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392 N. Nomikou et al.
https://t.me/med1917
minimal (Zhang et al. 2017). Drug molecules can also be attached to polymer chains
via enzyme-cleavable linkers to create amphiphilic polymer-drug conjugates that can
self-assemble into micelles. Specific enzymes have the capacity to break the conjugation, destabilizing the micelles, and liberating the drug. For example, Wang et al.
(
2020) conjugated doxorubicin to Angelica sinensis polysaccharide via an MMP-2
sensitive peptide to form amphiphilic conjugates that self-assemble into micelles.
The presence of the enzyme increased drug release by 90%, compared to the
enzyme-free environment.
(c) Enzyme-sensitive “gateke epers”.
Mesoporous silica nanoparticles (MSNs) are becoming an extremely popular
alternative to organic enzyme-sensitive nanoparticles. In these systems, therapeutic
agents can be loaded into the porous structures and can be trapped by capping the
pores. These caps, also called “gatekeepers”, can be engineered to release or degrade
in the presence of specific enzymes. For instance, Eskandari et al. (
MMP-2-responsive MSNs capped with gold nanoparticles (AuNPs) for the entrapment of doxorubicin. The AuNP “gatekeepers” were attached to the surface of the
MSNs via an MMP-2-sensitive peptide, which was conjugated onto the MSNs
through an amidation reaction with carboxyl groups that were pre-functionalized
onto the MSN surface. In the presence of MMP-2, 90% of the entrapped doxorub icin
was released in 8 h, while release was minimal in the absence of the enzyme. Iron
oxide nanoparticles have also been used as “gatekeeping” moieties to trap therapeutic agents within MSN pores. When the gatekeeping/capping is stabilized via
macromolecular linkers containing trypt ophan residues, it becomes breakable by
indoleamine 2,3-dioxygenase 1, an enzyme that is present in the tum our microenvironment and can metabolize tryptophan to N-formylkynurenine, causing drug
release selectively in the tumour mass (Qiao et al. 2019).
2019) described
(d) Enzyme-triggered size change.
Nanoparticle-based drug delivery in solid tumours significantly relies on the
enhanced permeability and retention (EPR) effect. The EPR effect is intensely operational when nanoparticles of sizes 100 – 200 nm are injected. These nanoparticles
are too large to be cleared by the kidneys or escape from normal vasculature to
healthy tissues, but they can permeate the dysfunctional tumour vasculature and be
retained in the tumour mass. However, upon extravasation, nanoparticles of this size
range show poor diffusion due to the dense nature of tumour tissues and they lack the
penetrative ability of smaller nanoparticles. Nanoparticulate drug carriers have been
designed to undergo enzyme-triggered size reduction in the tumour microenvironment, aiming to improve diffusion of the therapeutic system throughout the tumour
mass and potentially improve cellular uptake (Hadi et al.
2020) described the assembly of nanoclusters using folate functionalized
et al. (
albumins and carboxylesterase-sensitive tetrachloroperylene monoamide which,
under the action of tumour-expressed carboxylesterase, can break down from
100 nm nanoclusters into 10 nm nanoparticles, with improved diffusion capability.
2021). For example, Cai
15 Stimulus-Responsive Nanoparticles for Drug Delivery 393
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Alternatively, in terms of enzyme-induced size-shifting, it may be desirable to
have nanoparticles that aggregate to form larger structures. For mucosal drug
delivery, nanoparticles that can penetrate through the mucus layer, can also backdiffuse into outer mucus regions, reducing their accumulation. In order to prevent
back-diffusion, a solid–lipid nanocarrier (120 nm), formed with a phosphate ester
surfactant and octadecylamine, was developed, to penetrate through the mucus and
form aggregates in the presence of alkaline phosphatase, which is a cell membranebound enzyme (Le-Vinh et al. 2021). Microaggregates of 5–8 mm were formed due
to particle electrostatic interactions induced by the enzymatic phosphate removal,
after approx. 30 min of exposure to the enzyme. This innovative system may be used
for mucosal drug delivery to alkaline phosphatase-ex pressing tissues, such as the
colon, cervix, vagina, lung, and mucus-secreting tumours.
(e) Enzyme-driven charge shift.
Finally, enzymes can be used to degrade away negatively charged coatings that
protect nanocarriers in blood circulation, to expose the positive charge underneath
the coating and enhance cellular uptake or promote lysosomal/endosomal escape. He
et al. (
containing hyaluronic acid shell. These nanoparticles undergo an enzyme-dependent
charge shift in the presence of overexpressed hyaluronidase. This is an enzyme that
is abundant both in the extracellular tumour microenvironment and in the cytoplasm
of tumour cells. In the corresponding study, it was found that in the extracellular
hyaluronidase-rich environment, the coating of the nanoparticles was partially
degraded, and the charge shifts from -12 mV to neutral. This shift, along with the
active targeting ability of hyaluronic acid, promotes nanoparticle cellular uptake.
Once internalized, the intracellular action of hyaluronidase further degrades the
coating, increasing the charge to +39 mV, while releasing more doxorubicin.
Moreover, this drastic increase in surface positive charge accelerates the lysosomal/endosomal escape of doxorubicin into the cytoplasm (Fig.
2021) fabricated polymeric nanoparticles coated with a doxorubicin-
15.1).
15.2.2 pH-Responsive Nanoparticles
A common hallmark in pathology is an abnormal level of acidity in the extracellular
environment of the affected tissue(s). Cancer, inflammatory disease, ischemia, as
well as stroke are a few examples of where this phenomenon is known to occur. The
pH of the tumour microenvironment, for instance, is around pH 6.4–pH 6.8, while
blood and healthy tissue, however, present neutral environments (pH 7.0–pH 7.4).
There are also disparities in pH within individual cells themselves. For example, the
pH of the cytosol is between pH 7.0 and pH 7.4, while the pH of lysosomal/
endosomal compartments varies from pH 4.0 to pH 6.0 (Mi 2020).
These wide pH differentials can be exploited in nanomedicine to achieve targeted
therapeutic activity, at either the tissue or cellular level. pH-responsive
nanomedicines commonly utilize (a) acid-labile linkers, (b) pH-dependent charge
394 N. Nomikou et al.
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Fig. 15.1 Enzyme-triggered charge shift can promote both cellular uptake and lysosomal escape of
doxorubicin (DOX) (He et al.
2021)
shifts, and (c) pH-stimulated hydrophilic to hydrophobic trans itions. Another interesting strategy is the employment of (d) pH (low) insertion peptide (pHLIP) technology (Reshetnyak et al.
2020).
(a) Acid-labile linkers.
Acid-labile bonds or linkages can be used to directly conjugate drugs onto
nanocarriers, or they can be integrated into the structure of a nanocarrier itself. In
the former, acidic environments can promote the break of pH-sensitive bonds,
resulting in the release of drug molecules. In the latter, acidic conditions can trigger
the disassembly or degradation of the nanocarrier to release the encapsulated therapeutic agent(s). To date, multiple pH-labile bonds have been reported. These
include 2,3-dimethylmaleic anhydride linkages, imine, benzoic-imine, ketal, ortho
ester, acetal, cyclic acetal, citraconic amide, b-thiopropionate oxime, and hydrazone
bonds (Zhuo et al.
2020; Mi
2020
).
Hydrazone bonds can readily undergo hydrolysis under acidic conditions, a
reaction thought to be initiated by the protonation of the imine nitrogen (Kölmel
and Kool
2017). The most common means of forming hydrazone bonds involves a
condensation reaction between hydrazine and an aldehyde or ketone group
(Sonawane et al. 2017). For example, Gatti et al. (2018) exploited the ketone
group on doxorubicin to form a hydrazone conjugated prodrug with poly(lactic
acid). The polymeric micelles formed from the prodrug and a block copolymer
surfactant were found to release over two-fold more free doxorubicin in acidic
conditions, when compared to the release at physiological pH, over a 2-day period.
15 Stimulus-Responsive Nanoparticles for Drug Delivery 395
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However, direct drug conjugation using hydrazone bonds is quite limiting since
there are not many drugs that contain aldehyde or ketone functional groups
(Sonawane et al. 2017). To overcome this limitation, hydrazone bonds can solely
be used to build pH-sensitive polymeric nanoparticles for carrying therapeutic
agents, hence, avoiding any drug conjugation. Yang et al. (
with paclitaxel-loaded polymeric micelles by incorporating a hydrazone bond
between the hydrophilic and hydrophobic portions of an amphiphilic polymer
capable of self-assembling into micelles. In acidic pH, the hydrazone linkage breaks,
destabilizing the micelles and releasing the encapsulated paclitaxel.
(b) pH-dependent charge shifts
Nanoparticles can be engineered to carry a neutral or negative charge in pH 7.0–
pH 7.4 environments, and a positive charge in acidic environments. This would
allow nanoparticles to remain stable in circulation, while promoting cellular uptake
in acidic conditions or stimulating lysosomal/endosomal escape. Nanoparticles can
be equipped with a pH-triggered charge reversal ability in a few different ways. One
way is the breakage of pH-sensitive bonds such as the b-carboxylic amide bond,
which can be incorporated via the addition of various anhydrides, such as cis-4cyclohexene-1,2-dicarboxinic anhydride, carboxy dimethylmaleic anhydride, cisaconitic anhydride, 2,3-dimethylmaleic anhydride, and citraconic anhydride
(Zhang et al. 2022;Mi 2020). Hu et al. (2020) was inspired by this technique and
used a pH-sensitive polymer containing 2,3-dimethylmaleic anhydride (DA), which
was used to form a nanocarrier of two prodrugs, a-cyclodextrin (a-CD)-nitric oxide
and a-CD-chlorine e6. In physiological pH, the b-carboxylic amide bond of the
polymer is stable, the DA group remains in place and the nanoparticles have a
negative charge due to the exposed carboxyl groups. Upon entering the acidic pH
environment of bacterial biofilms, the amide bond is hydrolysed leading to the
exposure of protonated amino groups and a positive charge. This charge switch
allows deeper penetration of the nanocarriers into biofilms and better interaction with
the negatively charged bacterial surface.
Protonation is another way of increasing positive charge. It is possible for anionic
polymers containing carboxyl or sulfonyl groups to be deprotonated at physiological
pH and protonated when exposed to increased acidity (Zhang et al. 2022). For
example, Lin et al. (2018) functionalized nanoparticles with histamine, which
contains imidazole groups that can become protonated in acidic pH (pKa = 6.95),
leading to a neutral to positive surface charge transition.
2019) attempted this
(c) pH-dependent hydrophobicit y.
The environmental pH can have an impact on a polymer’s hydrophobicity. For
instance, at low pH, protonation can cause a polymer to switch from a hydrophobic
to a hydrophilic polymer. Such behaviour can be used to disrupt nanosystems and
cause them to release their payload under specific pH conditions, as shown by Wang
et al. (2019a, b). For example, when drug-containing nanoparticles were prepared
using two amphiphilic polymers, and the pH-responsive polymer poly(ethylene
glycol)-block-poly diisopropylaminoethyl methacrylate-block-poly dopamine,
396 N. Nomikou et al.
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Fig. 15.2 (a) Steps for the formation of a pH-responsive nanoparticle decorated with pHLIPs
(Palanikumar et al.
tured peptide. But as acidity increases, the peptide begins to spontaneously fold into an alpha helical
shape and inserts itself into the lipid bilayer
2020), (b) In healthy tissue with neutral pH, the pHLIP resembles an unstruc-
protonation of the latter at acidic pH (pKa = 6.8) caused a shift from hydrophobic to
hydrophilic, thereby destabilizing the nanoparticles and resulting in drug release
(Albuquerque et al. 2021).
Hydrophobicity-based conformational changes in certain peptides can be induced
by pH changes. This phenomenon is exploited in the pH (Low) Insertion Peptide
(pHLIP) technology which enables specific peptides to be inserted into cell membranes in response to an acidic pH. pHLIPs predominantly contain protonatable,
negatively charged residues, and hydrophobic residues. When residues are protonated at low pH, the prior negative charge is countered and the hydrophobicity of the
overall peptide increases, enabling peptide insertion. Commonly used residues are
asparagine and glutamate and their pKas would not allow for protonation or
deprotonation to occur at physiological pH. The rate of protonation and
deprotonation is influenced by the dielectric environment, leading to a rise in
carboxyl group pKas near the cell membrane. As peptides dissolve further into the
membrane, the dielectric environment shifts to favour intrahelical hydrogen bonds.
Thus, the peptide undergoes a conformational change from an unstructured peptide
outside the cell membrane into an alpha helical structure spanning across the cell
membrane (Fig.
15.2). pHLIPs have been used recently by Palanikumar et al. (2020)
to enhance the cellular internalization of nanoparticles in tumours. The group
utilized polylactic-co-glycolic acid (PLGA) as a backbone for pH-responsive hybrid
15 Stimulus-Responsive Nanoparticles for Drug Delivery 397
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nanoparticles loaded with doxorubicin-triphenylphosphine. The PLGA centre is
enveloped in a cross-linked bovine serum albumin coating that is decorated with
pHLIPs. The pHLIPs contain glutamate residues (with a measured pKa of 6.5) that
are protonated to increase the hydrophobicity of the peptide causing its transition
into a transmembrane a-helix structure, which can be translocated across the lipid
bilayer. This membrane insertion allowed the nanoparticles t
more direct passage. Alternatively,
nanoparticles could internalize via clathrin-mediated endocytosis. Here, the mode
of action is thought to involve pHLIPs activation within the acidic endosomal pH
and this then drives endosomal escape (Nguyen et al.
it was found that a small proportion of the
o enter the cytosol via a
2019; Reshetnyak et al. 2020).
15.2.3 Glucose-Responsive Nanoparticles
Currently, diabetes is managed by regular insulin injections taken dail y, which is
associated with pain and complications (e.g. infection). On that basis, self-regulating
insulin delivery systems to deli ver precise doses of insulin in response to changes in
blood glucose levels are of significant value. To this end, many glucose-responsive
insulin delivery nanosystems have been developed for the treatment of diabetes.
Overall, these strategies can be broken down into two main categories: a) the use of
glucose oxidase (GOx) enzyme, and b) the use of phenyl boronic acid (PBA)containing polymers (Zhao et al. 2017).
(a) Targeting glucose using GOx.
GOx is the oxidoreductase enzyme that catalyses the conversion of b-D-glucose
into D-glucono-d-lactone, giving off hydrogen peroxide as a by-product. D-gluconod-lactone then undergoes non-enzymatic hydrolysis into D-gluconic acid (Kornecki
et al. 2020). GOx can be immobilized onto an insulin-loaded nanocarrier that
degrades in acidic pH, and in the presence of glucose, D-gluconic acid is produced,
dropping the pH, and relea sing insulin from the nanocarrier. This was well demonstrated by Jamwal et al. (
acryloyl crosslinked dextran dialdehyde nanoparticles. These nanoparticles were
able to show a modest increase in insulin release in response to glucose (Das et al.
2020). A similar result is achieved when GOx is co-loaded along with insulin and
catalase into pH-responsive acetalated dextran nanoparticles (Volpatti et al. 2020).
(b) Targeting glucose using PBA-containing materials.
Nanoparticles based on PBA-containing polymers are often preferred to proteinbased systems like GOx due to their improved stability in the body, long-term
storage stability, and better glucose-sensitivity (Das et al.
derivatives have high affinity for cis-diol-containing compounds and can bind
these through reversible 5- or 6-membered cyclic boronic ester formation. Thus,
PBA-based nanoparticles can be designed to interact with the cis-diols found in
glucose molecules. PBA moieties are Lewis acids, and in aqueous solution, they
2019), who immobilized GOx via Schiff base linkage onto
2020). PBA and its
398 N. Nomikou et al.
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exist in equilibrium between two forms: a relatively hydrophobic (i.e. neutral)
trigonal-planar form and a relatively hydrophilic negatively charged tetrahedral
boronate form. Cyclic boronic esters with cis-diols tend to occur with the hydrophilic
form because it pushes the ionization equilibrium towards the hydrophilic PBA
making it the more stable and favourable route. Then, upon binding with cis-diols,
the PBA-containing polymer is made more hydrophilic. Th
licity subseq
uently causes PBA-based nanoparticles to swell and eventually disas-
is change in hydrophi-
semble in response to glucose, releasing their payload. It is, however, important to
mention that PBA has a reported pKa of approximately 8–9, and only at a pH above
the pKa will the equilibrium shift to the hydrophilic form. Hence, at physi ological
pH = 7.4, the dominant form will actually be the hydrophobic counterpart. This
o t
explains why early research was limited t
sivene
respon
ss in alkaline conditions. It was only until researchers were able to
he demonstration of glucose-
find ways of lowering the pKa of PBA-based polymers that practical designs began
appearing in the scientific literature. Essentially, the pKa of PBA can be lowered by
the addition of electron-withdrawing groups, such as the a ddition of an amino group
adjacent to the PBA moiety or a halo, nitro, or carbonyl group into the phenyl ring of
PBA (Zhao et al.
2017; Das et al. 2020; Gaballa
and Theato
2019). For example,
FITC-insulin loaded polymeric micelles were developed using the PBA-containing
block copolymer, poly[(N-acryloylmorpholine-block-(N-acryloylmorpholine-coPBA)], and a cis-diol containing polymer. By the addition of 2 mg/mL glucose,
boronate ester bonds with glucose began to form causing the hydrophobic core to
swell, consequently increasing the size of the micelles from 80 nm to 190 nm and
releasing insulin. When glucose concentration was in
contai
diol
micelle collapse (Gaballa and Theato
ning polymer was sufficiently replaced by glucose to cause complete
2019).
creased t
o 5 mg/mL, the cis-
15.2.4 Redox-Responsive Nanoparticles
The use of redox-responsive drug delivery systems is a promising approach for the
targeted treatment of cancer. The reducing environment of tumours is regulated by
the reduction–oxidation states of NADPH/NADP
GSH refers specifically to the reduced form of glutathione, while GSSG is the
oxidized form that consists of two glutathione molecules linked via a disulfide
bond. The concentration range of intracellular GSH is between 0.5 and 10 mM,
where it is mainly found in the mitochondria and cytosol, while the extracellular
concentration range is 2–20 μM (Guo et al.
achieved through the structural design of the nanoparticle carrier (Wang et al.
2019a, b; Lu et al. 2021), the GSH-dependent activation of a prodrug to an active
drug (Luo et al. 2022, Li et al. 2020) or both (Ling et al. 2019). The functional
components of the nanoparticle carrier that interacts with GSH are the chemical
bonds that are cleaved when GSH levels reach a certain threshold, allowing for a
targeted release of their cargo (Gisbert-Garzaran and Vallet-Regi
+
and glutathione (GSH/GSSG).
2018). Targeting based on GSH can be
2021). Amongst
15 Stimulus-Responsive Nanoparticles for Drug Delivery 399
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the different types of redox-responsive bonds used, the disulfide (S–S) bonds are the
most extensivel y studied (Guo et al. 2018) and can be incorporated in both polymerbased (Wang et al. 2019a, b) and metallic (Lu et al. 2021
nanoparticles.
or can be used to modify the main polymer molecule by addition of side-chains that
are intended to mediate drug conjugation. The disulfide bonds can also be used to
directly conjugate therapeutic agents on to the surface of nanoparticles (particularly
for metallic nanoparticles). Polymers with disulfide within the backbone (connecting
individual repeating units of monomer
have
but
disulfide bonds linking their side-chains to drugs on the other hand are relatively
easier to manufacture and more stable (Gisbert-Garzaran and Vallet-Regi
Specially designed polymer-drug conjugates, where the drug is linked to the
polymer via disulfide bonds, or polymerization of small molecule drugs with
redox-responsive bonds can lead to the formation of nanoparticulate structures,
and the latter may have the capacity to accommodate additional therapeutic agents
(Luo et al.
consisting of doxorubicin linked to poly[N-(2-hydroxypropyl) methacrylamide] via
disulfide bonds, that formed micelles through self-assembly due to its amphiphilic
properties, allowing the encapsulation of a photosensitizer. These nanoparticles
degraded in the presence of GSH intratumourally, subsequently releasing doxorubicin and the photosensitizer. In another study by Li et al. (
agent 10-hydroxycamptothecin was polymerized with disulfide bonds among individual molecules to produce a polymer-prodrug or poly-prodrug, that was used as
the inner core of nanoparticles with an amphiphilic lipid-PEG shell, allowing for a
prolonged circulation in blood.
The application of platinum (Pt)-based anticancer drugs has faced many challenges due to their poor blood circulation, non-specific targeting, and the irreversible
deactivation (via reduction) induced by thiol-containing species (e.g. GSH) that are
abundant in the cytoplasm (Ling et al.
enhanced activity owing to molecular targeting (Johnstone et al. 2016
the aforementioned GSH-induced reduction renders them ineffective. This issue may
be addressed by substituting the Pt(II) species for Pt(IV) prodrugs, while incorporating these agents into nanoparticles improves blood circulation time and tumour
accumulation levels. As proof of this concept, a study has reported the formulation of
Pt(IV) prodrug-containing lipid-PEG nanoparticles via self-assembly. Upon intracellular disintegration of the nanoparticles and in the presence of GSH, the released
Pt(IV) prodrug is converted to the active Pt(II) metabolite, successfully inhibiting the
progression of cisplatin-resistant tumours (Fig.
the efficacy of GSH-sensitive nanoparticles, in terms of improving anticancer
activity and tumour targeting, would depend on the type of cancer, since GSH levels
can greatly vary among different cancers.
The disulfide bonds can be inserted in the backbone of the polymer
s v
ia such bonds) are highly sensitive to GSH
poor stability, as well as being complex to synthesize. Polymers with
2022; Li et al. 2020). Luo et al. (2022) reported a polymeric prodrug
2019). Pt(II)-based substances allow for
15.3) (Ling et al.
) drug-carrying
2021).
2020), the anti-cancer
); however,
2019). Importantly,
400 N. Nomikou et al.
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Fig. 15.3 Pt(IV) prodrug-carrying lipid-PEG NPs for targeted treatment of cancer cells. Pt
(IV) prodrugs are prepared by oxidation of cisplatin [Pt(II)-based drug], followed by the addition
of super hydrophobic ligands to the intermediate. These are then self-assembled into nanoparticles
with lipid-PEG allowing efficient circulation in blood. At the tumour site, nanoparticle structure is
degraded into irregular-shaped debris in the presence of GSH and subsequently releases the Pt
(IV) prodrug. Pt(IV) undergoes GSH-induced reduction to its active form Pt(II) and triggers
apoptosis, as well as lowering the intracellular GSH levels
15.2.5 Reactive Oxygen Species (ROS)-Responsive
Nanoparticles
ROS play an important role in many biological mechanisms and intracellular
signalling. ROS levels can be elevated in certain pathological microenvironments,
such as inflammatory diseases, neurodegenerative diseases, and cancer.
ROS-responsive nanoparticles offer a more-specific targeting system for cancer
treatment compared to GSH-dependent redox responsive systems due to tumour
hypoxia, which results in much higher ROS levels in tumour cells (as high as
100 × 10
-6
M), compared to healthy tissues (~20 × 10
The elevated levels of ROS can be exploited by using ROS-responsive nanoparticles
that have oxidation-labile linkages/bonds embedded within the nanoparticle structure or on its surface. Examples of such linkages are boronic acids/esters, alkylene
sulphide, aminoacrylate, peroxalate ester, polyproline thioether, and thioketal
(Rinaldi et al. 2022; Liang and Liu 2016). Based on these moieties, ROS-i nduced
-9
M) (Xu et al. 2017a, b).
15 Stimulus-Responsive Nanoparticles for Drug Delivery 401
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drug release from nanocarriers is possible through nanocarrier destabilization/degradation and/or through carrier-drug linker cleavage.
A category of ROS-responsive nanocarriers utilize thioether-containing polymers, such as poly(ethylyene sulfide), which undergo phase transition from hydrophobic to hydrophilic states under oxidative environments. The ROS-induced phase
transition of the nanoparticle building block leads to nanoparticle disintegratio n with
subsequent payload release (Napoli et al.
2004). A similar effect was achieved when
poly (propylene sulfide) or an amphiphilic diblock copolymer of propylene sulfide
and N,N-dimethylacrylamide were used to form the nanoparticulate carriers (Gupta
2012; Poole et al. 2015).
et al.
Polymers containing selenium or tellurium have also been exploited for
ROS-induced drug delivery. ROS oxidation of monoselenium and monotellurium
compounds can lead to phase transition from hydrophobic to hydrophilic compounds, similarly to sulfur-containing polymers, and can be used for the production
of ROS-responsive drug nanocarriers (Ren et al.
2012; Wang et al. 2015). For
instance, doxorubicin-carrying self-assembled micelles, constructed using a hydrophobic monoselenide-containing block polymer and PEG, have been designed for
controlled drug release upon exposure to 0.1% H
(Ma et al. 2010). The “solu-
2O2
bility switch” strategy described above has also been applied using boronic estercontaining polymers in drug nanocarriers. In a non-“solubility switch” approach,
when aryl boronic ester is incorporated in polymers that form nanoparticles, the latter
can be degraded in response to ROS and release the payload (de Gracia et al.
2012).
Boronic acids/esters have previously been shown to be highly responsive to
, which plays an important role in the pathogenesis of chronic inflammation.
H
2O2
Hence, while the majority of studies involving ROS-responsive nanoparticles are on
cancer therapeutics, these systems offer exciting prospects for the diagnosis and
treatment of inflammatory vascular diseases, such as atherosclerosis. A recent study
by Tang et al. (
2021) developed a ROS-responsive biomimetic nanoparticle formu-
lation for the treatment of atherosclerosis. The co-polymer was synthesized by
conjugation of ROS-sensitive boronic ester onto dextran, which was used to encapsulate rapamycin.
Thioketal linkers are also known to be readily cleaved by ROS and are employed
in various drug delivery systems with potential applications in inflammatory disease
or cancer treatments. For instance, complexation of a cationic poly(amino thioketal)
and DNA has been used in gene delivery (Shim and Xia 2013). Alternatively, the
thioketal group can be used as a linker between the nanocarrier and the drug, to
enable ROS-induced drug release (Yuan et al. 2014).
Alternatively, ROS may be exploited to activate prodrugs, and similarly to
GSH-sensitive prodrugs, different types of ROS-sensitive prodrugs have been
reported. A representative example is the prodrugs formed by polymerizing individual monomers of a drug using ROS-responsive linkages to produce a “poly-prodrug”
that can then be coated with an appropriate polymer-based coating to enhance blood
circulation. In a proof of concept study, the anticancer agent mitoxantrone (MTO)
was polymerized using ROS-cleaved linkages, and the resulting polydrug formed
nanoparticles through self-assembly with lipid-PEG (Fig. 15.4) (Xu et al. 2017a, b).
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