A review of stimuli-responsive nanocarriers for drug and gene delivery
Srinivas Ganta, Harikrishna Devalapally, Aliasgar Shahiwala, Mansoor Amiji⁎
Department of Pharmaceutical Sciences, School of Pharmacy, Northeastern University, 110 Mugar Life Sciences Building, Boston, MA 02115, United States
Received 15 September 2007; accepted 3 December 2007
Available online 11 January 2008
Nanotechnology has shown tremendous promise in target-specific delivery of drugs and genes in the body. Although passive and active
targeted-drug delivery has addressed a number of important issues, additional properties that can be included in nanocarrier systems to enhance the
bioavailability of drugs at the disease site, and especially upon cellular internalization, are very important. A nanocarrier system incorporated with
stimuli-responsive property (e.g., pH, temperature, or redox potential), for instance, would be amenable to address some of the systemic and
intracellular delivery barriers. In this review, we discuss the role of stimuli-responsive nanocarrier systems for drug and gene delivery. The
advancement in material science has led to design of a variety of materials, which are used for development of nanocarrier systems that can
respond to biological stimuli. Temperature, pH, and hypoxia are examples of “triggers” at the diseased site that could be exploited with stimuli-
responsive nanocarriers. With greater understanding of the difference between normal and pathological tissues and cells and parallel developments
in material design, there is a highly promising role of stimuli-responsive nanocarriers for drug and gene delivery in the future.
© 2008 Elsevier B.V. All rights reserved.
Keywords: Nanotechnology; Targeted delivery; Stimuli-responsive nanocarriers; pH; Temperature; Redox potential
1.Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
1.1. Target-specific pharmacotherapy: need for nanocarrier delivery systems . . . . . . . . . . . . . . . . . . . . . . . . . . .
1.2.Passive and active targeting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
1.3.Intracellular delivery and sub-cellular distribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
Stimuli-responsive delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
2.1. pH differences for stimuli-responsive delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
2.2.Temperature differences for stimuli-responsive delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
2.3.Changes in the redox status at the disease site . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
Illustrative examples of pH-responsive nanocarriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.1. pH-responsive polymeric nanocarriers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.2.pH-responsive polymer–drug conjugates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.3.pH-responsive liposomes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.4.pH-responsive micellar delivery systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
3.5.pH-responsive dendrimers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
llustrative examples of temperature-responsive nanocarriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
4.1. Temperature-responsive polymeric nanocarriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
4.2.Temperature-responsive liposomes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
Available online at www.sciencedirect.com
Journal of Controlled Release 126 (2008) 187–204
⁎Corresponding author. Tel.: +1 617 373 3137; fax: +1 617 373 8886.
E-mail address: firstname.lastname@example.org (M. Amiji).
0168-3659/$ - see front matter © 2008 Elsevier B.V. All rights reserved.
5.Illustrative examples of redox-responsive nanocarriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 199
5.1. Disulfide cross-linked polymeric nanocarriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 199
5.2.Disulfide cross-linked liposomes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200
Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200
1.1. Target-specific pharmacotherapy: need for nanocarrier
With parallel recent breakthroughs in molecular under-
standing of diseases and controlled manipulations of material at
the nanometric length scale, nanotechnology offers tremendous
promise in disease prevention, diagnosis, and therapy .
Among the various approaches for exploiting developments in
nanotechnology for biomedical applications, nanoparticulate
carriers offer some unique advantages as delivery, sensing and
image enhancement agents . Many bioactives used for
pharmacotherapy, while have a beneficial action, can also
exhibit side-effects that may limit their clinical application.
There has long been the desire to achieve selective delivery of
bioactives to target areas in the body in order to maximize
therapeutic potential and minimize side-effects. For example,
cytotoxic compounds used in cancer therapy can kill target
cells, but also normal cells in the body resulting in undesired
side-effects. For achieving better therapeutic application,
nanocarriers are considered for target-specific delivery of
drugs and gene to various sites in the body in order to improve
the therapeutic efficacy, while minimizing undesirable side-
effects. Improvements in target-to-non-target concentration
ratios, increased drug residence at the target site, and improved
cellular uptake and intracellular stability are some of the major
reasons for greater emphasis on the use of nanoparticulate
delivery systems. With nucleic acid-based therapeutic modal-
ities, there is substantial need for the therapeutic molecules to be
delivered to desired sub-cellular compartments in an efficient
and reproducible manner .
andinorganic materials includingnon-degradableandbiodegrad-
able polymers, lipids (liposomes, nanoemulsions, and solid-lipid
nanoparticles) self-assembling amphiphilic molecules, dendri-
mers, metal, and inorganic semiconductor nanocrystals (quantum
dots) [1,3]. The selection of material for development of
nanoparticulate carriers is mainly dictated by the desired
diagnostic or therapeutic goal, type of payload, material safety
profile, and the route of administration. Preponderance of
lipid, self-assembling, and a variety of inorganic nanoparticulate
opportunity for drug and gene delivery where the delivery system
becomes an active participant, rather than passive vehicle, in the
optimization of therapy. Several families of molecular assemblies
or active targeting. Liposomes, polymeric nanoparticles, block
copolymer micelles and dendrimers are colloidal molecular
assemblies (Fig. 1). The composition of each class of these
molecular assemblies can be manipulated to obtain nanocarrier
of desired stimuli-responsive property. The benefit of stimuli-
responsive nanocarriers is especially important when the stimuli
are unique to disease pathology, allowing the nanocarrier to
respond specifically to the pathological “triggers”. Select
examples of biological stimuli that can be exploited for targeted-
drug and gene delivery include pH, temperature, and redox
microenvironment [2,5–8]. The extracellular and intracellular pH
Fig. 1. Different types of stimuli-responsive nanocarriers.
188S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
profile of biological system is greatly affected by diseases. For
instance, in solid tumors, the extracellular pH tends to be
significantly more acidic (∼6.5) than the pH of the blood (7.4) at
37 °C . In addition, the pHvalues of endosomal and lysosomal
vesicles inside the cells are also significantly lower that the
cytosolic pH. By selecting the right material composition, it is
possible to engineer nanocarriers that can exploit these pH
differences and allow for delivery of the encapsulated payload to
specifically occur in select extracellular or intracellular sites.
Temperature is another variable that can be exploited in
specifically releasing the nanocarrier-delivered drugs or genes to
a select target site . For instance, using temperature-sensitive
nanocarriers one could envision a delivery system that will only
release the payload at temperatures above 37 °C. Such a system
of hyperthermic stimuli to the disease area, the drug would be
readily available in a localized region . Lastly, intracellular
glutathione (GSH) levels in tumor cells are 100–1000 fold higher
than the extracellular levels . This concentration gradient can
be exploited using disulfide cross-linked nanocarriers that will
release the payload inside the cell. Such a system is especially
relevant in delivery of nucleic acid-based therapies, such as
plasmid DNA, small interference RNA, or oligonucleotide, since
these moleculeshave to reach intracellular targets in a stable form
for efficient therapeutic effect.
Another possible strategy is physical targeting of drugs and
genes by external stimuli (magnetic field, ultrasound, light and
heat) [12–16]. An interesting example is targeted delivery of
iron oxide nanoparticles using magnetic field. Upon the
administration, the drug immobilized magnetite carrier can
accumulate at targeted site under the direction of external
magnetic field . During the last decade, ultrasound has
attracted growing attention in the targeted-drug delivery.
Ultrasound has been used to achieve the targeted delivery to
the tumor by local sonication after the injection of micellar
encapsulated drugs [15,16]. In addition to tumor uptake, this
technique also allows the uniform distribution of micelles and
drug throughout the tumor tissue . Light-responsive
nanocarriers have also gained recent attention. Designing of
light sensitive polymeric systems that undergo reverse micelli-
zation/disruption under the action of light is an attractive idea
that would allow external control of drug release .
Discussion of extensive current literature on the external stimuli
is beyond the scope of the present review. Several examples
illustrating approaches to designing external stimuli are
available for further reading from the references [12–16].
1.2. Passive and active targeting
For systemic therapy, passive and active targeting strategies
are utilized. Passive targeting relies on the properties of the
delivery system and the disease pathology in order to
preferentially accumulate the drug at the site of interest and
avoid non-specific distribution. For instance, poly(ethylene
glycol) (PEG)- or poly(ethylene oxide) (PEO)-modified nano-
carrier systems can preferentially accumulate in the vicinity of
the tumor mass upon intravenous administration based on the
hyper-permeability of the newly-formed blood vessels by a
process known as enhanced permeability and retention (EPR)
effect, schematically illustrated in Fig. 2. Maeda and colleagues
[19,20] first described the EPR effect in murine solid tumor
models and this phenomenon has been confirmed by others.
When polymer–drug conjugates are administered, 10–100 fold
higher concentrations can be achieved in the tumor due to EPR
effect as compared to administration of free drug . The EPR
effect has been also present in other diseases such as chronic
inflammation and infection. Thus, the application of nanocar-
riers is expected to have therapeutic benefits for treating these
diseases as well . The tendency of nanocarriers to localize in
the reticuloendothelial system also presents an opportunity for
passive targeting of bioactives to the macrophages present in the
liver and spleen. For example therapies can be used to treat
intracellular infections such as candidiasis, leishmaniasis and
listeria; where macrophages are directly involved in the disease
process . Other approaches for passive targeting involve the
Fig. 2. Schematic illustration for passive targeting using the enhanced permeability and retention (EPR) effect.
189 S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
use of specific stimuli-sensitive delivery system that can release
the encapsulated payload only when such a stimuli is present.
For instance, the pH around tumor and other hypoxic disease
tissues in the body tends to be more acidic (i.e., ∼5.5 to 6.5)
relative to physiological pH (i.e., 7.4). Using pH-sensitive poly
(beta-amino ester) (PbAE) nanoparticles, we have found
significant enhancement in drug delivery and accumulation in
the tumor mass as compared to drug administration in PCL
nanoparticles, a non-pH-sensitive polymer, and in aqueous
solution [5–7]. Further approaches for passive targeting involve
size of the nanocarriers and surface charge modulation.
Nanoparticles of b200 nm in diameter and those with positive
in the tumor mass for longer duration than either neutral or
negatively charged nanoparticles . Recently, we have
examined the role of combination paclitaxel and the apoptotic
second messenger, C6-ceramide, when administered concur-
rently in PEO-modified PCL nanoparticles to overcome multi-
drug resistance in cancer [21,24].
type B gelatin-based nanoparticles has been very effective in
in nanoparticulate formulations is able to physically encapsulate
plasmid DNA at neutral pH . The physically encapsulated
effective in vitro and in vivo in transfection of reporter plasmid
DNA expressing green fluorescent protein and beta-galactosidase
[27,28]. Upon systemic administration in C57/BL6J mice bearing
Lewis lung carcinoma, the PEG-modified gelatin nanoparticle
thiolated gelatin nanoparticles could also encapsulate DNA and
transfect tumor cells in response to higher intracellular glutathione
levels [8,29]. When PEG-modified thiolated gelatin nanoparticles
encapsulated with plasmid DNA encoding for soluble vascular
endothelial growth factor receptor 1 (sVEGFR-1 or sFlt-1), highly
cells and in vivo in an orthotopic tumor model. In addition, the
expressed sFlt-1 was very effective in suppressing tumor growth
and angiogenesis .
Active targeting to the disease site relies, in addition to PEG
modification of nanocarriers to enhance circulation time and
achieve passive targeting, coupling of a specific ligand on the
surface that will be recognized by the cells present at the disease
site . Using solid tumor as an example again, there are
several strategies that can be adopted for surface modification of
nanocarrier systems for effective targeted delivery to the tumor
cells or to endothelial cells of the tumor blood vessels. Since
tumor cells are rapidly proliferating, they over-express certain
receptors for enhanced uptake of nutrients, including folic acid,
vitamins, and sugars. When the surface of nanocarriers is
modified with folic acid, they can be targeted to the tumor cells
that over-express folate receptors (Fig. 3). In addition, Fig. 3
also illustrates the intracellular delivery of folate anchored
nanocarrier through endocytosis process and releasing its
contents in response to internal stimuli. Tumor and capillary
endothelial cells also express specific integrin receptors, such as
αvβ5or αvβ3that can bind to arginine–glycine–aspartic acid
(RGD) tripeptide sequence. RGD-modification, therefore, has
been used to direct nanocarriers to tumor cells and capillary
endothelial cells of the angiogenic blood vessels. The phage
display method has been used to identify specific peptide
sequences that can be used for targeting tumors and other
disease areas in the body . For example, Schluesener et al.
Fig. 3. Schematic illustration of active drug targeting with surface-modified micelles.
190S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
 used in vivo phase display of recombinant M13 phages as a
tool to select peptides targeting pathological endothelium of
experimental rat brain tumors. One of the FDA approved
targeted therapeutics is Adalimumab® antibody; a human anti-
TNF IgG1 used against rheumatoid arthritis is generated by
phage display technique . Recently, Farokhzad et al. 
have elegantly described the use of aptamers, nucleic acid
constructs that specifically recognize prostate membrane
antigen on prostate cancer cells. The aptamer technology
provides an additional strategy for active targeting of tumor
cells in the body. Development of monoclonal antibodies
against specific epitopes present only on tumor cells allows for
other targeting strategies. For instance, HER2 specific antibody
(Trastuzumab® or Herceptin®) modified nanoparticles were
able to localize and deliver the therapeutic payload specifically
in HER2 expressing tumor cells . Using a monoclonal
antibody 2C5 that specifically recognizes anti-nuclear histones,
Torchilin's group has developed various strategies for active
targeted delivery of drugs to the tumor mass using liposomes
and micellar delivery systems [36,37]. Other groups have used
transferrin, an iron-binding protein, for surface modification of
nanocarriers for delivery to tumors . Epidermal growth
factor receptors are over-expressed in breast or prostate cancers,
making it a good candidate for targeting gene-delivery
complexes . Apart from tumor, epithelial surfaces of the
lungs and gastrointestinal tract, endothelial cells lining the
blood vessels, muscle myoblasts and skin fibroblasts are also
potential targets for gene delivery . Along with targeting
ability, the nanocarriers also can be tailored of stimuli-
responsive property to enhance the transfection ability of the
carriers. For example, Oishi et al.  showed that pH-
responsive and PEGylated nanogels bearing a lactose group at
the PEG end display endosomolytic abilities and achieving the
enhanced transfection efficiency.
1.3. Intracellular delivery and sub-cellular distribution
Once the nanocarriers are delivered to the specific diseased
organ or tissue, they may need to enter the cells of interest and
ferry the payload to sub-cellular organelles (Fig. 4). In this case,
non-specific or specific cell penetrating strategies need to be
adopted . Non-specific cell uptake of nanocarriers occurs by
endocytotic process, where the membrane envelops the
nanocarriers to form a vesicle in the cell called an endosome.
The endosome then shuttles the content in the cell can fuse with
enzymes. Endocytosed nanocarriers usually travel in a specific
direction and converge at the nuclear membrane. Endosomal
acidic condition is deterrent to therapeutic molecules present in
the nanocarrier. This bottleneck in gene delivery can be
responsible for the degradation of N99% of the internalized
DNA. Efficient gene delivery is achieved by buffering the
endosomes for safe release of its contents. For example, the
buffering capacity of the polycationic carriers can hamper the
acidification of theendosomes, causing itto swell and burst,as a
consequence safe release of trapped contents [43,44]. Specific
cellular uptake can occur through receptor-mediated endocy-
tosis, where upon binding of the ligand-modified nanocarrier
nanocarrier–receptor complex and vesicular transport through
the endosomes. Following dissociation of the nanocarrier–
receptor complex, the receptor can be re-cycled back to the cell
membrane. Recently,inorder toenhance cellular uptake,asurge
of research effort has been directed towards development of
arginine-rich cell penetrating peptides (CPP's) . Based on
the pioneering work of Dowdy's group HIV-1 Tat peptide was
identified to promote non-specific intracellular localization of
has been supported by other groups and a number of cationic
Fig. 4. Intracellular delivery and sub-cellular localization of nanocarriers.
191S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
peptides have been identified, including penetratin, to enhance
CPP's enhance cell permeation is still a subject of controversy,
but recent data show that it may be through endocytosis as well
. Following cellular internalization, stability of the payload
in the cytosol and uptake by specific organelle, such as the
nucleus, is also essential for nucleic acid-based therapeutics.
Weissig's group has attempted to direct various nano-sized
delivery systems to mitochondria using delocalized cationic
amphiphiles and other mitochondriotropic vector systems .
For efficient systemic gene therapy using non-viral vectors,
nuclear import of plasmid DNA in non-dividing cells is
considered to be the major limiting factor.
2. Stimuli-responsive delivery
2.1. pH differences for stimuli-responsive delivery
The pH profile of pathological tissues, such as upon
acquisition of inflammation, infection, and cancer, is signifi-
cantly different from that of the normal tissue . The pH at
systemic sites of infections, primary tumors, and metastasized
tumors is lower than the pH of normal tissue. For example, pH
of the regiondrops from 7.4 under normal conditions to 6.5 after
60 h following onset of inflammatory reaction . This
behavior can be utilized for the preparation of stimuli-
responsive drug or gene-delivery systems, which can exploit
the biochemical properties at the diseased site for targeted
delivery. The cellular components display trans-membrane pH
gradient in normal as well as pathological conditions, which can
also be used for intracellular delivery of macromolecules. Since
for many macromolecules with therapeutic potential, such as
antigenic peptides targeting the major histocompatibility
complex-I pathway, antisense oligonucleotides targeting
mRNA, or corrective genes targeting the nucleus, delivery to
the cytosol becomes necessary for therapeutic effect. Cellular
components such as the cytoplasm, endosomes, lysosomes,
endoplasmic reticulum, golgi bodies, mitochondria and nuclei
are known to maintain their own characteristic pH values .
The pH values range from 4.5 in the lysosome to about 8.0 in
the mitochondria. Given these pH gradients, therapeutic
compounds with pKabetween 5.0 and 8.0 can exhibit dramatic
changes in physicochemical properties.The apparent membrane
partitioning of a weak acid can increase significantly as pH
decreases its pKaand decreased membrane partitioning was
observed for a weak base due to the neutralization of charges.
These pH-responsive compounds can be incorporated into
nanocarriers or conjugated as such to macromolecules to
achieve efficient intracellular delivery and sub-cellular localiza-
tion of macromolecules.
Anticancer therapy is normally associated with unwanted
side-effects which are largely due to non-specific distribution of
drugs into normal tissue. The selective targeting of drugs or
macromolecules in tumor tissue is of high interest in cancer
therapy. The pH , surface charge , and density of low
density lipoprotein receptors  are the factors that show
notable differences among the normal and tumor tissues. All
these properties are known to influence the drug's physico-
chemical properties and are exploited for enhanced delivery to
the target site.
The pH is on average lower in the tumor mass and than
normal tissue . Since tumors proliferate very rapidly, the
vasculature of tumor is often insufficient to supply enough
nutritional and oxygen needs for the expanding population of
tumor cells. This results in difference in metabolic environment
between the various solid tumors and the surrounding normal
tissue. The insufficient oxygen in tumor leads to hypoxia and
causes production of lactic acid and hydrolysis of ATP in an
energy-deficient environment contributes to an acidic micro-
solid tumors have lower extracellular pH (b6.5) than the
surrounding tissues (pH 7.5). The pH is compartmentalized in
tumor tissue into an intracellular component (pHi), which is
similar in tumor and normal tissue and an extracellular
component (pHe), which is relatively acidic in tumors .
This gives rise to cellular trans-membrane pH gradient
difference between normal tissue and tumor tissue, which may
be exploited for the delivery of drugs to tumor that are weak
of weakly acidic drugs diffuses freely across the cell membrane,
upon reaching a relatively basic intracellular compartment,
becomes trapped within a cell, leading to substantial difference
in drug concentration between normal and tumor tissue.
2.2. Temperature differences for stimuli-responsive delivery
The use of hyperthermia as an adjunct to radiation or
chemotherapy of various types of solid tumors has become an
area of active investigation for the past 20 years, in part due to
the improvements in instrumentation and temperature monitor-
ing technique, as well as, an increasing understanding of the
biology of hyperthermia . In addition, tumor cells seem to
be more sensitive to heat-induced damage than normal cells.
Recently the majority of clinical studies of hyperthermia have
used super-paramagnetic iron oxide-containing liposomes or
nanoparticles [55,56]. The liposomes and nanoparticles provide
a method for intracellular delivery and localization of the iron
oxide particles. Unlike the external probes that can heat the
surrounding normal tissues, the magnetic nanoparticle
hyperthermia is appealing because it offers a way to ensure
that only the intended target is heated. A typical in vivo dose of
100–120 kHz alternating magnetic field is applied to experi-
mental tumor models for about 30 min to achieve temperatures
between 40 and 45 °C. For example, radiofrequency ablation in
combination with liposomal doxorubicin (DOX) has been
examined as an approach to promote the DOX accumulation in
highly refractory tumors .
Studies by Dewhirst and colleagues  at Duke University
have shown that a greater fraction of the intravenously
administered liposomes and other nanocarriers of up to
400 nm in diameter were able to extravasate from the tumor
microcirculation and into the tumor mass upon heating to 42 °C
in human ovarian carcinoma (SKOV-3) xenograft model. In
addition, the same authors have shown a higher concentration
192S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
and enhanced efficacy of doxorubicin in tumor mass upon
delivery in heat-sensitive liposomal formulations . Both of
these studies and others clearly point to the fact that nanocarrier
mediated delivery of anticancer drugs can be positively
influenced by localized hyperthermia. Heat-responsive trig-
gered drug delivery using polymeric nanocarriers with low
melting point offers additional advantage of creating micro-
emboli in the tumor vasculature as a drug delivery depot and for
restricting oxygen and nutrient supply to tumor mass.
2.3. Changes in the redox status at the disease site
The potential for nucleic acid therapeutics has grown rapidly
over the past decade. Such therapeutics include the use of
plasmid DNA, antisense oligonucleotides, ribozymes, peptide
nucleic acids and silencing RNAs for gene therapy to treat
genetic and acquired diseases . However, limited cytoplas-
mic delivery of genes is a challenge facing the development of
gene therapies that act on intracellular compartments. Gene-
delivery systems sensitive to intracellular stimuli mechanisms
have been the subject of intensive scrutiny as they would allow
formation of gene carriers stable in blood circulation that
disintegrate after intracellular uptake . Redox potential as a
stimuli mechanism has been put forward in gene delivery apart
from other stimuli mechanism like pH change and temperature
[11,60]. A high redox potential difference (∼100–1000 fold)
exists between the reducing intracellular space and oxidizing
extracellular space that enables it as a potential stimuli for
delivery of gene therapeutics . Redox-sensitive nanocarriers
rely on the higher intracellular reduction capacity compared to
the extracellular milieu . Gene-delivery systems containing
disulfide linkages that are known to be taken up by endocytosis
Illustrative examples of pH-responsive nanocarriers used for drug and gene delivery
Stimuli-responsive polymer or lipidDrug or genePropertyTherapeutic outcome Ref.
Paclitaxel Rapid sol-to-gel
change in pH
Insulin was released
in a pH-dependent
fusion at endosomal
Induces a coil-to-
globule transition at acidic
pH and is exploited
to destabilize the
Higher drug release under
pH conditions (5-5.5)
Hyrdazone bond cleaved
intracellularly at low pH
(5–6) and release the drug
Drug release was faster at low
pH and able to target
Micelles were destabilized in
the pH range of 7.2–6.6
Showed good antitumor
effect in melanoma tumor
This system has a potential
as a polypeptide drug carrier
N-Ac-poly(l-histidine)-graft-poly(l-lysine) Plasmid DNAShowed higher transfection
efficacy in 293T cells
anticancer activity in
murine tumor models
Micelles Self-assembling amphiphilic block copolymers, poly
higher antitumor activity
in C-26 bearing mice
These multifunctional micelles
enhanced the KB cellular
uptake and increased cell kill
Significantly enhanced KB cell
Mixed micelles — PLLA/PEG block copolymer
with PEG via disulfide linkage
Histidine-modified galactosylated cholesterol
derivative — cationic liposome
Transferrin-modified liposomes (Tf-L) with a
pH-sensitive fusogenic peptide (GALA)
DoxorubicinShowed effective MCF-7 cell
inhibition and uptake in vitro
Effective cytotoxicity on B-
Increased gene transfection
This system could be used in
selective mitochondrial targeting
for cancer therapy
Effective in the treatment
of viral infections cancer
or inflammatory diseases
Anionic liposomes containing
At low pH it allows fusion of
and destabilization of
Drug was released in a pH-
internalization of the carrier
Dendrimers Polyamidoamine dendrimerChlorambucil
Drug linked to polyamidoamine
dendrimer via pH-sensitive linker
DoxorubicinImproved cytotoxicity on
193 S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
may undergo disulfide cleavage in the lysosomal compartments
. The glutathione pathway which controls the intracellular
redox potential , is significantly involved in this stimuli
mechanism. The use of reducible polymers for gene delivery
achieves efficient gene transfection in vivo while limiting
effects of toxicity. For example polyethylene imine(PEI)/DNA
polyplex formed by disulfide bond enhanced the transfection
efficiency . Redox-sensitive carriers with long circulating
ability were obtained by coupling PEG to polycations. Redox-
sensitive PEGylated polylysine polyplexes exhibited enhanced
blood levels of DNA .
3. Illustrative examples of pH-responsive nanocarriers
Polymeric nanoparticles, polymeric conjugates, liposomes,
micelles, and dendrimers have been developed to provide
responsive drug release behavior. A summary of the different
types of pH-responsive nanocarriers is shown in Table 1.
3.1. pH-responsive polymeric nanocarriers
Nanocarriers constructed from the stimuli-responsive poly-
mers have been proposed as anticancer drug delivery systems.
The physical properties, such as swelling/deswelling, particle
disruption and aggregation of stimuli-responsive nanocarriers
change in response to the changes in environmental condition.
In turn, these properties alter the interactions of the nanocarriers
with the cells and trigger the drug release from slow to fast at the
tumor site. The pH-sensitive poly(β-amino ester) (PbAE)
constitutes a novel class of biodegradable cationic polymers
for development of site-specific drug and gene-delivery
systems. In the acidic microenvironment of tumor (pHb6.5)
PbAE undergoes rapid dissolution and releases its content at
once. Using PbAE nanoparticles, we have found significant
enhancement of paclitaxel accumulation in the tumor tissue as
compared to PCL nanoparticles containing paclitaxel, a non-
pH-sensitive polymer, and in aqueous solution [5–7,79]. In
another study, pullulan acetate, a linear polysaccharide has been
introduced with sulfadimethoxine (SDM, pKa6.1) to prepare
pH-sensitive and self-assembled hydrogel nanoparticles, which
also demonstrated enhanced adriamycin (ADR) release in
response to lower pH and increased cytotoxicity .
particular challenge mainly because of its high degradative
activities of the endosomal/phagosomal compartment. The
inflammatory reactions as well as in the foreign body responses
are mediated through macrophages.Inaddition, macrophages are
T and B lymphocytes. Therefore, macrophages are an important
target to treat inflammatory and immune responses. There is a
growing interest in developing pH-sensitive nanocarriers that can
enhance cytoplasmic entry of genes in macrophages, such as
antisense oligonucleotides, and antigenic protein and peptides.
that incorporate three functionalities of viruses and toxin: 1)
targeting ligands that direct receptor-mediated endocytosis, 2) a
pH-responsive agent that selectively disrupts the endosomal
membrane, and 3) the therapeutic component which is delivered
as an active agent into the cytoplasm termed as “encrypted
polymers” by the authors . The encrypted polymer has
provided pH-sensitivity through acid-cleavable acetal bonds that
polymer backbone. These systems can target and direct cellular
uptake, as well as enhance cytosolic delivery by disrupting
endosomal membranes in a pH-dependent manner. A functiona-
lized monomer, pyridyl disulfide acrylate was incorporated into
an amphiphilic copolymer of methacrylic acid and butyl acrylate.
This resulted in a pH-sensitive and glutathione-membrane
disruptive terpolymer with thiol functional groups that can
allow rapid conjugation of peptidal moieties. Oligonucleotides
and peptides conjugated to this polymer show the significant
enhancement in cytoplasmic delivery of these molecules .
The pH-sensitivity of polymeric carriers can be manipulated
by controlling the length of hydrophobic alkyl segments. Stayton
et al.  demonstrated that the pH profile of poly(alkyl acrylic
acid) polymers is controlled by the choice of the alkylacrylic acid
monomer (Fig. 5) and by the ratio of the carboxylate-containing
alkylacrylic acid monomer to alkylacrylate monomer. Poly
(propylacrylic acid) (Fig. 5) is from the family of poly
disruption at pHs below 6.5 and significantly enhances in vitro
transfections of lipoplex formulations .
3.2. pH-responsive polymer–drug conjugates
to exploit the acidic environment of tumor. Presence of acid-
sensitive spacers between the drug and polymer enables release
of drug either in relatively acidic extracellular fluids or, after
endocytosis in endosomes or lysosomes of tumor cells. Kamada
and colleagues  synthesized a pH-sensitive polymeric
carrier, in which a poly(vinylpyrrolidone-co-dimethyl maleic
anhydride) (PVD) was conjugated to doxorubicin (DOX), that
could gradually release free drug in response to changes in pH
Fig. 5. Representative chemical structures of stimuli-responsive polymers. A. pH-
sensitive polymer blocks: a) poly(acrylic acid), b) poly(methacrylic acid), c) poly
(2-ethyl acrylic acid, d) poly(2-propyl acrylic acid. B. Temperature-sensitive
polymer blocks: a) N-isopropylacrylamide, b) poly(organophosphazenes).
194S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
[i.e., from near neutral (∼7.0) to slightly acidic pH (∼6.0)]. The
authors concluded that the superior anticancer activity of PVD–
tumor accumulation of the drug. Water-soluble polymeric drug
carriers based on copolymers of N-(2-hydroxypropyl)methacry-
a group of potential drug delivery systems capable of delivering
drugs to model tumors or tumor cells in mice . HPMA
been described . Ulbrich et al.  have synthesized and
examined the properties of antibody-targeted pH-sensitive
polymer–DOX conjugates in which DOX was attached to a
water-soluble polymer carrier via a simple hydrolytically labile
linker containing the hydrazone bond. Hydrazone linkage
hydrolytically controls the release of DOX from the carrier
and its activation after transfer of the polymer–drug from the
blood circulation and extracellular environment into intracel-
lular compartments. Unlike classic conjugates, these conjugates
do not require lysosomal enzymes for biological activity of the
hydrazone linker can be released in the acid environment of
endosomes and/or lysosomes after cellular uptake of the
conjugates by endocytosis . Anticancer drugs conjugated to
serum albumin displayed greater anticancer activity, for example,
in vitro studies of acid-sensitive chlorambucil and anthracycline
conjugates with serum albumin showed higher antiproliferative
activity, and acid-sensitive DOX albumin conjugates displayed
to free drug [88,89]. Polyacetals rapidly undergo hydrolysis at
acidic pH and have potential for development as biodegradable
ethers, serinol and PEG can be used to synthesize biodegradable,
hydrolytically labile amino-pendent polyacetals (APEGs) suita-
DOX (APEG–DOX) and compared to HPMA copolymer–DOX
conjugates which are in clinical development and found to be
more promising .
3.3. pH-responsive liposomes
Extensive research has been carried out during the past
30 years to use liposomes, bilayered phospholipid vesicles with
the anticancer drugs and gene delivery and are principle areas of
interest. As a consequence some of the liposomal formulations
are approved for clinical trials or in the market (e.g.,
DaunoXome, Doxil\Caelyx, Mycet, AmBisome). To achieve
the pH-sensitivity to release active contents, the liposomes can
be tailored from the pH-sensitive components. The pH-sensitive
liposomes are endocytosed in the intact form and fuse with the
endovascular membrane as a consequence of the acidic pH
inside the endosome, and release its active contents into the
cytoplasm . Recent studies mainly focus on the construct of
new lipid compositions that attribute pH-sensitivity to lipo-
somes or modification of liposomes with various pH-sensitive
polymers  and imparting hydrophilicity to the liposomal
surface for longevity and ligand-mediated targeting. This
combination of pH-sensitivity, longevity and targeting ability
of liposomes can effectively deliver their contents into the
Anionic pH-sensitive phosphatidylethanolamine (PE) incor-
porated liposome are used to deliver the antisense oligonucleo-
tides intracellularly. These are stable in the blood, however, they
undergo phase transition at acidic endosomal pH, in turn this
facilitates cytoplasmic delivery of oligos . Long circulating
PEG-modified pH-sensitive liposomes have been prepared using
a combination of PEG and a pH-sensitive, terminally alkylated
copolymer of poly(N-isopropylacrylamide) (NIPAAm) and
methacrylic acid . pH-sensitive liposome attached with
target-specific ligands (folate and Tf) has been described for
cytosolic drug delivery . Imparting hydrophilic property,
particularly by incorporating PEG by covalent linking or physical
coating of the carrier surfaces, prolongs the residence time of
and thus minimizing the clearance in reticuloendothelial system
(RES) organs. This long circulating property of nanocarrier
promotes EPR effect and more valuable in passive targeting of
drugs or genes for cancer therapy. This finding has lead to
designing of pH-sensitive liposomes by anchoring the surface
with carboxylic groups, but these failed to respond to small
changes in physiological pH because of intrinsic pKa of
The pH-sensitive liposomes have been designed to deliver
highly hydrophilic molecules or macromolecules into the
cytoplasm. These pH-sensitive liposomes destabilize under
acidic conditions present in the endosomes and usually contain
PE and titrable stabilizing amphiphiles. The well investigated
class of pH-sensitive liposomes consists polymorphic lipids
such as unsaturated PE with mildly acidic amphiphiles that act
as stabilizers at neutral pH . The amphiphile head group
gets protonation in the acidic environment and causes a
destabilization of liposomal bilayer which is usually accom-
panied by the release of liposomal contents . These systems
have been successfully investigated for the in vitro cytoplasmic
delivery of antitumor drugs, protein toxins, antigen, antisense
oligonucleotides and plasmid DNA [98,99]. Lipids other than
PE incorporated in liposomes also show pH-sensitive behavior
such as cholesteryl hemisuccinate (CHEMS) and poly(organo-
phosphazenes) (Fig. 5) [93,100]. pH-sensitive liposomes
prepared from the hydrophobically-modified copolymers of
NIPAAm bearing pH-sensitive moiety have been examined for
the release of water-soluble fluorescent marker, pyranine, and
an amphipathic cytotoxic anticancer drug DOX. The release
from the copolymer modified liposomes is found to be pH
dependent . Lipids like dioleoyl phosphatidyl ethanola-
mine (DOPE) form a non-bilayer structure in an aqueous
medium at neutral pH but when combined with stabilizing
components such as CHEMS are able to assemble into a bilayer
. The liposomes based on these bilayered components are
destabilized in the acidic environment of the endosomes and
rapidly released their contents . Ishida et al.  have
found that DOPE liposomes stabilized with cleavable lipid
derivative of PEG (mPEG-S-S-DSPE) enhanced the DOX
delivery into the nuclei of the CD19 epitope on B-lymphoma
195S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
cells and increased cytotoxicity compared to non-pH-sensitive
A pH stimuli release of drugs encapsulated in liposomes can
be achieved both with drugs that increase as well as decrease
membrane permeabilities upon acidification, as long as the
intraliposomal buffer strength and pH is rationally selected. Lee
et al.  investigated the folate receptor-targeted liposomes
with three different compounds whose pKais dependent on pH.
Anionic 5(6)-carboxyfluorescein converts into non-ionic at
endosomal pH and releases at endosomes. These compounds
can be encapsulated into liposomes at neutral pH. As a result of
decreasing pH of intraliposomal at endosomal acidic pH, the
liposomal contents can be released to endosome. Another
category of compounds retains both anionic and cationic
charges at endosomal pH and retains at endosomes for a long
time like sulforhodamine B. DOX in its cationic form in strong
acidic buffer when loaded into liposomes, displays endocytosis
triggered release, since sufficient uncharged DOX remains at
3.4. pH-responsive micellar delivery systems
Micelles are spherical supramolecular nanoassemblies ran-
ging from 20 to 100 nm in size that have attracted considerable
interest as potential drug nanocarriers due to their unique
properties such as high solubility, high drug loading capacity
and low toxicity . The small size of micelles is responsible
to avoid rapid renal exclusion and uptake by the RES .
This in turn prolongs the blood circulation of micelles and
facilitates their passive accumulation at tumor tissue. Many
approaches have been described to the development of pH-
sensitive micelles that can exploit acidic environment at tumor
tissue to unload its contents . One such approach is
attaching the “titratable” groups such as amines or carboxylic
acids into the block copolymers such that the micelle formation
is controlled by the protonation of these groups. Despite the fact
that, only a few of these have been found to undergo transitions
in the physiologically relevant pH range of 5.0–7.4 and proved
to encapsulate drugs [72,101].
In their recent work, Bae and co-workers have reported the
intracellular pH-sensitive polymeric micelles that can release
the anticancer drug, DOX, in response to acidic pH at
endosomes (pH 5.0–6.0) and lysosomes (pH 4.0–5.0), which
can maximize the DOX delivery efficiency to the tumor tissue
. Bae and co-workers have also found that water-soluble
polymers modified with a new pH-sensitive functional group (a
weak acid of sulfonamide) self-assembled nanoparticles
showed increased drug release, interaction with and internaliza-
tion into cells at tumor pH . A mixed micellar formulation
has been explored for the paclitaxel, in which micelle-unimer
transition will takes place due to the ionization–non-ionization
of sulfamethazine oligomer in the pH range (pH 7.2–8.4) above
the CMC . A polyamine, poly(L-histidine) is investigated
in pH-sensitive micellar systems because of its amphoteric
property and fusogenic activity of the imidazole group and its
interaction between endosomal membrane and poly(L-histidine)
. Folate receptor present on tumor has been utilized for
targeting of anticancer drugs, genes, and radiopharmaceuticals
via folate receptor-mediated endocytosis [109,110]. Folate
receptors are highly expressed on various tumors such as
ovarian, lung, breast, brain, colon and kidney cancers . In
one such study, poly(L-histidine) based micelles were further
conjugated with ligands either folic acid or biotin to enhance the
tumor uptake by folate receptor-mediated or biotin receptor-
mediated endocytosis . The activity is further enhanced by
preparing the mixed micelles [poly(L-histidine)/PEG and poly
(L-lactic acid)/PEG block copolymer] conjugated with folic
acid, enhanced cytotoxicity has been found in vitro using these
Water-soluble block copolymers that can exhibit different
forms in aqueous environment have been in recent focus for
drug delivery . Block copolymers based on poly[4-
vinylbenzoic acid (VBA) and 2-N-(morpholino)ethyl metha-
crylate (MEMA) can dissolve in pH above 6.0. VBA forms
micelles if the pH is below 6.0, whereas MEMA micelles can
form in alkaline pH . Triblock copolymers such as poly
(acrylic acid)-b-polystyrene-b-poly(4-vinyl pyridine) (PAA-b-
PS-b-P4VP) have also been shown to exists in different forms
by changing the aqueous pH. Poly[2-(dimethylamino)ethyl-
(DEA-MEMA) forms MEMA-core micelles at pH 6.5 in the
presence of 1.0 M Na2SO4, whereas the DEA block forms a
micellar core at pH 1.0 . These unique characteristics of
amphiphilic block copolymers are responsible for the special
advantage of using micelles for various drug delivery applica-
tions. NIPAAm (Fig. 5) is a thermosensitive polymer that
exhibits a low critical solution temperature (LCST) of 32 °C,
below which the polymer is soluble but precipitates if the
temperature is raised above the LCST . By incorporating a
hydrophilic titrable monomer (e.g., MAA), it is possible to
increase the LCST of NIPAAm and make the polymer pH-
sensitive . Polymeric micelles based on poly(L-lactide)-b-
PLLA) ABA triblock copolymers and diblock copolymers
(PEOz-PLLA) have been successfully investigated for tumor
targeting of DOX . Polymeric micelles consisting of
random copolymers of NIPAAm, methacrylic acid, and
octadecyl acrylate , also with adding an additional
hydrophilic monomer, N-vinyl-2-pyrrolidone (VP) decrease
the mononuclear phagocyte system (MPS) uptake and improve
accumulation in tumors . Anthracycline antibiotics, DOX
and daunorubicin are known to form a dimer in an aqueous
solution due to π–π stacking of their planar aromatic ring .
Polymeric micelles were prepared from poly(ethylene glycol)-
poly(beta-benzyl-L-aspartate) block copolymer (PEG-PBLA),
in which DOX was converted partially into its dimeric form
during the drug loading . These polymeric micelles
exhibited two-stage release profile, an initial rapid release was
followed by a state of slow and prolonged release of DOX.
DOX release was accelerated by decreasing the surrounding pH
from 7.4 to 5.0, indicating a pH-sensitive release of DOX from
the micelles. This rapid release may be due to the protonation of
the 3′-NH2group as well as an acid-catalyzed cleavage of the
azomethine bond in DOX–DOX.
196S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
3.5. pH-responsive dendrimers
High molecular weight polymers that incorporated with
anticancer drugs can significantly improve tumor targeting of
drugs due to the EPR effect. However, the availability of well-
defined water-soluble polymers with uniform dispersities and
that are nontoxic and biocompatible is rather limited. Dendritic
polymers are promising polymeric drug carriers due to their
well-defined molecular architecture. Dendrimers are synthetic,
spherical, highly branched, macromolecules of nanometer
dimensions developed as a result of pioneering work of Tomalia
et al. .Theyhave excellent control over polydispersity,and
the ability to display a high surface functionality along with
water solubility is one of the factors that makes dendrimers
attractive for drug delivery and biomedical applications .
Dendrimers are capable for use in many applications mainly due
to incorporating a variety of physical properties in its structures
through functional group modifications at the core, branches
and the periphery. Imparting a stimuli-sensitive property into
dendrimers could significantly expand the scope of these
macro-molecules in biomedical and drug delivery applications.
The synthetic dendritic polyester systems composed based
on the monomer unit 2,2-bis(hydroxymethyl)propanoic acid are
possible versatile drug carriers . DOX was attached via a
pH-sensitive linkage to these carriers. This demonstrated the
feasibility of using these polyester dendritic polymers to prepare
a viable polymer–drug conjugate. The dendritic polymer can be
attached to the DOX through several functional groups that are
available with the DOX. The hydrolysable amide linkage with
the polymer may be too stable toward acid-catalyzed hydro-
lysis. On the other hand, the keto group of DOX can be used to
form an acid labile hydrazone linkage, the resulting compound
showed excellent water solubility and making it an excellent
candidate for further biological evaluation .
Recently, a new approach has been investigated using PEO–
dendrimer hybrids as backbones for acid-sensitive micelles .
the periphery of the core-forming dendrimer block using an acid-
sensitive acetal linkage. It loses hydrophobic groups upon hydro-
lysis of the linkage, the core-forming block becomes hydrophilic,
thus destabilizing the micelle and allowing the release of the drug
from its encapsulating micellar compartment. The stepwise
synthesis of the PEO–dendrimer backbone allows a high degree
of control over the polymer structures. This in turn controls the
properties such as the rate of micelle disruption, the critical micelle
been used to deliver the DOX for tumor targeting .
A summary of the different types of temperature-responsive
nanocarriers is shown in Table 2.
4.1. Temperature-responsive polymeric nanocarriers
Temperature-sensitivity is one of the most interesting
characteristics in stimulus-responsive polymeric nanocarriers
and has been extensively investigated to exploit the hyperther-
mia condition for drug and gene delivery [10,136]. A
thermosensitive polymer displays a lower critical solution
temperature (LCST) in aqueous solution, below which the
polymers are water soluble and above which they become
water-insoluble. This property has been exploited in targeted
delivery of anticancer drugs. For example, the rhodamine–poly
(N-isopropyl acrylamide-co-acrylamide) conjugates were selec-
tively accumulated in a tumor tissue using targeted hyperther-
mia .Thermosensitive amphiphilic polymers generally have
temperature-responsive hydrophilic segments and a suitable
hydrophobic segment. NIPAAm and its random copolymers are
the most intensively investigated temperature-sensitive hydro-
philic segments (Fig. 5) . Block copolymers of PEG as a
hydrophilic block and NIPAAm or poly(N-isopropylacryla-
mide)-co-N-(2-hydroxypropyl) methacrylamide-dilactate as a
thermosensitive block are able to self-assemble in water into
temperature-responsive nanocarriers above the LCST of the
thermosensitive block . An amphiphilic thermosensitive
nanocarrier was prepared from N-(2-hydroxypropyl) methacry-
lamide lactate and PEG. This system showed to be a promising
delivery system for the parenteral administration of paclitaxel
. An interesting review on triggered destabilization of
polymeric micelles and vesicles bychanging polymer polarity is
available for further reading, see Ref. . The hydrophobic
segments, poly(L-lactide), cholic acid, alkyl, and poly(γ-benzyl
L-glutamate) also have been used in diblock polymers with the
temperature-sensitive polyacrylamide derivatives being the
hydrophilic segments. Indomethacin incorporated in block
copolymeric nanospheres was prepared from poloxamer and
poly(epsilon-caprolactone) (PCL). These exhibited the rever-
sible change of size depending on the temperature and were able
to reduce the damage compared with the free indomethacin
when evaluated by MTT assay . In another interesting
work, gold nanoparticles were used to prepare shell cross-linked
Pluronic® (poloxamer) micelles that exhibit a reversibly
thermosensitive swelling/shrinking behavior. This property of
the micelles was caused by hydrophobic interactions of cross-
linked or grafted poloxamer copolymer chains in the micelle
structure with raise in temperature .
Polymeric micelles have been explored for temperature
induced release of actives for drug and gene delivery. Tempera-
ture-sensitive micelles can be formed as a result from the
assembly of block copolymers composed of a temperature-
property is possessed by the outer shell of the polymeric micelles
and the drug molecules are incorporated into the hydrophobic
inner core. NIPAAm is known to exhibit phase-transition
temperature at 32 °C termed as LCST . It shows a very
sharp change in hydrophilicity/hydrophobicity in a very narrow
temperature region around 32 °C. In addition, LCSTof NIPAAm
can be easily modified to get above 37 °C using copolymers with
varying hydrophilicity or hydrophobicity. PEG is the commonly
as well as for the coating of other colloidal nanocarriers, because
of its biocompatibility and good “stealth” properties . The
197S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
for temperature-sensitive drug release.Bothinner cores and outer
shell polymer chemistries were investigated to modify the
temperature-responsive behavior of micelles for specific drug
delivery . AB type block copolymers consisting of a
micellar structures below the NIPAAm LCST. The inner
hydrophobic core can be loaded with the water-insoluble drugs,
while the NIPAAm outer shell plays the role of temperature-
responsiveness and aqueous solubilization. Introduction of an
amino group to the NIPAAm chain raises the LCST and slows
down the rate of the phase transition . Cholesteryl end-
capped temperature-sensitive amphiphilic polymers were synthe-
sized from the hydroxyl-terminated random poly(N-isopropyla-
and cholic acid, conjugated with amine-terminated NIPAAm was
shown to have LCST 37.7–38.2 °C and 31.5 °C respectively
. In another study, amphiphilic NIPAAm-grafted-polypho-
sphazene (NIPAAm-g-PPP) was synthesized by stepwise co-
terminated NIPAAm oligomers and ethyl glycinate (GlyEt), the
LCSTwas found to be 30 °C in water.
Non-viral vectors based on thermosensitive polymers have
been proposed for effective in vitro or in vivo gene transfection.
Piskin and colleagues investigated different types of cationic PEI
and their block copolymers with NIPAAm as temperature-
sensitive carriers for in vitro and in vivo transfection of plasmid
DNA [144,145]. In this study , introduction of PEI units to
the NIPAAm chains increased the LCST values up to 37 °C. A
green fluorescent protein expressing plasmid to transfect Hela
cells in cell culture media was complexed to PEIs and
copolymers. Cytotoxicity was evident with PEIs especially with
the branched higher molecular weights. Copolymerization
reduced the cytotoxicity. The effective gene expression without
any significant toxicity was achieved with the complex prepared
plasmid DNA from polycationic polymers in response to
temperature stimuli and enhanced in vivo transfection efficiency
. Elegantly discussed reviews on gene therapy with main
emphasis on polycationic stimuli-responsive carriers as non-viral
gene vectors are available from Refs. [146–148]. Twaites et al.
 reported DNA binding behavior of pH and temperature-
responsive NIPAAm copolymers. Plasmid DNA complexed to
NIPAAm copolymers displayed variation in gel retardation
behavior above and below polymer phase-transition tempera-
tures. High molecular weight NIPAAm copolymer forming
complexes with reduced affinity above LCST where a branched
PEI–NIPAAm bound with higher affinity above NIPAAm phase
transition. However, PEI–NIPAAm conjugate showed low level
transfection efficiency. Cell viability studies showed that the
complexed to DNA or at certain DNA:polymer ratio . PEI
copolymers with side chain grafted NIPAAm were shown to be
less toxic than PEI alone or NIPAAm copolymer and the effects
were concentration dependent .
Illustrative examples of temperature- and redox-responsive nanocarriers for drug and gene delivery
Nanocarrier typeStimuli-responsive polymer or lipid Drug or geneProperty Therapeutic outcomeRef.
PaclitaxelRapid sol-to-gel transition with
change in temperature
Showed good antitumor effect
in melanoma bearing mice
Doxorubicin Temperature-sensitivity was
observed at 42 °C
Showed enhanced activity
against Lewis lung
DoxorubicinEnhanced drug release in
response to temperature fluctuation
Low temperature-sensitive liposomes
that triggers complete release in 39–40 °C
Complete drug release was
achieved at 40 °C
DoxorubicinIncreased therapeutic efficacy in
FaDu human tumor xenografts
vesicles bearing poly
Poloxamer F127 containing liposomes
90% release was achieved at 42 °CShowed 2.5 folds increase
in fluorescence in CT-26
tumor bearing mice
Effective tumor therapy
Effective intracellular deliver of pGFP
than the non-pH-sensitive
TAT-modified PEGylated liposomes
Plasmid DNA Thiopolyplexes releases DNA in
Plasmid DNA DNA release depends on the redox state
Thiopolycation PESCEfficient gene transfection
Redox-sensitiveCationic lipoic acid Increased several folds
of transgene expression
Enhanced gene transfection
Poly(ethylene glycol)-modified thiolated
Interconvertible dihydripyridine to
pyridinium salt lipoidal carrier
Plasmid DNA Intracellular DNA delivery in response
to glutathione redox environment
EstradiolThis redox system traps inside the brain
and undergoes subsequent hydrolysis
and release of estradiol
Effective in treating the
198S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
4.2. Temperature-responsive liposomes
Temperature-sensitive liposomes that show the response to the
temperature above physiological temperature have been studied
intensively to achieve targeted-drug delivery, since temperature-
sensitive liposomes can release their contents to the target site
wheretheheat is applied. It isa wellrecognizedfactthat the
membranes of various types of phospholipids undergo phase
transitions, such as gel-to-liquid crystalline and a lamellar-to-
hexagonal transition and become highly leaky to small water-
sensitive liposomes using dipalmitoylphosphatidylcholine
(DPPC) as the primary lipid . These DPPC liposomes
become leaky to small water-soluble molecules at a gel-to-liquid
crystalline phase transition. The gel-to-liquid crystalline transition
of DPPC membrane takes place at the clinically achievable
temperature of 41 °C. Adding small amount of distearoylpho-
sphatidylcholineasa co-lipidtotheliposomal membranecanhelp
in adjusting the desired transition temperature. Recently, tempera-
ture-sensitization of liposomes has been employed using thermo-
sensitive polymers  which exhibits a LCST . At the
transition as the temperature changes. Attachment of these
temperature-sensitive polymers to the liposomes can provide the
liposomes with temperature controlled functionalities. For exam-
ple, highly hydrated polymer chains attached to the liposomes
stabilize the liposomes below the LCST, but above the LCST, the
resulting in release of its contents . As a consequence of
changeintemperature dependent alteration of hydrophobicityand
conformation of the polymer chains, the surface properties of
liposomes can change and which exhibit temperature controlled
fusion or affinity to cells . Therefore in addition to targeted
delivery by temperature-sensitive drug release, the modified
liposomes using temperature-sensitive polymer may offer another
strategy for targeted-drug delivery, which is targeting by the
temperature induced control of interaction with cells.
Various researchers have investigated the modification of
liposomes with NIPAAm copolymers to obtain liposomes with
temperature-sensitivity . An interesting review on thermo-
sensitive polymer-modified liposomes is available for further
reading . Han et al.  investigated the surface
modification of liposomes by using poly(N-isopropylacryla-
mide-co-acrylamide) (NIPAAm-AAM) and PEG. The release of
DOX from the NIPAAm-AAM/PEG modified liposomes was
increased around the transition temperature of the polymer. In
addition, modified liposomes were found to be stable in the
serum compared with unmodified liposomes suggesting that
NIPAAm-AAM/PEG modified liposomes are suitable for
targeted-drug delivery. Chandaroy et al.  have studied
the temperature-sensitive di-oleoylphosphatidylcholine lipo-
somes containing Pluronic® F127 (P-127) molecules where
the P-127 interacts with the liposomal lipid bilayer at elevated
temperature and causes the release of its encapsulated
fluorescent markers. Concentration of the P-127 is found to
be critical in destabilizing the liposomal membrane and release
at precise temperature. Authors also further proved that the
stealth liposomes using di-stearoyl(polyethylene glycol 5000)
phosphatidylethanolamine (PEG5000DSPE) in conjunction with
P-127 showed similar effect on content release in comparison to
the non-stealth liposomes.
5. Illustrative examples of redox-responsive nanocarriers
A summary of the different types of redox-responsive
nanocarriers is shown in Table 2.
5.1. Disulfide cross-linked polymeric nanocarriers
spaces could provide an opportunity for programmed delivery of
drugs and genes. A drug or gene molecule can be entrapped or
encapsulated in a nanocarrier that is held together by disulfide
bonds. Once the disulfide bonds of nanocarrier are reduced in the
presence of low reducing potential due to an excess of reduced
glutathione (GSH) inside the cell, the drug or gene present in the
nanocarrier is released. One of the features required for a
successful gene-delivery system is a high plasma stability that
overcomes DNA release before reaching the target cells .
This can be achieved by incorporating the covalent linkages
between polymer chains of polymeric nanocarrier, i.e. intramo-
lecular and intermolecular cross-linking. This could stabilize the
polymeric nanocarriersand inhibits dissociation and early release
of DNA . Cavallaro et al.  have prepared a
polyaspartamide polymeric carriers for DNA delivery based on
polycation strategy. The positively charged groups introduced on
the polymer backbone for electrostatic interactions with DNA,
and thiol groups for the formation of disulfide bridge between
polymer chains. This resulted in formation of reversibly stable
thiopolyplexes. By incorporating disulfides between polymer
chains, it is expected to give polymeric nanocarrier systems of
environment-sensible properties, so that while dissociation of
of low reducing potential due to an excess of reduced glutathione
inside the cell, should lead to selective intracellular DNA release
. Recently, FDA has approved the anti-CD33 antibody-
conjugate optimized for targeting leukemic cells (Mylotarg®),
this consists of two cleavable sites in the linker, a disulfide bond
and an acylhydrazone bond for treatment of acute myeloid
leukemia . An interesting review article that focuses on
biological fate of disulfide bonds is available .
of plasmid DNA using thiolated gelatin nanoparticles. Gelatin
thiopolyplexes can release the DNA in the highly reducing
environment, such as in response to glutathione. Thiolated
gelatin was prepared by covalent modification of the primary
amino groups of type B gelatin using 2-iminothiolane (Traut's
plasmid DNA transfection in NIH-3T3 murine fibroblast cells
for enhanced green fluorescent protein (EGFP-N1) was done by
fluorescence confocal microscopy and fluorescence-activated
cell sorting (FACS). Qualitative results showed highly efficient
expression of GFP that remained stable for up to 96 h.
Quantitative results from FACS showed that the thiolated
199S. Ganta et al. / Journal of Controlled Release 126 (2008) 187–204
gelatin nanoparticles (SHGel-20) were significantly more
examined. The results of this study suggest that thiolated gelatin
nanoparticles would serve as a biocompatible intracellular
delivery system that can release the payloadin a highly reducing
environment. Recently, we have proved that a long circulating
PEG-modified thiolated gelatin nanoparticles could also deliver
glutathione concentrations [8,29]. These nanocarriers incorpo-
rated with plasmid DNA encoding for soluble vascular
endothelial growth factor receptor 1 (sVEGFR-1 or sFlt-1)
cancer cells in in vivo tumor model. In addition, the expressed
sFlt-1 was very effective in suppressing tumor growth and
Carlisle et al.  prepared nanoparticles of plasmid DNA
condensed with thiolated PEI, and these were coated with thiol-
reactive poly[N-(2-hydroxypropyl)methacrylamide] (PHPMA)
with 2-pyridyldisulfanyl or maleimide groups, forming reducible
disulfide-linked or stable thioether-linked coatings. Disulfide-
linked complexes showed 40–100 fold higher transfection
efficiency than thioether-linked ones. This is because of the
reduction with dithiothreitol (DTI) that allowed complete release
of DNA from disulfide-linked coated complexes. Transfection
efficiency was further improved by boosting of intracellular
glutathione using glutathione monoethyl ester or decreased using
buthionine sulfoximine. A glutathione-sensitive cross-linked PEI
gene carrier was prepared using dimethyl 3,3′-dithiopropionimi-
date dihydrochloride as a cross-linking agent . Thiol groups
prone to oxidation were immobilized on the polymeric backbone
of chitosan in order to incorporate the property of extracellular
stability and intracellular gene release by forming reversible
disulfide bonds . Chitosan–thiobutylamidine conjugate,
exhibiting 299.1+/−11.5 μmol of free thiol groups per gram
polymer, formed coacervates with pDNA and the highest
efficiency was observed in transfection studies performed in a
Caco-2 cell culture for chitosan–thiobutylamidine–DNA nano-
particles thatidentifieschitosan–thiobutylamidineasa promising
new vector for gene delivery.
5.2. Disulfide cross-linked liposomes
Disulfides (–S–S–) have found use in liposomal targeted
delivery approaches. Disulfide bonds can be used as linkers for
targeting conjugates and also to prepare lipids with disulfide
Upon reaching intracellular spaces, the thiolated liposomes
destabilize in response to glutathione, as this destabilizing effect
is attributed to the reduction of disulfide bridges of liposomes.
As a consequence, the active component encapsulated in
liposomes is released intracellularly and this intracellular
delivery of genes is necessary for efficient transfection.
Disulfide mediated redox stimuli-responsive liposomes are
prepared using standard phospholipids and a small lipid of
which the hydrophobic and hydrophilic parts are linked through
a disulfide bond. Such liposomes exhibited stability until
reaching to a reducing environment that cleaves the disulfide
bonds, disrupting the liposomal membrane and releasing the
liposomal contents . Thiocholesterol-based cationic lipids
(TCL) were synthesized which can be used into liposomes to
encapsulate DNA . The resulting lipoplexes are shown to
release their content in the presence of low concentrations of
reducing agents . Redox-sensitive liposomes with long
circulating property were reported by Kirpotin et al. . In
which liposomes were incorporated with detachable disulfide-
linked PEG polymer coating, whereby it is expected to show
longer circulation, therefore enhanced accumulation at tumor
site, and release of its contents into target cells in response to
redox stimulus. Stearically stabilized mPEF-DTP-DSPE ant-
CD19 liposomes were prepared to deliver the DOX into B-
lymphoma cell cytoplasm . This demonstrated the modest
increase in therapeutic activity in vivo. Lipid based mitomycin
C conjugates with cleavable disulfide linkage were prepared
and encapsulated in PEGylated liposomes . These
liposomes are found to be less toxic and superior in therapeutic
activity compared to free mitomycin C.
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