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Hydrogels for protein delivery in tissue engineering ⁎ Roberta Censi

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Hydrogels for protein delivery in tissue engineering ⁎ Roberta Censi
COREL-06235; No of Pages 14
Journal of Controlled Release xxx (2012) xxx–xxx
Contents lists available at SciVerse ScienceDirect
Journal of Controlled Release
journal homepage: www.elsevier.com/locate/jconrel
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Roberta Censi a,⁎, Piera Di Martino a, Tina Vemonden b, Wim E. Hennink b
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School of Pharmacy, University of Camerino, via S. Agostino 1, 62032, Camerino (MC), Italy
Department of Pharmaceutics, Utrecht Institute for Pharmaceutical Sciences,Utrecht University, PO Box 80082, 3508 TB Utrecht, The Netherlands
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Article history:
Received 13 January 2012
Accepted 2 March 2012
Available online xxxx
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Tissue defects caused by diseases or trauma present enormous challenges in regenerative medicine. Recently, a
better understanding of the biological processes underlying tissue repair led to the establishment of new approaches in tissue engineering which comprise the combination of biodegradable scaffolds and appropriate
cells together with specific environmental cues, such as growth or adhesive factors. These factors (in fact proteins) have to be loaded and sustainably released from the scaffolds in time. This review provides an overview
of the various hydrogel technologies that have been proposed to control the release of bioactive molecules of interest for tissue engineering applications.
In particular, after a brief introduction on bioactive protein drugs that have particular relevance for tissue engineering, this review will discuss their release mechanisms from hydrogels, their encapsulation and immobilization methods and will overview the main classes of hydrogel forming biomaterials used in vitro and in vivo to
release them.
Finally, an outlook on future directions and a glimpse into the current clinical developments are provided.
© 2012 Published by Elsevier B.V.
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Keywords:
Hydrogels
Growth factors
Protein delivery
Tissue engineering
Controlled release
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1. Introduction
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Controlled drug delivery has seen rapid advances in the last few decades with the introduction of novel biomaterials and technologies that
found application in all fields of pharmaceutical and biomedical sciences.
Particularly, with the advent of protein therapeutics, the need for controlled delivery systems able to enhance protein's pharmacokinetic and
pharmacodynamic properties became more urgent. Nowadays, proteins
are used in the treatment of many diseases but also in the area of tissue
engineering, where supply of biomolecular cues that mimic the environment of natural tissues and promote the communication between cells
proved crucial for achieving effective tissue repair or replacement. Therefore, modern tissue engineering aims at assisting the re-growth of functional tissues by combining cells and engineering materials with
signaling biomolecules [1,2]. Biomolecular signals that are mainly
growth factors or other cytokines, chemoattractants, adhesion proteins
and many others, must be locally delivered in their active form and
with a sustained release profile. Among the existing technologies, hydrogels – water-swollen, cross-linked polymer networks – have emerged as
particularly promising materials for tissue engineering, as they can act
both as scaffolding materials and/or releasing matrices for biologically
active and cell modulating substances. Their water content, soft nature
and porous structure mimic biological tissues and make them suitable
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Hydrogels for protein delivery in tissue engineering
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⁎ Corresponding author. Tel.: + 39 0737 402215; fax: + 39 0737 402457.
E-mail address: [email protected] (R. Censi).
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to accommodate cells and to encapsulate and release water-soluble compounds like proteins in a controlled fashion.
After highlighting the rationale behind the use of delivery technologies to improve the effectiveness of biomolecular cues, this minireview briefly describes the main hydrogel systems that have been
studied recently as releasing matrices for biomolecules involved in
healing processes. Strategies for encapsulation, (transient) immobilization and controlled release are also discussed. Finally, hydrogelbased releasing technologies with potential clinical impact in tissue
engineering are outlined and an outlook on future directions is
provided.
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2. Proteins in the healing process: an introduction to their
delivery approaches
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Cell fate and behavior is highly influenced by a number of factors
and interactions with the surrounding microenvironment. Signaling
molecules create an effective communication between cells and extracellular matrix (ECM) and orchestrate the complex cascade of events
that leads to successful regeneration of damaged tissues [3]. Of all molecules currently identified as key components of the wound healing
process, growth factors play a pivotal role in the information transfer
mechanism between cell and ECM. Although the natural dynamics of
the cellular microenvironment are difficult to emulate in an artificial
setting, it is currently widely accepted that the self-healing capacity of
an organ or tissue can be augmented by the integration of growth factors. In fact, they accelerate the proliferation and differentiation of
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0168-3659/$ – see front matter © 2012 Published by Elsevier B.V.
doi:10.1016/j.jconrel.2012.03.002
Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
j.jconrel.2012.03.002
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Physical encapsulation of growth factors in order to obtain a controlled release of the active is a frequently applied technique for local
growth factor delivery in tissue engineering. Its relatively simplicity
represents a clear advantage over more sophisticated encapsulation/
release methods and the majority of hydrogel systems relies on this
loading process. Loading is done by incubation of the preformed gel
with the protein of interest or by adding the protein to the hydrogel
forming monomers/prepolymers. In recent years, attention is shifted
to in situ gelling systems which preferably are liquid before administration and gellify at the site of injection [42]. For these gels, loading
with proteins can easily be established by dissolution of the biotherapeutic in the gel forming polymer solution before administration.
Typically, the mechanisms that govern the drug release from these
hydrogels are controlled by diffusion, swelling, erosion, external
stimuli or combinations thereof [43].
The release mechanism depends on both the characteristics of the
polymeric network and the protein. When the hydrogel pores are bigger than the hydrodynamic radius of the protein, diffusion is the driving force for release, with a diffusion rate depending on the protein
size and the water-content of the gel (‘free-volume’) [44]. On the
other hand, when the hydrogels pores are smaller than the protein diameter, swelling or erosion/degradation (bulk or surface) is needed
for release. Generally speaking, the majority of the gel matrices
reported to date exhibit diffusion controlled release, following Higuchi's kinetics, implying that the release is proportional to the square
root of time [45]. This release profile was demonstrated particularly
beneficial for the delivery of several growth factors for tissue engineering applications. For example, Brown et al. and Boerckel et al.
demonstrated that bone regeneration was enhanced when rhBMP-2
was released down a concentration gradient [46,47]. The diffusional
spatiotemporal growth factor presentation proved effective not only
for bone regeneration, but also for other engineered tissues like
blood vessels [20,21].
The Ritger–Peppas equation is often used to fit release data and determine the underlying release mechanism: M t/M∞ = ktn with M t/M∞
the fractional drug release at time, t and k a constant incorporating
Table 1
Typical examples of growth factors and their applications in the field of tissue engineering [11,14].
Growth factor
Abbreviation
Action
Tissue engineering applications
t1:4
t1:5
Vascular endothelial growth factor
Insulin-like growth factor
VEGF
IGF-1
Migration, proliferation and survival of endothelial cells
Proliferation, apoptosis inhibition
t1:6
t1:7
Hepatocyte growth factor
Epithelial growth factor
HGF
EGF
t1:8
t1:9
t1:10
NGF
BMP-2/3/7
PDGF-AA/BB/AB
t1:11
Nerve growth factor
Bone morphogenetic protein
Platelet derived growth factor (2 polypeptide
chains A and B forming homodimers
AA and BB and heterodimer AB)
Transforming growth factor
Proliferation, migration, differentiation of mesenchymal stem cells
Proliferation and differentiation of fibroblast, epithelial,
mesenchymal and glial cells
Proliferation and survival of neural cells
Differentiation and migration of osteoblasts, renal development
Embryonic development, proliferation and migration of
endhotelial and smooth muscle cells
Blood and lymphatic vessels
Cartilage, skin, nerve, kidney,
bone, muscle
Liver, muscle, bone
Skin, nerve
t1:12
Angiopoietin fibroblast growth factor
Ang-1/Ang-2/bFGF
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t1:2
t1:3
TGF-α/β
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3.1. Physical entrapment of growth factors into hydrogels: general 136
principles and release mechanisms
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Proteins can be loaded into hydrogels through a manifold of mechanisms and strategies. They can be physically entrapped in hydrogels or
adsorbed to the matrix by specific interactions, non-covalent or
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3. Protein release strategies from hydrogels
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covalent binding via degradable linkers and subsequently released via
diffusion, swelling, erosion, degradation or a specific trigger such as
pH, temperature, etc. Also, a combination of multiple delivery systems
can be used to achieve modular release of multiple proteins. Depending
on the release mechanism, different release profiles can be obtained.
Table 2 overviews the technologies further described in the following
sections.
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recruited or implanted cells and promote the re-growth of tissues and
organs not capable of self-healing otherwise [4–6].
Growth factors are large polypeptides that modulate cellular proliferation, differentiation, migration, adhesion and gene expression upon
binding to specific receptors on the surface of target cells. Examples of
growth factors and their applications in tissue engineering are listed
in Table 1. Unlike other proteins (i.e. hormones), growth factors act locally and possess short diffusion distances through the ECMs, owing to
their short half-lives. Typically, growth factors are enzymatically or
chemically degraded or deactivated in physiological conditions in a
very limited time frame. To mention, basic fibroblast growth factor
(bFGF) has a half-life of only 3 min after intravenous administration;
[7] similarly, less than 30 min are needed to halve the plasma concentration of vascular endothelial growth factor (VEGF) [8] and plateletderived growth factor (PDGF), isolated from platelets, has a half life of
less than 2 min when injected intravenously [9].
Therefore, the exogenous administration of growth factors for tissue
engineering faces important limitations associated to their rapid inactivation and/or elimination after intravenous delivery, their poor transdermal adsorption due to the large size and the hydrolytic and
proteolytic degradation upon oral administration. Hydrogels are frequently used to release growth factors in a controlled and effective
manner and to direct the protein specifically to the wound site
[10,11]. Spatiotemporal control over the delivery of these key signaling
molecules not only enhances tissue regeneration, but also prevents
unwanted and potentially harmful side-effects at other locations than
the target. Various strategies have been investigated to achieve controlled protein delivery from hydrogels. They comprise ‘direct’ and ‘indirect’ delivery approaches. Direct release occurs through physical
encapsulation, non-covalent binding, covalent immobilization to the
delivery system via hydrolytically or enzymatically degradable linkers
and the use of double carriers, where protein loaded micro/nanospheres are embedded in hydrogels and the release of the active is
based on a combination of mechanisms (diffusion and or degradation).
‘Indirect’ approaches rely on gene therapy and cell transplantation.
Gene therapy is realized through the expression of genetic material
encoding for the desired protein that is delivered in the target tissue,
while in cell transplantation, specific proteins are secreted by cells
that are encapsulated in a hydrogel [12,13].
In the following section the ‘direct’ approaches for protein delivery in tissue engineering will be discussed.
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Proliferation and differentiation of basal, neural, bone-forming
cells, keratinocytes, anti-proliferation epithelial cells
Blood vessels maturation
Proliferation endothelial cells
Nerve, brain, spine
Bone, cartilage, kidney
Bone, skin, muscle, blood vessels
Brain, skin, cartilage, bone
Hearth, muscle, blood vessels,
bone, skin, nerve
Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
j.jconrel.2012.03.002
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Release mechanism from hydrogels
t2:4
t2:5
Some examples
✓
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Fickian diffusion
Mt/M∞ = k √t
Concentration gradient
Generally limited to relatively short time spans
Protein size b mesh size
Geometry determines release rate
Swelling degree controls release rate
Water uptake allows protein to diffuse out
Swelling can be mediated by degradation
PEG/p(HPMAmlac) [15–17]
PLA-PEG-PLA [18]
RADA 16 networks [19]
alginate [20,21]
dex-VS/PEG-4-SH [22]
✓
✓
✓
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Controlled by physical or chemical erosion
Surface erosion mostly for physical gels
Constant release rate
Hydrolytical or enzymatic degradation
Protein size > mesh size
dextran/P(D,L)LA [23]
oppositely charged dextran microspheres [24]
PLGA-PEG-PLGA [25,26] PEG/chol-PEG/β-CD [27].
PEG-P(D,L)LA [28]
✓ Stimuli lead to gel changes and drug release
✓ Stimuli: temperature, pH, biomolecules, drugs, magnetic field, etc…
nitrilotriacetic acid/GyrB [29]
PEG-VS-MMP [30]
PEG acrylate-MMP [31]
✓ Adsorption of proteins through weak interaction (i.e. electrostatic, hydrophobic, H-bonding)
✓ Affinity with binding sites: heparin, antibodies, chelating ions, peptides,
molecular imprinting
Fibrin-heparin [32,33]
PEG-IDAA/Ni+ 2Cu+ 2 [34]
thiol acrylate PEG [35]
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t2:6
Characteristics
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t2:2
t2:3
Table 2
Overview of the main release mechanisms and strategies to load and modulate the kinetics of proteins released from hydrogels.
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t2:10
✓ Combination of multiple carriers: i.e. protein-loaded micro/nanoparticles
embedded in hydrogels
✓ Release of multiple proteins through combination of mechanisms and rates
PEG-GMA/gelatin [39]
Fibrin-PLGA [40]
HAMC/PLGA [41]
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structural and geometric characteristics of the device. If n = 0.5, the release is governed by Fickian diffusion. If n = 1, molecules are released by
surface erosion, while both mechanisms play a role if n has a value between 0.5 and 1 [48,49].
Swelling-controlled systems depend on water uptake and changes in
drug diffusivity within the matrix. Swelling increases polymer flexibility
and makes pores bigger, resulting in higher drug mobility with sometimes n values that exceed 1. As a consequence, drug release depends
on Fickian diffusion, polymer disentanglement and dissolution in water
[50].
Erosion-controlled systems have increased in number since the
development of synthetic biodegradable polymers. In these systems,
the mobility of the drug in the homogeneous non-degraded polymer
matrix is limited and drug release is then governed by the degradation rate of the polymer, porosity increase, and drug diffusion mainly
in the surface of the polymer matrix.
The possibility to tailor the release kinetics of proteins from hydrogels
can be achieved through the use of excipients or by changing the crosslinking density of the polymer network. In this particular instance, the
use of synthetic polymers is extremely advantageous because they offer
the opportunity to fine tune their chemical structure to achieve modular
release. However, also natural polymer networks can be tailored to some
extent by changing polymer concentration and crosslink density. Also extensive work on the development of hybrid networks in which natural
polymers are synthetically modified or combined with synthetic polymers, has been proposed to combine beneficial properties of both kinds
of polymers with respect to mechanical properties, tailorability in terms
of biodegradability as well as tissue and cytocompatibility.
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polyacrylamide [36,37]
PEG-RGD-MMP [38]
t2:9
✓ Attachment of drugs to matrix via covalent bonds
✓ Cleavable linker
✓ Hydrolytical or enzymatic cleavage
The geometry of the hydrogel-based depot also affects the release rate
and the duration. Generally, the bigger is the device, the longer the diffusion distances become and the longer the release consequently lasts. In
recent years, patterning techniques utilized in tissue engineering (i.e.
stereolithography, bioprinting) have been used to adjust surface area
and consequently modulate release profiles of the biomolecules [51–53].
Deviation from the described release behaviors is observed when a
specific stimulus – typically temperature, pH, presence of certain molecules – leads to a physical or chemical change in the network structure,
for instance, hydrogels swell or shrink in response to a certain trigger
thereby modulating the release of encapsulated drugs/proteins.
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3.1.1. Synthetic polymers for the physical encapsulation of growth factors
Among the various synthetic polymers studied for the delivery of
growth factors in tissue repair, PEG-based networks play a prominent
role.
Burdick et al. investigated the feasibility of a PEG-based hydrogel
platform for the replacement and delivery of neurotrophins to the central nervous system. For this purpose, triblock copolymers based on
PEG and acrylated poly(lactic acid) (PLA-b-PEG-b-PLA) were synthesized
and hydrogels formed by photopolymerization. Three neurotrophins, ciliary neurotrophic factor (CNTF), brain-derived neurotrophic factor
(BDNF) and neurotrophin-3 (NT-3) were encapsulated and released in
vitro mainly by diffusion in 20 to 80 days depending on polymer concentration (10 to 30%) and neurotrophin size. In vitro, the released proteins
stimulated the proliferation of a TF-1 cells transfected with the α-subunit
of the CNTF receptor and stimulated outgrowth of a greater number of
neuritis from retinal explants than in control culture conditions [18].
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Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
j.jconrel.2012.03.002
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3.1.2.1. Polysaccharides. Polysaccharides are in general hydrophilic
polymers and are therefore very suitable for the design of hydrogels.
The most commonly used polysaccharides in recent hydrogel research
aimed for protein delivery are alginate, hyaluronic acid, chitosan and
dextran (Fig. 1) [73,74]. Representatives of the hydrogel technologies
based on these polysaccharides are briefly discussed below with respect
to their use for growth factor delivery.
3.1.2.1.1. Alginate. Alginate is a linear polysaccharide composed of
homopolymeric blocks of 1-4′-linked β-D-mannuronic acid (M) and
its C-5 epimer α-L-guluronic acid (G), respectively, covalently linked
in different sequences or blocks. This polymer block consists of consecutive G-residues, consecutive M-residues or alternating M and G
residues. Depending on the block composition, alginate has different
conformational preferences and behavior. G-rich blocks of the polymer are able to bind divalent cations of which Ca 2+ is frequently
used for crosslinking alginate to obtain hydrogels. Ca 2+ ions bind to
the G units of different polymers chains cooperatively in a so-called
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3.1.2. Natural polymers for the physical encapsulation of growth factors
Hydrogels designed with natural polymers as building blocks display
multiple advantages over synthetic polymer networks with respect to
their biocompatibility, biodegradability and good cell adhesion properties. Therefore, extensive work is available on biopolymer-based hydrogels for cell and growth factor encapsulation in regenerative medicine
[69–72].
The main classes of natural polymers studied in hydrogel formulations are polysaccharides and proteins/polypeptides.
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sensing hydrogel based on gyrase sub-unit B (GyrB), reversibly crosslinked by coumermycin and able to release VEGF upon addition of an aminocoumarin antibiotic, novobiocin. The polymer forming the hydrogel is
based on polyacrylamide functionalized with nitrilotriacetic acid chelating a Ni2+ ion to which GyrB can bind through a hexahistidine sequence.
The addition of novobicin causes the cross-links to be partly broken,
resulting in opening of the network structure and release of VEGF [29].
Hubbell et al. synthesized hydrogels comprising multiarm vinyl
sulfone-terminated PEG, a monocysteine containing adhesion protein, and a bis(cysteine) metalloprotein substrate protein (MMP).
These hydrogels were studied for tissue engineering purposes and
both cells and vascular endothelial growth factor (VEGF105) were encapsulated in the hydrogel. The release of the encapsulated factor is
triggered by matrix metalloproteinases (MMPs) [30]. MMPs are are
zinc-dependent endopeptidases capable of degrading extracellular
matrix proteins and they influence cell behavior such as cell proliferation, migration (adhesion/dispersion), differentiation, angiogenesis,
apoptosis, and host defense. MMPs sensing hydrogels were developed with the aim to mimic the natural processes involved in tissue
repair. Other examples of semi-synthetic hydrogels were described
by Phelps et al. whose design provides incorporated VEGF, enzymedegradable sites and arginine-glycine-aspartic acid (RGD) cell adhesive ligands. Similarly to the system proposed by Hubbell, these matrices use PEG as the main building block because it has been shown
to act as a protein-resistant and cell non-adherent background. The
network is based on PEG diacrylate which is photocrosslinked in the
presence of MMP-degradable polypeptide functionalized with two
PEG-acrylate groups, a mono-PEG-acrylate RGD and a mono-PEGacrylate VEGF. Recently, Phelps et al. have demonstrated the benefits
of this approach in a mouse hind limb femoral artery ligation model.
After 7 days, mice treated with implanted hydrogels showed a 50% increase in perfusion to the limbs and 100% increase in perfusion to the
feet as compared to untreated mice. The matrix alone (not containing
VEGF) performed as well as soluble VEGF injections, indicating that
the engineered degradable and adhesive hydrogel has itself a beneficial healing or supportive effect. Engineered matrix containing VEGF
performed better than injections of soluble VEGF, acting synergistically as a directive scaffold and a growth factor delivery vehicle [31].
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These types of hydrogels, initially described by Hubbell and coworkers
[54], have been extensively investigated for growth factor delivery
[55,56] because they offer several benefits, including the possibility to
easily modify the network properties by changing the macromer chemistry and solution concentration. For instance, the degradation and swelling of the hydrogels are dictated by the number of lactic acid repeat units,
the acrylation efficiency, the molecular weight of the PEG core and the
concentration of macromer in the prepolymer solution [57]. Also, the
type of degradable unit (e.g., lactic vs caproic acid) alters network degradation rates [58].
Similar release and network concepts apply to other PEG/polyesters
hydrogels, like those based on PEG-poly(lactic-co-glycolic acid)
(PLGA)-PEG triblock copolymers [59]. These materials display lower
critical solution temperature behavior and were processable avoiding
the use of high temperatures to dissolve the polymer. It was shown
that upon subcutaneous injection (rat model) the hydrogels were stable
for one month [60]. TGF-β1 was loaded into these hydrogels and used as
a slow releasing drug reservoir aimed for wound healing purposes. Significant levels of re-epithelialization, cell proliferation and collagen organization were observed [61]. Porcine growth hormone (pGH) and
Zn-pGH were sustainably released from PLGA-PEG-PLGA hydrogels for
10–14 days in vitro with no initial burst and a good correlation with in
vivo data, that confirmed a weight gain in hypophysectomized rats
equivalent to conventional daily injections of 5 mg pGH for 14 consecutive days. Approximately 85% of the glycosylated granulocyte colony
stimulating factor loaded in PLGA-PEG-PLGA hydrogels was released
over a 12-day period and similarly to pGH, showed similar efficacy in
rats as compare to daily i.v. injections [62]. Censi et al. studied the protein release from photopolymerized thermosensitive PEG-based networks formed upon gelation of methacrylated poly(hydroxypropyl
methacrylamide lactate)-PEG- poly(hydroxypropyl methacrylamide
lactate) (p(HPMAm-lac)-PEG-p(HPMAm-lac)) triblock copolymers
[63]. Proteins are released from this network by Fickian diffusion and,
similarly to previously described systems, network properties and release behavior can be tailored by tuning polymer chemistry and concentration [16,17,64]. The gels showed a good biocompatibility in rats [65]
and their potential for cartilage tissue engineering application has been
demonstrated [51].
Peptide based hydrogels were also used for protein delivery and release. A class of ionic self-complementary oligopeptides named RAD 16
and commercialized as PuraMatrix was introduced by Zhang et al. These
peptides consist of alternating hydrophilic and hydrophobic amino
acids that form β-sheet structures having one polar surface with complementary charged ionic side chains and a non-polar surface with alanines. These peptides are able to spontaneously self-assemble into
stable macroscopic matrices or nanofibers. The monovalent cations
that are needed for the peptide assembly are sequestered from the
physiological environment [19]. The described family of peptides has
been used to encapsulate and deliver several proteins for intramyocardial delivery. This peptide-based hydrogel has also been used to deliver
platelet-derived growth factor BB (PDGF-BB) [66], stromal cell-derived
factor-1 (SDF-1) [67] and insulin-like growth factor I (IGF-I) [68] to decrease myocardial infarct. The observed slow and controlled release of
the active proteins has been ascribed to diffusion and to some extent
to interaction between protein and polymer. The amphiphilic nature
of the self-assembling polypeptide leads to interactions with the loaded
protein, resulting in reduced diffusivity and slower release kinetics.
Hydrogels that showed degradation mediated release of model
proteins and that potentially can be used to physically encapsulate
and release growth factors with a zero order release kinetics are for
example those formed by stereocomplexation between star PEG copolymerized with the stereoforms D or L of PLA [28] or those based on
PLGA-PEG-PLGA (ReGel) systems [25,26] and inclusion complexes of
PEG/cholesterol-PEG/β-cyclodextrin [27].
Stimuli sensing hydrogels have also been investigated for growth factor release in tissue engineering. Recently, Weber et al. proposed a drug-
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protein or plasmid expressing SDF-1 and the kinetics of SDF-1 release
were measured both in vitro and in vivo in mice. They demonstrated
that SDF-1 plasmid- and protein-loaded patches were able to release
therapeutic product over hours to days, with faster release for SDF-1
protein (in vivo K(d) 0.55 days) than SDF-1 plasmid (in vivo K(d)
3.67 days). Prolonged (induction of) SDF-1 release of SDF-1 resulted
in accelerated healing (9 days) and reduced scarring of acute surgical
wounds in Yorkshire pigs [77].
Although the biological properties of alginate are well recognized,
some limitations in tailoring the mechanical properties, the pore size
and the release kinetics of proteins from these gels still remain. Many
researchers investigated several modified alginates to overcome these
limitations [78–83]. For example, Ruvinov et al. examined the release
of hepatocyte growth factor (HGF) from modified alginate hydrogels
and studied its potential to induce angiogenesis in vivo. They developed
a cross-linkable affinity-binding alginate by introducing sulfoester
groups into the uronic acids of alginate. It was shown that alginatesulfate was capable to bind heparin-binding proteins, like growth factors, with equilibrium binding constants similar to that of heparin. The
affinity binding to alginate-sulfate retarded the release of HGFThe bioconjugate HGF-alginate-sulfate was combined with alginate in an aqueous solution and a hydrogel was formed ex vivo by partial cross-linking
in presence of CaCl2. Upon injection into ischemic myocardial tissue the
injectable hydrogel completed its cross-linking by sequestration of
endogeneous Ca 2+. A cumulative protein release of 75% was observed
after 6 h and in the following 5 days a cumulative release of approximately 85% was achieved [84]. This profile is typical of diffusion-based
delivery systems that release the drug down a concentration gradient,
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egg-box arrangement. The partial oxidation of alginate using sodium
periodate makes the polysaccharide biodegradable [75,76], and these
oxidized alginates are therefore particularly appealing for biomedical
applications.
An alginate hydrogel patch to deliver stromal cell-derived factor-1
(SDF-1), a naturally occurring chemokine that is rapidly overexpressed
in response to tissue injury, was described by Rabbany et al. Alginate
patches were loaded with either purified recombinant human SDF-1
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Fig. 1. Most commonly used polysaccharides for hydrogel preparation for biomedical
applications (M = mannuronic acid, G = guluronic acid).
Fig. 2. Biophysical properties and vascular endothelial growth factor (VEGF) release from alginate gels. Degradation of gels formed from high molecular weight non-oxidized alginate (triangles), binary molecular weight non-oxidized alginate (white squares) and binary molecular weight partially oxidized alginate (black squares) in phosphate-buffered saline (A). The viscosities of solutions of high (white bar), low (black bar) and binary (grey bar) molecular weight partially oxidized alginate were assessed (B). Release kinetics of
VEGF165 from gels formed from binary molecular weight alginate, either partially oxidized (black squares) or non-oxidized (white squares) (C). Greater endothelial cell proliferation was observed in cells cultured with VEGF released from binary molecular weight partially oxidized alginate gels (grey), as compared to medium without VEGF (black) and
control VEGF added to medium (white), as determined from the cell number counts from day 0 to day 4 (D). Reproduced from ref. [85].
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and subsequent physical or chemical crosslinking [90]. For example,
Hennink et al. developed a hydrogel system potentially suitable for the
delivery of proteins relevant for tissue engineering, where a dextran
backbone is derivatized with hydroxyethyl methacrylate (HEMA) moieties. This polymer in aqueous medium was chemically cross-linked by
radical polymerization of the dextran-bound methacrylate groups to
yield hydrogel based microspheres [91,92]. These microspheres released
model proteins as well as therapeutic relevant proteins (IL2 [93]; hgH
[94]) in a controlled manner for days to weeks. The same group also developed a physically crosslinked hydrogel system by grafting the dextran
backbone with either D- or L-oligo-lactate chains containing 5 to 15 lactic
acid units. The hydrogel formation was driven by stereocomplexation
and preclinical studies demonstrated the suitability of this hydrogel for
the controlled delivery of IL-2. The density of oligo-lactates and the
grafting density allowed tailoring of the release and degradation rate.
The release of the encapsulated protein was dependent on diffusion
[23]. Further efforts were addressed by Hennink et al. towards the development of dextran based delivery platforms for proteins, potentially suitable for growth factor and cytokine controlled release. Self-assembled
macroscopic hydrogels based on oppositely charged dextran microspheres were prepared and their release behavior investigated. It was
found that native proteins were released by diffusion/swelling [24].
More recently, dextran-peptide bioconjugates were synthesized by
Shoichet et al. In their approach, the polysaccharide was modified
with p-maleimidophenyl isocyanate (PMPI) thereby introducing maleimide functionalities in the backbone (Dex-PMPI). A peptide crosslinker, derived from collagen and susceptible to gelatinase A digestion,
was synthesized with bifunctional cysteine termini and used to crosslink Dex-PMPI (Fig. 4). It was demonstrated that this type of hydrogel
was cytocompatible and mimicked the degradation and remodeling of
the ECM through the activity of cell-secreted enzymes [95].
Michael addition cross-linking reaction was used to induce network
formation between dextran vinyl sulfone conjugates (dex-VS) and tetrafunctional mercapto poly(ethylene glycol) (PEG-4-SH). The release
of several proteins including bFGF from hydrogels of different polymer
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as described for another alginate hydrogel described in Fig. 2 [85]. A recent method employed to make alginate networks more stable and tailorable is through the preparation of interpenetrating polymer
networks (IPN's) [83]. An IPN is defined as a network comprising two
or more polymers that are partially interlaced on a molecular scale
but not chemically bounded to each other and that can be separated
only after chemical bond breakage [86]. IPNs are designed to combine
the advantageous properties of different polymers. This combination
often translates in good modulation of the mechanical and biological
characteristics of the final hydrogel, derived from the additive or synergistic effect of the polymers' properties. Pescosolido et al. developed IPN
hydrogels based on Ca2+-alginate and photocross-linked dextranmethacrylate derivatives, of which the mechanical properties were notably improved by the synergistic effect of the polymers composing the
IPN system. These hydrogels proved effective in the controlled release
of proteins [83]. In another study, it was demonstrated that complete
delivery and a better control over VEGF release kinetics were achieved
by cross-linking alginate microparticles with Zn2+ instead of Ca 2+
[87,88].
Silva et al. described alginate hydrogels that were covalently modified with RGD sequences and loaded with growth factor VEGF, and
that were lyophilized to form microporous scaffolds to induce neovascularization. The delivery system, modified with adhesive sequences
and loaded with growth factor and cells, was used to induce the formation of a depot of vascular progenitor cells (outgrowth endothelial cells
(OECs)) in vivo. It was observed that OECs were viable during the
studied time frame and that the encapsulated VEGF induced the cells
to migrate out of the scaffold. Local and controlled delivery of the morphogen was effective in stimulating the cells to repopulate the damaged
tissue and participate in regeneration of a vascular network, while bolus
injections was ineffective in preventing necrotic toe (Fig. 3) [89].
3.1.2.1.2. Dextran. Dextran consists of α-1,6-linked Dglucopyranoses with some degree of 1,3-branching. Various methods
to cross-link dextran hydrogels have been developed and its high number of available hydroxy groups presents many options for derivatization
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Fig. 3. Analysis of angiogenesis in ischemic hindlimbs after OEC transplantation. (A) Implantation of blank scaffolds, bolus injection of OECs and VEGF (same quantities as placed in
scaffolds), transplantation of OECs on scaffolds lacking VEGF (alginate scaffold OEC), and transplantation of OECs on scaffolds presenting VEGF121 [alginate scaffold (VEGF) OEC].
Photomicrographs of tissue sections from ischemic hindlimbs of SCID mice at postoperative day 15, immunostained for the mouse endothelial cell marker CD-31 (B), and
human CD-31 (C). Reproduced from ref. [89].
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Fig. 4. (A) Reaction of dextran with p-maleimidophenyl isocynate (PMPI) and (B) preparation of dextran-peptide hydrogel through the Michael-type reaction of peptide crosslinker to Dex-PMPI. Reproduced from ref [95].
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concentrations was studied. Also these networks showed diffusional
release and a certain extent of tailorability of the bFGF [22].
In a recent study, Sun and coworkers demonstrated that the manipulation of the cross-link density of dextran based hydrogels loaded with
multiple growth factors affects the neovascularization of ischemic and
wounded tissues. Dextran-allyl isocyanate-ethylamine (Dex-AE)
derivatives with different degrees of substitution were synthesized
and mixed with PEG diacrylate at different ratios and subsequently, hydrogel formation was induced by photopolymerization. It was found
that tissue ingrowth was favored by reducing the degree of substitution
of cross-linking groups, which resulted in reduced rigidity, increased
swelling, increased VEGF release rate, and rapid hydrogel disintegration. Furthermore, the release of multiple angiogenic GFs increased
the size and number of newly formed functional vessels [96]. In another
study, hybrid hydrogels based on combinations of glycidyl methacrylated dextran and gelatin, processed into different physical structures
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(microspheres, cylinders) have been synthesized and used to deliver
growth factors, including BMP-2and IGF-1 [97].
3.1.2.1.3. Hyaluronan. Hyaluronan or hyaluronic Acid (HA) is a
linear glycosaminoglycan composed of repeating disaccharide units
of D-glucuronic acid and N-acetylglucosamine [98,99]. This naturally
occurring polysaccharide is negatively charged under physiological
conditions has molecular weights up to 10 7 Da. It is found mainly in
the extracellular matrix (ECM) and in the synovial fluids of joints
where it acts as a lubricant by which it reduces the friction of bones
due to its unique viscoelastic properties [100].
Because of its biocompatible, attractive physical properties and
possibilities for further chemical modifications, HA is an extensively
studied polysaccharide for the development of delivery systems for
tissue engineering applications [101–107]. However, unmodified
hyaluronic acid in aqueous solution displays poor mechanical properties. Therefore, to be suitable for protein delivery, HA-gels need
Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
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3.1.2.2. Proteins. Collagen, fibrin and gelatin are naturally occurring,
enzymatically degradable proteins that have been extensively
exploited as main component in matrices for drug delivery and tissue
engineering purposes. Many of the hydrogel systems based on gelatin
and collagen are cross-linked using harsh conditions that exploit gluteraldehyde or water-soluble carbodiimides and that can be detrimental for the encapsulated proteins [126]. However, also milder
preparation conditions were investigated. For instance, noncovalently crosslinked fibrillar collagen can be used to create hydrogels by entanglements of collagen fibers [127]. The mesh sizes of
these entangled collagen fibers are quite large and therefore diffusion
of only very large proteins can be controlled. Nevertheless, the protein release rate may still be lower than expected based on diffusion
coefficients due to weak interactions between collagen and loaded
proteins [127].
Gelatin has a long history in clinical use with examples of gelatin
based systems with a variety of physical geometries. Gelatin derived
from collagen can be positively or negatively charged depending
whether collagen depolymerization is carried out at acidic or basic conditions. The charge of gelatin network can be utilized to complex oppositely charged proteins serving as additional feature to control release
rate [128]. Gelatin based hydrogels have been used for the delivery of
FGF-2 [129], HGF and VEGF [130]. Nowadays, gelatins can be produced
in yeast cells recombinantly leading to slightly different polymer properties. The advantage is that properties such as molecular weight, amino
acid sequences and isoelectric points can be tailored precisely. Degradation times, swelling properties and ultimately also drug release kinetics
are affected by the design of recombinant gelatin. Only a few recombinant gelatin hydrogels have been investigated for their protein delivery
properties [131,132]. Fibrin is a fibrous, non-globular protein derived
from fibrinogen, a protein of the blood coagulation cascade. Fibrinogen,
a water soluble protein, is converted into fibrin that polymerizes to
yield a gel due to the action of thrombin. Fibrin networks are gradually
resorbed in vivo by the secretion of fibrinolytic enzymes and as such has
been widely used for formulating depot systems for proangiogenic
growth factor delivery. Commercial fibrin glues are also used in the clinic for in situ application.
In an early clinical trial, VEGF was injected in an in situ forming fibrin matrix into the right popliteal region of a man suffering from
claudication. Although substantial growth of newly formed vessels
was observed, a high initial burst occurred within the first 24 h. To
overcome this issue, VEGF was later covalently bound to the matrix
and released by enzymatic cleavage [133,134]. Similarly, FGF-2 was
electrostatically immobilized in fibrin hydrogels using heparin
[135]. These strategies showed in animal model enhanced neovascularization and reduced fibrosis and inflammation.
Recently, Drinnan et al. reported on an elegant approach on the
multimodal release of different proteins from the same hydrogel.
They developed an injectable PEGylated fibrin gel designed for the release of PDGF-BB and TGF-β1 with distinct kinetics (Fig. 5). Growth
factors were loaded into PEGylated fibrin gels via 3 mechanisms: entrapment, conjugation through a homobifunctional amine reactive
PEG linker, and physical adsorption on the fibrin matrix. PDGF-BB
was entrapped during thrombin-mediated crosslinking leading to its
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[116–118]. The interactions between the nitrene groups and the
amine groups of the proteins were identified as the reason for the incomplete release from these types of hydrogel [119,120]. Yeo et al.
reported that 80% of initially loaded VEGF was retained in the network
after cross-linking [118]. In situ gelling double networks of oxidized
dextrans and thiolated chitosan (semi-interpenetrating networks)
were recently proposed for healing purposes by Chen et al. These networks showed faster gelation kinetics as compared to auto-gelling thiolated chitosan hydrogels [121]. Recently, comprehensive and complete
reviews on chitosan hydrogels were published concerning their drug/
protein delivery aspects [122–125].
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physical and/or chemical stabilization. Usually, the carboxylic groups
of hyaluronic acid are derivatized to introduce functionalities that
make the polymer suitable for cross-linking [108–113].
In situ photopolymerization was applied by Park et al. to crosslink
thermosensitive hyaluronic acid/Pluronic composite hydrogels that released human growth hormone with kinetics correlated with the mass
erosion [105]. Subsequently, Patterson et al. synthesized glycidyl methacrylate modified hyaluronan hydrogels with different crosslink densities and thus of different degradation rates and used them to prepare
photopolymerized hydrogels as BMP-2 and/or VEGF releasing matrices
for bone regeneration. They demonstrated that all matrices displayed
diffusion-like release behavior of entrapped proteins and their degradation and release rates modulated the formation of mature bone, specifically affecting the organization of the collagen matrix. Additionally, the
co-delivery of an angiogenic molecule (VEGF) in conjunction with an
osteoinductive molecule (BMP-2) increased the extent of formed mineralized tissue [102].
Crosslinked, PEG diacrylate/thiolated hyaluronan hydrogels were
studied for the delivery of multiple growth factors (VEGF and/or
Ang-1) in both presence or absence of heparin, a molecule that forms
complexes and stabilize the structure and modulates the release of
growth factors, according to the dissociation constant of the complex
growth factor/heparin. Greater neovascularization was observed
when the hydrogels were loaded with both growth factors [106,113].
Similar studies were conducted by the same group both in vitro and
in vivo using in situ gelled thiolated gelatin/thiolated hyaluronic acid/
PEG diacrylate based networks with or without disulfide bridged heparin as part of the network and releasing combinations of growth factors like VEGF, Ang-1, PDGF and KGF. It was found that release
followed first order kinetics, with gelatin slightly increasing the
growth factor release rate, and heparin, slowing down the kinetics, because of its complexation with the growth factors. The extent of revascularization and blood vessel maturation was found greater in gels
containing both heparin and gelatin than in those containing only gelatin or only heparin [104,114].
Other cross-linked hyaluronic acid networks developed for tissue
engineering purposes are for example those studied by Varghese et
al. based on aldehyde-HA/hydrazide-HA for the burst-free release of
rhBMP-2 [107] or those developed by Leach et al. composed of
photocross-linked glycidyl HA or photocross-linked glycidyl HA/
acrylate 4-arms PEG [108]. Other HA cross-linkable hydrogels for
protein delivery that have been proposed recently are based on
tyramine-HA [109], on disulfide bridged thiolated-HA [111], divinyl
sulfone-HA/divinyl sulfone PEG [110] and adipic acid dihydrazide-HA/
methacrylate-HA [112].
Since HA is negatively charged at physiological pH, the protein release rate will be affected by the charge of the protein. Generally
speaking, cationic proteins are slower released than anionic or neutral proteins. The release of proteins is not only influenced by charge
interactions with the polymer matrix, but also by the enzymatic degradation of HA-based gels in the presence of hyaluronidase, which is
present in biological tissues [109,112].
3.1.2.1.4. Chitosan. Chitosan is a copolymer of glucosamine and Nacetylglucosamine and is derived from the natural polymer chitin
(=poly N-acetylglucosamine) by (partial) deacetylation. The polymer
is positively charged at low pH's and uncharged and insoluble at neutral
and high pH's. Chitosan has mucoadhesive properties originating from
its cationic nature [92]. To make this polymer suitable for hydrogel formation, water-soluble and cross-linkable derivatives were synthesized.
One of these approaches was realized by grafting 4-azidobenzoic acid to
the available free amine groups of lactose-modified chitosan. In the
presence of UV light the azido groups are converted into nitrenes that
are highly reactive species with the free amine groups on chitosan
yielding networks [115]. VEGF and FGF-2 were loaded into photo
cross-linkable modified chitosan hydrogels and their release and efficacy in inducing neovascularization in ischemia models were studied
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Fig. 5. Gelation mechanism and protein encapsulation methods for PEGylated fibrin gels. Reproduced from ref. [136].
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as compared to fibrin hydrogels. In a similar approach adopted by Kiick
et al. four arms thiolated PEGs were reacted with monomaleimidefunctionalized low molecular weight heparin and cross-linked in aqueous medium after addition of heparin-binding peptide modified PEGs.
This delivery system as well other similar hydrogels have been extensively used for the controlled delivery of many growth factors for different applications [145–148].
Heparinization was applied also to collagen [149,150], alginate
[151], hyaluronan [152] and pluronic [153] using different chemistries and cross-link methods.
Polymers were complexed with growth factors binding heparin also
by electrostatic interactions. Sulfonation of hyaluronan and uronic acid
monomers of alginate allowed non-covalent interaction with heparinbinding proteins [80,84]. Although heparin bound hydrogels are definitely the most studied affinity systems to date, some challenges still remain such as unstable protein-heparin bonding in vivo, large amounts
of heparin necessary to achieve sufficient drug loading and heparininduced thrombocytopenia [154]. Affinity hydrogel systems based on
the mechanism of metal-ion-chelation and binding of histidine tagged
proteins have been proposed as an effective alternative to the
heparin-containing systems. To mention, Metters et al. developed
PEG-co-methacrylated iminoadiacetic acid displaying affinity binding
for histidine tagged proteins through divalent metal ions such as nickel
and copper. The authors also elaborated a mathematical model to predict the release rate of proteins from the hydrogels [34].
Phage display is a technique that has been exploited to identify peptides with affinities for growth factors that could be coupled to the gelling polymer [155].
As an example Anseth et al. developed an affinity peptidefunctionalized PEG hydrogel with the ability to sequester monocyte chemotactic protein 1 (MCP-1), a chemokine that induces the chemotaxis of
monocytes, dendritic cells, and memory T-cells. Affinity peptides were
immobilized in PEG hydrogels via a thiol-acrylate photopolymerization.
The release of encapsulated recombinant MCP-1 from PEG hydrogels
was tailored by altering the spacer distance between the affinity peptide
and the crosslinking site [35]. Antibodies were also exploited for the specific binding of proteins onto scaffolds [156].
Molecular imprinting is a technique used to synthesize biomimetic
polymer networks with template-shaped cavities that have affinity for
specific molecules of interest (the templates). The nature of the interaction can vary from non-covalent bonds to metal coordination and even
covalent bonds. This technique has been applied for drug and protein
delivery and recognition. However, although some success has been
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diffusion-controlled release over 2 days. TGF-β1 was both conjugated
through the PEG linker and bound to the matrix via physical adsorption, delaying the release rate of TGF-β1 up to 10 days. Further, the
release rate was highly correlated to gel degradation rate, indicating
that TGF-β1 release was degradation-controlled [136].
Fibrin hydrogels were also used to fabricate circuits for the controlled release of chondroitinase ABC (a bacterial enzyme capable of
digesting the glycosaminoglycan (GAG) side chains from chondroitin
sulphate proteoglycans) in spinal cord lesion treatment. Sustained delivery of the protein in vivo for about 3 weeks was demonstrated as
compared to bolus injection. Furthermore, after 3 weeks, six times
more bioactive chondroitinase ABC and 37% lower levels of inhibitory
glycosaminoglycans (GAG) were found in the spinal cord when hydrogels were used vs intraspinal injection of enzyme solution [137].
3.1.2.2.1. Affinity-based drug release. Affinity based drug delivery
systems utilize physical interactions between the therapeutic drug
and the delivery system to manipulate drug loading and control release
kinetics.
The easiest method to modulate release kinetics of loaded proteins is
physical adsorption, involving the formation of ionic complexes by electrostatic interactions between charged polymers oppositely charged proteins [138–140]. Many charged polymers like hyaluronan, chitosan,
alginate, acidic gelatin can be potentially used to retard drug release [89].
Heparin sulfate is a naturally occurring, highly sulfated anionic
glycosaminoglycan found in the ECM that is a natural matrix responsible for immobilizing and releasing various proteins that influence the
natural processes of adhesion, migration, proliferation, differentiation.
This glycosaminoglycan possesses a specific binding domain for many
growth factors that interact with heparin via non-covalent bonding
[32,141]. Besides enhancing matrix ability to retain the encapsulated
protein, it has been shown that affinity binding with heparin also stabilizes the drug, preserving its structure and functionality during handling
of the material [142]. Heparin is generally incorporated within the delivery system through many strategies as addressed in recent state of
art reviews of affinity-based drug delivery [11,143,144]. Therefore,
this paper deals only with some general aspects of this topic.
One approach to incorporate heparin in a hydrogel structure is to covalently link the heparin moiety to the polymer. Many examples of
growth factor delivery from heparin modified polymer hydrogels have
been reported. Hubbell used a bi-domain peptide that was covalently
attached to fibrin on one site and to heparin on the other extremity.
Growth factors of different affinity for heparin (bFGF >bNGF, BDNF,
neutrophin-3) [32,33] were loaded and released with slower kinetics
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Proteins, mostly exploiting their reactive amine and thiol groups,
can be covalently bound to polymer matrices via functional groups
like hydroxy, amine, carboxyl groups that, if not naturally present in
the structure of the polymer, have to be introduced by functionalization reactions, blending or co-polymerization. The release of the
protein is mediated either via hydrolysis or reduction reactions or
(cell-mediated) enzymatic cleavage. This type of mechanism leads to
on-demand release of loaded proteins, mimicking the enzymatic activity naturally occurring in ‘healthy’ ECM. However, stability and maintenance of biological activity of the protein may represent an issue.
Verheyen et al. recently reported on an efficient strategy to introduce methacrylamide groups on the lysine residues of a model protein
(lysozyme) for immobilization and triggered release from a polyacrylamide or dextran hydrogel network. They designed a novel spacer
unit containing a disulfide bond, such that the release of the protein
can be triggered by reduction [36,37].
PEG hydrogels containing pendant RGD peptides can be formed
using amine-functionalized MMP ligands as a cross-linker. Close proximity of the gel to cells was mediated by RGD peptide and BMP-2
bound to the matrix was delivered to the site of the bone defect [38].
Other popular systems that rely on covalent cross-linking are those containing peptide linkers susceptible to the enzymatic activity [158–160].
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3.3. Dual/multiple delivery systems
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Biomedical hydrogels that can deliver multiple growth factors in a
multimodal mode and provide desirable pore structure and porosity
to potentially encapsulate cells, have considerable potential as future
therapeutic tools in tissue engineering. The use of growth factors
loaded microspheres embedded in the hydrogel structure is a common approach to multimodal protein delivery.
Microparticles of acidic and basic gelatin were used as carriers for
the individual delivery of two different growth factors and embedded
in a hydrogel matrix, composed of glycidyl methacrylated dextran
(Dex-GMA)/gelatin. This hybrid system allowed the independent release of BMP-2 or IGF-1(Fig. 6) that facilitated cell attachment, proliferation, metabolism, and osteoblastic differentiation of cells in a
synergistic manner [39,161].
Fibrin hydrogels were complexed with heparin-functionalized PLGA
nanoparticles for the release of VEGF for revascularization purposes
[40].
A similar approach was adopted by Burdick et al., who formulated
PLGA microspheres into PLA-PEG-PLA hydrogels for the delivery of
multiple neutrophins with individual release rate [18]. Similarly,
Biondi et al. used PLGA microspheres in collagen gels and collagenHA semi-interpenetrating networks. The protein release kinetics and
delivery onset strongly depended on the complex interplay between
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The importance of growth and adhesive factors and their synergistic interplay during healthy tissue development is nowadays a wellestablished concept. However, their administration in the form of
bolus injection to promote tissue regeneration and repair has shown
to be often ineffective and potentially harmful due to the short duration of action of the encapsulated protein drug. Such a delivery approach of growth factors has been applied in a number of clinical
trials of unsatisfactory outcome. While some phase I clinical trials
(angiogenic gene therapy and human growth factor FGF-I infusion
for coronary artery disease) [165,166] have reported promising results, phase II clinical trials, recruiting a larger number of patients,
have not shown the expected benefits to patients. To mention, VIVA
(Vascular endothelial growth factor in Ischemia for Vascular Angiogenesis) trial conducted on 178 patients by intravenous and intracoronary infusion of VEGF [167] and FIRST trial comprising intracoronary
infusions of FGF-2 on 337 patients [168], did not result in effective
commercially available formulations for the treatment of cardiovascular diseases. On the other hand, positive outcomes resulted from
clinical trials on growth factors administered through the use of delivery systems that controlled to some extent the protein release
rate. In this way the problems associated with the applications of
bolus injections are overcome. For example, after successful clinical
trials, two growth factors BMP-2 (clinical trial BESTT) [169,170] and
BMP-7 (clinical trial OP-1 Putty) [171,172] immobilized in collagen
sponge and matrix, respectively, are currently commercially available
for the treatment of bone fractures and defects. Also in the field of
cardiovascular diseases, encapsulation of FGF-2 into alginate microcapsules led to some positive results in clinical trial phases (clinical
trial Polymer) [173,174]. Finally, Regranex® Gel is the only growth
factor delivery system of PDGF that obtained full FDA approval for
clinical use.
Generally speaking, formulation of proteins in controlled delivery
systems allow to sustain the release and lower the doses of potent
growth factors, that, if administered by infusion, need supraphysiological concentrations, leading to severe side-effects and do not assure
therapeutic efficacy over a sufficient time-span, because of rapid degradation/elimination. Moreover, the design of smart delivery systems
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4. Conclusions and outlook
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3.2. Covalent binding
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protein transport through the PLGA matrix and in the collagenbased release media, and water sequestration within the scaffold
[162,163]. Physical hydrogel blends composed of hyaluronan (HA)
and methyl cellulose (MC) were designed by Shoichet et al. for independent delivery of one or more therapeutic proteins, from 1 to
28 days, for the ultimate application of spinal cord injury repair. To
achieve a diversity of release profiles, they exploited the combination
of fast diffusion-controlled release of dissolved solutes from the
HAMC itself and slow drug release from PLGA particles dispersed
within the gel [41]. Methacrylated hyaluronan networks encapsulating alginate microparticles enhanced mesenchymal stem cell chondrogenesis following delivery of TGF-β3 in vitro and in vivo [164].
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achieved in the molecular imprinting of small molecule drugs, huge
challenges still exist with imprinting of proteins [42,77,125,157].
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Fig. 6. Cumulative protein(s) release of BMP-2 and/or IGF-1 from scaffolds containing BMP-2/IGF-1 combination (BMP-2 and IGF-1) (C); or from scaffolds containing a mixture of
microparticles loaded with BMP-2 or IGF-1 (F). Reproduced from ref. [39].
Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
j.jconrel.2012.03.002
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This project was financially supported by L'Òreal – Unesco, “For
Women in Science 2011”.
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may offer the possibility to deliver multiple growth factors with independent release rates and to load or attract cells for the regeneration
of damaged tissues. This review summarizes the efforts that have
been addressed on facilitating the delivery of single or multiple
growth factors with appropriate control over temporal and spatial
presentation of these biomolecular cues. Current available technologies that have demonstrated potential in the modulation of release
profiles, according to the specific therapeutic needs, including traditional diffusion/swelling/degradation mediated, on-demand, affinity
and covalent binding based delivery are presented.
However, some challenges still exist. For example, highly inconvenient encapsulation methods often involving post-loading techniques
and low encapsulation efficiency, the stability of the protein, the safety
and biocompatibility of the delivery devices and the difficulties to translate in vitro results to in vivo situation. Further efforts in the design or
optimization of materials with minimal tissue response, sufficient stiffness, appropriate degradation and release profiles are needed in order
to make the translational step between academia and clinics successful.
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Please cite this article as: R. Censi, et al., Hydrogels for protein delivery in tissue engineering, J. Control. Release (2012), doi:10.1016/
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