Designing hydrogels for controlled drug delivery
Jianyu Li and David J. Mooney
John A. Paulson School of Engineering and Applied Sciences, and the Wyss Institute for
biologically Inspired Engineering, Harvard University, Cambridge, Massachusetts 02138, USA
Abstract
Hydrogel delivery systems can leverage therapeutically beneficial outcomes of drug delivery and
have found clinical use. Hydrogels can provide spatial and temporal control over the release of
various therapeutic agents, including small-molecule drugs, macromolecular drugs and cells.
Owing to their tunable physical properties, controllable degradability and capability to protect
labile drugs from degradation, hydrogels serve as a platform in which various physiochemical
interactions with the encapsulated drugs control their release. In this Review, we cover multiscale
mechanisms underlying the design of hydrogel drug delivery systems, focusing on physical and
chemical properties of the hydrogel network and the hydrogel–drug interactions across the
network, mesh, and molecular (or atomistic) scales. We discuss how different mechanisms interact
and can be integrated to exert fine control in time and space over the drug presentation. We also
collect experimental release data from the literature, review clinical translation to date of these
systems, and present quantitative comparisons between different systems to provide guidelines for
the rational design of hydrogel delivery systems.
Conventional drug administration often requires high dosages or repeated administration to
stimulate a therapeutic effect, which can lower overall efficacy and patient compliance, and
result in severe side effects and even toxicity
1–3
. For example, intravenously administered
Interleukin-12 (IL-12) resulted in systematic toxicities, including deaths in a clinical trial
4
.
Oral administration, which is the most common approach for delivering pharmaceuticals, is
frequently limited by poor targeting and short circulation times (<12 hours)
5
. Peptide and
protein drugs often have short serum half-lives of only minutes to hours
6
. To address these
issues, controlled drug delivery systems, including membranes, nanoparticles, liposomes and
hydrogels have been focused on in recent decades
7,8
. These drug delivery systems can
control how the drugs are available to cells and tissues over time and in space. They can, in
principle, leverage beneficial outcomes of therapeutics by enhancing their efficacy, and
reducing their toxicity and required dosage. The clinical use of drug delivery systems is
appreciable
7
, with a global market of over $150 US billion in 2013.
Hydrogels are a particularly appealing type of drug delivery system, and have been used in
many branches of medicine, including cardiology, oncology, immunology, wound healing,
and pain management. Hydrogels are composed of a large amount of water and a cross-
Correspondence to D. J. M. via mooneyd@seas.harvard.edu.
Competing interests
The authors declare no competing interests.
HHS Public Access
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Published in final edited form as:
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linked polymer network. The high water content (typically 70–99%) provides physical
similarity to tissues, and can give the hydrogels excellent biocompatibility and the capability
to easily encapsulate hydrophilic drugs. Moreover, because they are typically formed in
aqueous solutions, the risk of drug denaturation and aggregation upon exposure to organic
solvents is minimized. The cross-linked polymer network makes hydrogels solid-like, and
they can possess a wide range of mechanical properties. For example, their stiffness can be
tuneable
9
from 0.5 kPa to 5 MPa, allowing their physical properties to be matched with
different soft tissues in the human body
10–12
. The cross-linked network can impede
penetration of various proteins
13
, and thus is believed to protect bioactive therapeutics from
premature degradation by inwardly diffusing enzymes. This feature is particularly critical for
highly labile macromolecular drugs (for example, recombinant proteins and monoclonal
antibodies), which comprise an increasing percentage of new drugs approved, with many
others under development
14
. Since the introduction of human insulin, more than 130 protein
therapeutics have been approved by the Food and Drug Administration (FDA)
15
. The
attributes mentioned above make hydrogels attractive material systems for the delivery of a
large range of therapeutics.
Hydrogels differ in size, architecture and function, and together these features dictate how
they are utilized for drug delivery. In hydrogels, there are features with length scales
spanning from centimetres to sub-nanometres (FIG. 1). The macroscopic design largely
determines the routes by which hydrogels can be delivered into the human body (FIG. 1a).
Hydrogels can be formed into almost any overall size and shape. Micropores, if present, will
dramatically affect the overall physical properties (for example, the deformability), while
allowing for convective drug transport. On the several-nanometre scale, a cross-linked
polymeric network surrounds the water contained in the hydrogel network. Such networks
contain open spaces, the size of which is referred to as the mesh size of the network.
Importantly, the mesh size governs how drugs diffuse inside the hydrogel network (FIG. 1b).
Finally, at the molecular and atomistic scale, various chemical interactions may occur
between the drugs and the polymer chains (FIG. 1c). The polymer chains can possess
numerous sites for binding interactions with the drugs, and these can be pre-designed using a
diversity of physical and chemical strategies. The features at the mesh scale and the
molecular and atomistic scale are essential for controlled drug release. Because they are
decoupled from the macroscopic properties of the hydrogel, desirable features at each length
scale can often be designed independently of the other. This multiscale nature enables the
modular design of hydrogels, which can serve as a versatile platform to meet specific
application-based requirements.
Although some design requirements are common to all hydrogel delivery systems, others are
specific to the desired therapeutic application. In general, the fabrication of hydrogel
delivery systems needs to maintain the drug bioactivity, and through packaging, transport
and storage, both the drug and hydrogel must be chemically and physically stable.
Hydrogels can be delivered in a variety of manners, such as surgical implantation, local
needle injection or systemic delivery via intravenous infusion. The choice of delivery
method for a given application is based on maximizing the overall efficacy and patient
compliance. How the hydrogel releases the drug is often essential to achieve desirable
therapeutic outcomes, and the required duration of drug availability (short term versus long
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term) and its release profile (continuous versus pulsatile) depend on the specific application.
When the drug is exhausted, the hydrogel should be designed to either degrade to avoid
surgical removal, or to be re-used by drug refilling. The degradation of the hydrogel may
also need to be tailored to coordinate with tissue regeneration. Besides the general
requirements, there exist other application-based requirements. For example, in the treatment
of skin wounds, hydrogels are placed on dynamic surfaces to which they need to be adhesive
and conform, while being tough enough to tolerate the surface movement (for example,
strain of knee bending up to 50%) and deformation derived from the environment (for
example, compression and scratching).
16–18
The multiscale properties of hydrogels are essential for their functions to protect, target and
locally deliver drugs in a controllable manner. This Review will first illustrate the multiscale
structural properties of hydrogels, and how they affect encapsulation, delivery and release of
therapeutic agents. The discussion will start from the macroscopic scale, in which the key
design parameters include architectural factors (such as hydrogel size and porous structure).
The Review will proceed to the mesh scale, in which drug diffusion is regulated by the mesh
size and its temporal or stimuli-responsive evolution. The Review will then focus on the
molecular and atomistic scale. Any affinity or binding between the drugs and the polymer
chains will have a crucial role in the sustained or on-demand release of the drugs. Various
mechanisms to bind drugs with the polymer network, ranging from covalent conjugation to
secondary bonds such as electrostatic interactions and hydrophobic associations, will be
discussed. After a detailed discussion of the three length scales, this Review will provide
quantitative comparisons of the release kinetics between systems utilizing one or a
hybridization of different mechanisms. The use of hydrogels for the delivery of cells that
secrete therapeutic agents will also be discussed. The goal of this Review is to link a
fundamental understanding of hydrogels and their interactions with drugs to the rational and
practical design of hydrogel drug delivery systems.
Macroscopic design and delivery routes
The size of a hydrogel matters. Hydrogels can be cast or formed into practically any shape
and size, according to the requirements of the delivery route into the human body. Hydrogel
delivery systems can be classified into three main categories based on their size:
macroscopic hydrogels, microgels and nanogels (FIG. 2). Microgels and nanogels are
particulate hydrogels with dimensions on the order of micrometres and nanometres,
respectively.
Macroscopic hydrogels
The size of macroscopic hydrogels is typically on the order of millimetres to centimetres.
Correspondingly, they are usually either implanted surgically into the body or are placed in
contact with the body for transepithelial drug delivery (FIG. 2a)
7
. Success has been achieved
with surgically implanted hydrogels for drug delivery in the clinic, as exemplified by
INFUSE, a type I collagen gel that releases recombinant human bone morphogenetic
protein-2 (BMP-2), which is implanted surgically into the body for the treatment of long
bone fracture and spinal fusion
19
.
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Epithelial barriers that have been exploited for drug delivery include the skin, intestinal
epithelium and mucosa. Although these are impenetrable to macroscopic hydrogels, they can
be permeable to drugs released from the hydrogels. Hydrogels, including those fabricated
from synthetic polymers such as poly(vinyl alcohol) and poly(hydroxyl alkyl
methacrylate)
20
, and biopolymers such as alginate
21
, collagen
22
and chitosan
23
, are widely
used as wound dressings. Adaptation of these materials for transdermal drug delivery is
highly appealing, and these hydrogels have been used to deliver proteins such as insulin and
calcitonin
7
. Alginate hydrogels, similar to those in use for decades to treat wounds
24
, have
been shown to controllably deliver potential therapeutics like substance P to promote wound
healing
25
. Some of these materials are currently being evaluated in clinical trials. Indeed, a
recent clinical trial showed the efficacy and safety of a hydrogel delivering recombinant
human granulocyte-macrophage colony-stimulating factor (rhGMC-SF) for the healing of
deep, second-degree burn wounds
26
.
Placement of the hydrogel in the body may be necessary when the target site for drug
delivery is located deep within the tissue or when the biological barriers have low
permeability to the drug of interests. Placement of hydrogels within the body directly
bypasses the epithelial barriers and concentrates drug release at the target site
27,28
. However,
surgical implantation is invasive and potentially leads to patient discomfort and the risks
associated with surgery
29
. To address this issue, there has been great emphasis placed on the
development of injectable macroscopic hydrogels. These are largely designed on the
principles of gelation in the body (
in situ
-gelling hydrogels), gelation outside of the body but
with a transition to a flowable state upon application of sufficient shear stress (shear-thinning
hydrogels) to allow injection, or gelation outside of the body to a physical form that can be
readily collapsed for minimally invasive delivery followed by shape recovery
in vivo
(shape-
memory hydrogels).
In situ-gelling hydrogels—These systems can be injected in liquid form, followed by a
sol-gel transition inside the human body. The resulting hydrogels will take the shape of the
available space at the injection site and the sol–gel transition can be achieved with different
strategies. The simplest strategy is to use slow-gelling systems that allow gelation to be
initiated outside the body. Because this process occurs so slowly, the solution can be injected
before solidification occurs. Ideal kinetics are slow enough to prevent needle clotting, but
fast enough to prevent dilution of the pre-gel solution by body fluids once in the body. This
strategy has been applied with numerous gelation mechanisms, including charge
interaction
30,31
, stereocomplexation
32
and Michael addition
33
. As an example of a system
exploiting charge interactions, elastin-like polypeptides have been cross-linked via
electrostatic interactions between their cationic lysine residues and anionic
organophosphorus cross-linkers
34
. Non-covalent interactions between heparin and heparin-
binding peptides and proteins can be used to form hydrogels for growth factor delivery
35,36
.
Another
in situ
-gelling hydrogel was formed with a polyelectrolyte complex, which showed
a sustained release of proteins (for example, insulin and Avidin) over two weeks
37
. Certain
pre-gel solutions may require monomers or catalysts that are harmful to cells and tissues,
which provides a strong impetus to develop bio-orthogonal cross-linking reactions such as
copper-free click reactions for gelation (FIG. 2b)
38–40
. Click chemistry provides high
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specificity, quantitative conversion and modest gelation kinetics that are tuneable from a few
minutes to one hour, with no side reactions with biomolecules in the body.
Temperature-responsive systems that gel at body temperature have also been explored for
in
situ
gelation. Most natural polymers like gelatin form a gel upon lowering of temperature,
which would require their introduction to the body at supraphysiologic temperatures. By
contrast, certain synthetic polymers such as poly(N-isopropylacrylamide) (PNIPAm) and
poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) (PEO-PPO-PEO) undergo
reverse thermogelation (that is, gelation at elevated temperature)
41,42,43
and remain in a
flowable state at room temperature, thus providing a significant practical advantage.
Temperature-responsive polymers often contain hydrophobic domains that allow for the
inclusion of hydrophobic drugs (for example, doxorubicin)
43
. However, challenges exist for
these systems, such as their relatively low mechanical strength, poor physical stability, and
synthetic polymers such as PNIPAm may be non-biodegradable
44
. The stability of
temperature-responsive systems may be improved by adding additional covalent cross-links
after initial gelation
45,46
.
Shear-thinning hydrogels—Certain hydrogels can be pre-gelled outside of the body, and
then injected by application of shear stress. These shear-thinning hydrogels flow like low-
viscosity fluids under shear stress during injection, but quickly recover their initial stiffness
after removal of shear stress in the body. The shear-thinning behaviour is a result of
reversible physical cross-links (FIG. 2c). In contrast to covalent bonding, physical cross-
links are reversible, resulting from a dynamic competition between pro-assembly forces (for
example, hydrophobic interactions, electrostatic interactions and hydrogen bonding) and
anti-assembly forces (for example, solvation and electrostatic repulsion)
47
. Self-assembling
peptides have been extensively explored to make shear-thinning hydrogels owing to the
diversity of amino acids and the ease of sequence-specific peptide modification
48–50
. A
family of β-hairpin peptides (namely, MAX peptides) has been developed to make injectable
hydrogels for drug delivery
51,52
. These peptides contain two blocks of alternating
hydrophobic and charged amino acids in the form of long fibrils (up to 200 nm) that can
undergo a sol–gel transition. Peptide self-assembly can also be realized by taking advantage
of interactions between metal cations and amino acid residues of the peptides. This was
demonstrated with gelation of a β-sheet-rich fibrillar hydrogel with zinc ions
53
. Alginate
hydrogels are also shear-thinning and have been studied extensively. These hydrogels are
formed via electrostatic interactions between alginate and multivalent cations (for example,
calcium and zinc) (FIG. 2c); they can be readily injected via a needle after gelation in a
syringe and have been used to achieve sustained local delivery of bioactive vascular
endothelial growth factor (VEGF) in ischemic murine hindlimbs for 15 days, in contrast to
complete VEGF deprivation after 72 hours with bolus injection
30,31
.
Another approach to make shear-thinning hydrogels is to use dynamic covalent bonds
54,55
.
Such bonds are reversible and the hydrogels can behave similarly to self-assembling
peptides. The most widely used reactions are the dynamic exchange of C=N bonds in
imines, hydrazones and oximes. Other dynamic covalent reactions include reversible Diels–
Alder reactions (for example, furan and maleimide), boronic acid condensations and
disulfide exchange
56
. For example, complexation of boronic acids and diol compounds was
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