctoral Programme (LIDo). E.G. acknowledges a
Research Career Development Fellowship from the Wellcome Trust and
Conflict of Interest
The authors declare no conflict of interest.
advanced therapies, biomaterials, bioprinting, hydrogels, regenerative
medicine, tissue engineering
Received: August 4, 2017
Revised: October 24, 2017
Published online:
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... For instance, synthetic biomaterials allow the modulation of individual parameters (e.g., stiffness, viscoelasticity, degradation) on cell behaviour, in a reductionist and stepwise manner Lust et al., 2021). The most common synthetic biomaterial that has been used for organoid cultures is polyethylene glycol (PEG), owing to its high tunability, design flexibility, resistance to absorption, and low cost (Caliari and Burdick, 2016;Foyt et al., 2018;Kratochvil et al., 2019). PEG-based hydrogels are inherently inert but can be functionalized through the addition of adhesive peptides, including RGD and IKVAV motifs for fibronectin and laminin, respectively (Foyt et al., 2018). ...
... The most common synthetic biomaterial that has been used for organoid cultures is polyethylene glycol (PEG), owing to its high tunability, design flexibility, resistance to absorption, and low cost (Caliari and Burdick, 2016;Foyt et al., 2018;Kratochvil et al., 2019). PEG-based hydrogels are inherently inert but can be functionalized through the addition of adhesive peptides, including RGD and IKVAV motifs for fibronectin and laminin, respectively (Foyt et al., 2018). These integrin-binding domains can influence intercellular communication depending on their concentration, spacing, presentation timing, and patterning (Kratochvil et al., 2019). ...
Full-text available
Cell-cell interactions underlay organ formation and function during homeostasis. Changes in communication between cells and their surrounding microenvironment are a feature of numerous human diseases, including metabolic disease and neurological disorders. In the past decade, cross-disciplinary research has been conducted to engineer novel synthetic multicellular organ systems in 3D, including organoids, assembloids, and organ-on-chip models. These model systems, composed of distinct cell types, satisfy the need for a better understanding of complex biological interactions and mechanisms underpinning diseases. In this review, we discuss the emerging field of building 3D multicellular systems and their application for modelling the cellular interactions at play in diseases. We report recent experimental and computational approaches for capturing cell-cell interactions as well as progress in bioengineering approaches for recapitulating these complexities ex vivo . Finally, we explore the value of developing such multicellular systems for modelling metabolic, intestinal, and neurological disorders as major examples of multisystemic diseases, we discuss the advantages and disadvantages of the different approaches and provide some recommendations for further advancing the field.
... Hydrogels are water-swollen three-dimensional polymer networks [1], formed by covalent bonds or noncovalent bonds such as hydrogen bonds and van der Waals forces [2]. The interaction between hydrogels counterbalance bonds, leading to expansion of polymer networks, which also regulates hydrogels' internal transport, diffusion characteristics, and mechanical strength. ...
... As capsules in the second strategy may affect the structure of the material and limit self-healing, self-healing materials that apply dynamic interactions have gained increasing attention. Multiple types of reversible interactions can be used to design self-healing hydrogels, and variable concentrations and strengths of bonds contribute to the unique physicochemical properties of the hydrogel [2,77]. Here, we have mainly summarized the mechanisms of self-healing hydrogels based on reversible interactions and discussed their applications in the medical field. ...
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A hydrogel is a three-dimensional structure that holds plenty of water, but brittleness largely limits its application. Self-healing hydrogels, a new type of hydrogel that can be repaired by itself after external damage, have exhibited better fatigue resistance, reusability, hydrophilicity, and responsiveness to environmental stimuli. The past decade has seen rapid progress in self-healing hydrogels. Self-healing hydrogels can automatically self-repair after external damage. Different strategies have been proposed, including dynamic covalent bonds and reversible noncovalent interactions. Compared to traditional hydrogels, self-healing gels have better durability, responsiveness, and plasticity. These features allow the hydrogel to survive in harsh environments or even to be injected as a drug carrier. Here, we summarize the common strategies for designing self-healing hydrogels and their potential applications in clinical practice.
... They can be used for exosome encapsulation or protection against the dynamic circulation of tissue fluids and other environmental factors. There are several scaffold fabrication technologies used for exosome encapsulation, from bulk to micro-/nanofabrication methods, such as 3D printers [17]. The following sections will make a comprehensive review of the mentioned methods and their application in similar research works. ...
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Exosomes, a subset of vesicles generated from cell membranes, are crucial for cellular communication. Exosomes' innate qualities have been used in recent studies to create nanocarriers for various purposes, including medication delivery and immunotherapy. As a result, a wide range of approaches has been designed to utilize their non-immunogenic nature, drug-loading capacity, or targeting ability. In this study, we aimed to review the novel methods and approaches in exosome engineering for encapsulation and targeting in regenerative medicine. We have assessed and evaluated each method's efficacy, advantages, and disadvantages and discussed the results of related studies. Even though the therapeutic role of non-allogenic exosomes has been demonstrated in several studies, their application has certain limitations as these particles are neither fully specific to target tissue nor tissue retainable. Hence, there is a strong demand for developing more efficient encapsulation methods along with more accurate and precise targeting methods, such as 3D printing and magnetic nanoparticle loading in exosomes, respectively. Graphical abstract
... Biocompatible nanomaterials and hydrogels are increasingly applied in the biomedical field as medical devices, including drug delivery systems, cell therapy, tissue engineering, and wound dressings [1][2][3][4][5][6][7][8][9][10]. ...
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Hydrogel wound dressing makes easier the treatment of patients suffering from difficult wounds. A new process for the manufacturing of a sterile, packaged hydrogel wound dressing, based on an interpenetrating structure of calcium alginate, agar, and polyvinylpyrrolidone, was recently developed. The new formulation overtakes some previous technologies' drawbacks expressing a better resistance to mechanical deformations compared to products on the market. In this work, the 2.35 T proton density, spin-lattice relaxation time, spin-spin relaxation time, phase-coherence relaxation, and water apparent diffusion coefficient analysis in the new hydrogel and several alternative formulations, including a commercial one (Neoheal®), are reported. Specifically, the combination of agar, acting as a thermolabile forming agent, with calcium alginate and γ irradiated polyvinylpyrrolidone, acting, respectively, as physical, and chemical crosslinking agents with an irreversible (temperature independent) effect, have been investigated. The new hydrogel formulation brings a qualitative improvement in its handling due to its increased mechanical stiffness when compared to the commercial hydrogel reference. This comes together with a reduced water content (100 vs. 112 for proton density in arbitrary units) and swelling capacity (88% vs. 124%) but with improved water mobility (1.42 vs. 1.34 × 10−3 mm2 s−1 for the apparent diffusion coefficient).
The promise of cell therapy has been augmented by introducing biomaterials, where intricate scaffold shapes are fabricated to accommodate the cells within. In this review, we first discuss cell encapsulation and the promising potential of biomaterials to overcome challenges associated with cell therapy, particularly cellular function and longevity. More specifically, cell therapies in the context of autoimmune disorders, neurodegenerative diseases, and cancer are reviewed from the perspectives of preclinical findings as well as available clinical data. Next, techniques to fabricate cell-biomaterials constructs, focusing on emerging 3D bioprinting technologies, will be reviewed. 3D bioprinting is an advancing field that enables fabricating complex, interconnected, and consistent cell-based constructs capable of scaling up highly reproducible cell-biomaterials platforms with high precision. It is expected that 3D bioprinting devices will expand and become more precise, scalable, and appropriate for clinical manufacturing. Rather than one printer fits all, seeing more application-specific printer types, such as a bioprinter for bone tissue fabrication, which would be different from a bioprinter for skin tissue fabrication, is anticipated in the future.
A class of hydrophilic polymers known as "hydrogels" have extensive water content and three-dimensional crosslinked networks. Since the old period, they have been utilized as plant culture substrates to get around the drawbacks of hydroponics and soil. Numerous hydrogels, particularly polysaccharides with exceptional stability, high clarity, and low cost can be employed as plant substrates. Although numerous novel and functionalized hydrogels might assist in overcoming the drawbacks of conventional media and giving them more functions, the existing hydrogel-based plant growth substrates rarely benefit from the developments of gels in the previous few decades. Prospects include the development of new conduction techniques, the creation of potential new hydrogels, and the functionalization of the hydrogel as plant culture substrates.
Cardiovascular diseases have been the leading cause of death worldwide during the past several decades. Cell loss is the main problem that resulted in cardiac dysfunction and further mortality. Cell therapy aiming to replenish the lost cells is proposed to treat cardiovascular diseases especially ischemic heart diseases which lead to a big portion of cell loss. Due to the direct injection's low cell retention and survival ratio, cell therapy using biomaterials as cell carriers attracts more and more attention because of their promotion of cell delivery and maintenance at the aiming sites. In this review, we systematically summarized the three main factors involved in cell therapy for myocardial tissue regeneration: cell sources (somatic cells, stem cells and engineered cells), chemical components of cell carriers (natural materials, synthetic materials and electroactive materials), and categories of cell delivery materials (patches, microspheres, injectable hydrogels, nanofiber and microneedles, etc.). An introduction of the methods including magnetic resonance/radionuclide/photoacoustic and fluorescence imaging for tracking the behavior of transplanted cells in vivo is also included. Current challenges of biomaterials‐based cell therapy and their future directions are provided to give both beginners and professionals a clear view of the development and future trends in this area. This article is protected by copyright. All rights reserved
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Acute myocardial infarction and chronic heart failure rank among the major causes of morbidity and mortality worldwide. Except for heart transplantation, current therapy options only treat the symptoms but do not cure the disease. Stem cell-based therapies represent a possible paradigm shift for cardiac repair. However, most of the first-generation approaches displayed heterogeneous clinical outcomes regarding efficacy. Stemming from the desire to closely match the target organ, second-generation cell types were introduced and rapidly moved from bench to bedside. Unfortunately, debates remain around the benefit of stem cell therapy, optimal trial design parameters, and the ideal cell type. Aiming at highlighting controversies, this article provides a critical overview of the translation of first-generation and second-generation cell types. It further emphasizes the importance of understanding the mechanisms of cardiac repair and the lessons learned from first-generation trials, in order to improve cell-based therapies and to potentially finally implement cell-free therapies.
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Abstract Background There has been increased interest in one-step cell-free procedures to avoid the problems related to cell manipulation and its inherent disadvantages. We have studied the chondrogenic induction ability of a PAMPS/PDMAAm double-network (DN) gel and found it to induce chondrogenesis in animal osteochondral defect models. The purpose of this study was to investigate whether the healing process and the degree of cartilage regeneration induced by the cell-free method using DN gel are influenced by the size of osteochondral defects. Methods A total of 63 mature female Japanese white rabbits were used in this study, randomly divided into 3 groups of 21 rabbits each. A 2.5-mm diameter osteochondral defect was created in the femoral trochlea of the patellofemoral joint of bilateral knees in Group I, a 4.3-mm osteochondral defect in Group II, and a 5.8-mm osteochondral defect in Group III. In the right knee of each animal, a DN gel plug was implanted so that a vacant space of 2-mm depth was left above the plug. In the left knee, we did not conduct any treatment to obtain control data. Animals were sacrificed at 2, 4, and 12 weeks after surgery, and gross and histological evaluations were made. Results The present study demonstrated that all sizes of the DN gel implanted defects as well as the 2.5mm untreated defects showed cartilage regeneration at 4 and 12 weeks. The 4.3-mm and 5.8-mm untreated defects did not show cartilage regeneration during the 12-week period. The quantitative score reported by O’Driscoll et al. was significantly higher in the 4.3-mm and 5.8-mm DN gel-implanted defects than the untreated defects at 4 and 12 weeks (p
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Bioprinting has emerged as a novel technological approach with the potential to address unsolved questions in the field of tissue engineering. We have recently shown that Laser Assisted Bioprinting (LAB), due to its unprecedented cell printing resolution and precision, is an attractive tool for the in situ printing of a bone substitute. Here, we show that LAB can be used for the in situ printing of mesenchymal stromal cells, associated with collagen and nano-hydroxyapatite, in order to favor bone regeneration, in a calvaria defect model in mice. Also, by testing different cell printing geometries, we show that different cellular arrangements impact on bone tissue regeneration. This work opens new avenues on the development of novel strategies, using in situ bioprinting, for the building of tissues, from the ground up.
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Cartilage lesions can progress into secondary osteoarthritis and cause severe clinical problems in numerous patients. As a prospective treatment of such lesions, human-derived induced pluripotent stem cells (iPSCs) were shown to be 3D bioprinted into cartilage mimics using a nanofibrillated cellulose (NFC) composite bioink when co-printed with irradiated human chondrocytes. Two bioinks were investigated: NFC with alginate (NFC/A) or hyaluronic acid (NFC/HA). Low proliferation and phenotypic changes away from pluripotency were seen in the case of NFC/HA. However, in the case of the 3D-bioprinted NFC/A (60/40, dry weight % ratio) constructs, pluripotency was initially maintained, and after five weeks, hyaline-like cartilaginous tissue with collagen type II expression and lacking tumorigenic Oct4 expression was observed in 3D -bioprinted NFC/A (60/40, dry weight % relation) constructs. Moreover, a marked increase in cell number within the cartilaginous tissue was detected by 2-photon fluorescence microscopy, indicating the importance of high cell densities in the pursuit of achieving good survival after printing. We conclude that NFC/A bioink is suitable for bioprinting iPSCs to support cartilage production in co-cultures with irradiated chondrocytes.
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A major challenge in tissue engineering is to generate a functional microvasculature that ensures proper blood perfusion and connection with surrounding tissues. Strategies such as the incorporation of growth factors have been proposed to induce the growth of new blood vessels into engineered tissue, but limitations remain. Herein a novel chitosan–fibrin (CF)-based self-healing hydrogel with a modulus of ~1.2 kPa was developed. The self-healing hydrogel was found to be injectable and to degrade ~70% in 2 weeks. Vascular endothelial cells seeded in the CF hydrogel were able to form capillary-like structures. Moreover, the injection of the CF hydrogel alone promoted angiogenesis in the perivitelline space of zebrafish and rescued the blood circulation in ischemic hindlimbs of mice. The excellent self-healing and angiogenic capacities of the hydrogel may be associated with the formation of an interpenetrating polymer network structure between chitosan and fibrin. This unique self-healing hydrogel offers new possibilities for future applications to vascular repair.
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Osteoarthritis (OA) is a common cause of pain and disability and is often associated with the degeneration of articular cartilage. Lesions to the articular surface, which are thought to progress to OA, have the potential to be repaired using tissue engineering strategies; however, it remains challenging to instruct cell differentiation within a scaffold to produce tissue with appropriate structural, chemical and mechanical properties. We aimed to address this by driving progenitor cells to adopt a chondrogenic phenotype through the tailoring of scaffold composition and physical properties. Monomeric type-I and type-II collagen scaffolds, which avoid potential immunogenicity associated with fibrillar collagens, were fabricated with and without chondroitin sulfate (CS) and their ability to stimulate the chondrogenic differentiation of human bone marrow-derived mesenchymal stem cells was assessed. Immunohistochemical analyses showed that cells produced abundant collagen type-II on type-II scaffolds and collagen type-I on type-I scaffolds. Gene expression analyses indicated that the addition of CS – which was released from scaffolds quickly – significantly upregulated expression of type II collagen, compared to type-I and pure type-II scaffolds. We conclude that collagen type-II and CS can be used to promote a more chondrogenic phenotype in the absence of growth factors, potentially providing an eventual therapy to prevent OA.
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Spinal cord injuries (SCI) often lead to persistent neurological dysfunction due to failure in axon regeneration. Unfortunately, currently established treatments, such as direct drug administration, do not effectively treat SCI due to rapid drug clearance from our bodies. Here, we introduce a three-dimensional aligned nanofibers-hydrogel scaffold as a bio-functionalized platform to provide sustained non-viral delivery of proteins and nucleic acid therapeutics (small non-coding RNAs), along with synergistic contact guidance for nerve injury treatment. A hemi-incision model at cervical level 5 in the rat spinal cord was chosen to evaluate the efficacy of this scaffold design. Specifically, aligned axon regeneration was observed as early as one week post-injury. In addition, no excessive inflammatory response and scar tissue formation was triggered. Taken together, our results demonstrate the potential of our scaffold for neural tissue engineering applications.
In this study, by combining photopolymerization and particle leaching technique, in situ formation of porous hydrogel with pore interconnectivity was demonstrated in vivo upon subcutaneous injection into the back of mice as well as in vitro. A precursor solution containing thiolated heparin, PEG-diacrylate (PEG-DA), and gelatin microparticles (GMPs) as a fast dissolving porogen were photopolymerized by visible-light-initiated thiol-ene reaction with eosin Y (EY) as a photo initiator and triethanolamine (TEOA) as a co-initiator. Formation of porous structure of the hydrogel after subsequent leaching of GMPs was confirmed in an animal model as well as in a physiological environment. The physical characteristics of the hydrogel were analyzed, and the acute in vivo biocompatibility of this system was characterized.
While microporous scaffolds are increasingly used for regenerative medicine and tissue repair applications, the most common techniques to fabricate these scaffolds use templating or top-down fabrication approaches. Cytocompatible bottom-up assembly methods afford the opportunity to assemble microporous systems in the presence of cells and create complex polymer-cell composite systems in situ. Here, microgel building blocks with clickable surface groups are synthesized for the bottom-up fabrication of porous cell-laden scaffolds. The facile nature of assembly allows for human mesenchymal stem cells to be incorporated throughout the porous scaffold. Particles are designed with mean diameters of ≈10 and 100 µm, and assembled to create varied microenvironments. The resulting pore sizes and their distribution significantly alter cell morphology and cytoskeletal formation. This microgel-based system provides numerous tunable properties that can be used to control multiple aspects of cellular growth and development, as well as providing the ability to recapitulate various biological interfaces.
This review discusses supramolecular biofunctional materials, a novel class of biomaterials formed by small molecules that are held together via noncovalent interactions. The complexity of biology and relevant biomedical problems not only inspire, but also demand effective molecular design for functional materials. Supramolecular biofunctional materials offer (almost) unlimited possibilities and opportunities to address challenging biomedical problems. Rational molecular design of supramolecular biofunctional materials exploit powerful and versatile noncovalent interactions, which offer many advantages, such as responsiveness, reversibility, tunability, biomimicry, modularity, predictability, and, most importantly, adaptiveness. In this review, besides elaborating on the merits of supramolecular biofunctional materials (mainly in the form of hydrogels and/or nanoscale assemblies) resulting from noncovalent interactions, we also discuss the advantages of small peptides as a prevalent molecular platform to generate a wide range of supramolecular biofunctional materials for the applications in drug delivery, tissue engineering, immunology, cancer therapy, fluorescent imaging, and stem cell regulation. This review aims to provide a brief synopsis of recent achievements at the intersection of supramolecular chemistry and biomedical science in hope of contributing to the multidisciplinary research on supramolecular biofunctional materials for a wide range of applications. We envision that supramolecular biofunctional materials will contribute to the development of new therapies that will ultimately lead to a paradigm shift for developing next generation biomaterials for medicine.