Functionalization of silica nanoparticles for nucleic acid ... · nucleic acids, such as plasmid DNA (pDNA), small interfering RNA (siRNA), and antisense oligonucleotide (ASO) [22].
This document is posted to help you gain knowledge. Please leave a comment to let me know what you think about it! Share it to your friends and learn new things together.
Transcript
Nano Res
1
Functionalization of silica nanoparticles for nucleic
acid delivery
Rimpei Kamegawa1, Mitsuru Naito2, and Kanjiro Miyata1 ()
Nano Res., Just Accepted Manuscript • https://doi.org/10.1007/s12274-018-2116-7
metabolic stability of the nanoparticles. No adverse
side effects caused by the targeted SiNP were
apparent in any of the patients, and the SiNP was
observed in the tumor regions by positron emission
tomography imaging in some cases. Also, a
poly(ethylene glycol)‐modified (PEGylated)
mesoporous silica nanoparticle (MSN) with a
diameter of 50 nm was developed for the loading of
the anticancer drug doxorubicin (DOX), eliciting
efficient tumor shrinkage in human squamous
carcinoma xenograft model mice with minimal side
effects [3]. When the PEGylated MSN was
administered intravenously, 12% of the dose
accumulated effectively in the tumor lesion by
enhanced permeability and retention (EPR) effect [4,
5], as will be described below. These studies
demonstrate the great potential of SiNPs for in vivo
bioimaging and drug delivery applications via the
systemic route.
The excellent performances of SiNPs stem from
their unique features. SiNPs with controlled
monodispersity can be prepared by modified Stöber
methods or reverse microemulsion methods for
controlling the biodistribution after systemic
administration [6, 7]. The modified Stöber methods
allow for the preparation on a large scale of
monodispersed SiNPs with a diameter of 10 nm to 1
µm by changing the concentration of silica
precursors and catalysts, such as lysine or arginine
and ammonia, respectively (Fig. 1(a), (b)) [6, 8, 9]. On
the other hand, the reverse microemulsion methods
are advantageous for the preparation of sub‐100 nm
SiNPs and enable the silica coating of inorganic
nanoparticles, such as gold and iron oxide [10, 11].
The shape of SiNPs can also be tuned to generate
nanorods [12], hollow SiNPs [13], MSNs [14], and
dendritic SiNPs [15] using the appropriate templates
(Fig. 1(c)–(f)). Among these, porous architectures
possess a large surface area that allows for the
efficient loading of drugs and imaging agents [16, 17].
In addition, the composition of SiNPs can be
controlled by co‐condensation of silica precursors
with organosilane compounds to obtain the desired
functionalities. For bioimaging applications, organic
dyes can be incorporated into the silica matrix
through the reaction with organosilane compounds
without compromising the spectral characteristics
[18]. Although the SiNPs with a size of <7 nm are
eliminated from the body through renal excretion [2],
larger SiNPs may accumulate in the tissues/cells,
which could lead to accumulative toxicity. To
circumvent these drawbacks, SiNPs bearing cleavable
bonds or doped with inorganic ions have been
designed in order to ensure their degradation in the
body [19]. Furthermore, SiNPs have abundant silanol
groups on their surfaces, which can be further
modified with functional polymers for the active
targeting of specific tissue and cells and enhanced
environment responsiveness and biocompatibility
[20].
Although SiNPs have been mainly applied for the
delivery of low MW drugs [20, 21], their unique
characteristics are also appealing for the delivery of
nucleic acids, such as plasmid DNA (pDNA), small
interfering RNA (siRNA), and antisense
oligonucleotide (ASO) [22]. These nucleic‐acid‐based
drugs provide fundamental treatment modalities for
various intractable diseases, such as spinal muscular
atrophy and familial hypercholesterolemia. Nucleic
acids regulate the gene expression pattern
responsible for these diseases in a sequence‐specific
manner. However, when administered into a human
body, natural nucleic acids are readily degraded by
nucleases and eliminated from the body via renal
excretion. Furthermore, the negatively charged
macromolecular structures of nucleic acids hamper
significantly their cellular internalization because of
the electrostatic repulsion with the negatively
charged cytoplasmic membrane. Therefore, to ensure
an effective transfection into the target cells, the
incorporation of nucleic acids into appropriate
delivery systems is required. To this end, SiNPs are
promising candidates because of the aforementioned
characteristics. The surface of SiNPs can be readily
modified with cationic moieties through electrostatic
interactions for nucleic acid loading. In particular,
SiNPs can be designed to have large pores (e.g., 20
nm) for enhanced loading of small nucleic acids. In
addition, the silica surface can be functionalized with
environment‐responsive moieties, which allows for
the programmed release of nucleic acid payloads in
response to the intracellular environment or a
specific biosignal. Furthermore, SiNPs enable the
codelivery of nucleic acids with anticancer drugs,
generating a therapeutic synergy for cancer
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
3 Nano Res.
treatment.
As described above, SiNPs are attracting a great
deal of attention for nucleic acid delivery. To clearly
illustrate the strong potential of SiNPs for this
particular application, this review first describes the
biological mechanisms of various types of nucleic
acids and the functions required for their delivery.
Next, the review introduces a series of promising
SiNP‐based delivery systems for nucleic acid
delivery and codelivery with anticancer drugs or
probes. Finally, current challenges to address for
pharmaceutical implementation and potential target
diseases are discussed.
Figure 1. Transmission electron microscopy images of various types of SiNPs: nonporous SiNPs with a diameter of 19 nm (a) and 120 nm (b), nanorod (c), hollow SiNPs (d), MSN (e), and DSN (f). Adapted with permission from ref. [9], Copyright American Chemical Society, 2008 (a, b); ref. [12], Copyright American Chemical Society, 2011 (c); ref. [13] Copyright Wiley-VCH, 2006 (d); ref. [14], Copyright Wiley-VCH, 2009 (e); and ref. [16], Copyright Wiley-VCH, 2013 (f).
2 Nucleic acid delivery
2.1 Nucleic acids as potential drugs
Treatment of intractable diseases with nucleic acids
occurs via the regulation of the disease‐related gene
expression (Fig. 2). For example, pDNA, a coiled‐coil,
double‐stranded DNA with approximately 5,000 base
pairs (bp), expresses the encoded protein after being
delivered into the nucleus of target cells. Currently,
the clinical trial of beperminogene perplasmid, which
is a pDNA encoding the human hepatocyte growth
factor, has been completed for the treatment of
critical limb ischemia, and its approval application
for manufacture and market has been submitted in
Japan in 2018 [23].
Other nucleic acids smaller than pDNA have also
been presented as promising therapeutic candidates.
Thus, the double‐stranded RNA molecule siRNA,
which has 19–23 bp, has been shown to potently
suppress the target gene expression in a
sequence‐specific manner [24]. When introduced into
the cytoplasm, siRNA is bound to the RNA‐induced
silencing complex (RISC), which selectively cleaves
complementary mRNAs, thereby reducing efficiently
the protein expression. It is worth mentioning that
the discovery of this biological mechanism, which is
known as RNA interference (RNAi) [25], merited the
Nobel Prize in 2006. Currently, several clinical trial
programs regarding siRNA‐based drugs are ongoing
for treatments of various diseases, such as dry eye
mRNA in the nucleus, ASO can modulate the splicing
behavior of premature mRNA to alter the mature
mRNA (or exon) sequence, thereby improving the
therapeutic protein production [30]. To date,
Fomivirsen and Mipomersen, ASO drugs which are
based on the RNase H‐mediated gene silencing
mechanism without any delivery formulations, have
been approved by the US FDA. These approved
ASOs demonstrate the strong potential of nucleic
acid drugs for the treatment of intractable diseases.
Nevertheless, the target diseases are highly limited
because of the inherent biodistribution of naked
ASOs. Therefore, the use of appropriate delivery
systems for improving the bioavailability of ASOs is
still required.
Figure 2. Schematic illustration of the biological activities of pDNA, siRNA, miRNA, and ASO. pDNA induces expression of gene encoded in its sequence. siRNA, miRNA, and ASO suppress target gene expression in a sequence-specific manner through the degradation or hybridization of complementary (pre) mRNA.
2.2 Nucleic acid delivery to target sites
Since nucleic acids are instantaneously decomposed
by nucleases when administered into a body, they
clearly require protection from enzymatic attacks.
Two major approaches for the protection of nucleic
acids from degradation are available: chemical
modification of nucleic acids and their incorporation
into a delivery system. Thus, the stability against
nuclease can be dramatically improved by replacing
a nonbridging oxygen in the phosphate backbone
with sulfur (i.e., phosphorothioate modification) and
by methylation of the 2′‐OH group in the ribose ring
(i.e., 2′‐O‐methylation) [28]. All four approved ASO
drugs employ different chemical modifications [28].
Although chemical modifications are highly useful
for improving the stability of nucleic acids, certain
serious limitations can be encountered. Chemical
modification of nucleic acids may increase the
toxicity while decreasing the biological activity. In
addition, the chemical modification of pDNA is
relatively difficult because of the use of bacteria or
enzymatic systems, which may not recognize
unnatural modified nucleotides.
On the other hand, the incorporation of nucleic
acids into delivery systems can also help to their
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
5 Nano Res.
stabilization by blocking the access of nucleases.
More importantly, delivery systems can alter the
biodistribution or intracellular distribution of nucleic
acids, thereby affording an efficient delivery to the
target site (Fig. 3). In case of systemic administration,
the stable circulation of nucleic acids in the
bloodstream prior to extravasation into the target
organ/tissue is desired. Small nucleic acids, such as
siRNA and ASO, are rapidly eliminated from the
bloodstream through kidney filtration, which
exhibits a filtration threshold of ~8 nm in size (or ~40
kDa in MW) [29, 30]. Thus, the kidney excretion of
small nucleic acids can be avoided by their
incorporation into nanoparticles larger than 8 nm.
The extravasation of nucleic acids or
nucleic‐acid‐containing nanoparticles from the blood
to the organ/tissue is determined by the vascular
endothelial structure. The continuous endothelium in
muscle and skin substantially hampers the
extravasation of circulating nanoparticles larger than
10 nm. On the other hand, the discontinuous
endothelium in the liver and spleen, which has
significantly large defects (or pores) of several tens to
hundred nanometers, permits the distribution of 10–
100 nm‐sized nanoparticles from the blood to the
organs. This fact has prompted the development of a
variety of liver‐targeted nanoparticle delivery
systems, including LNPs [31]. An important finding
that is worth mentioning in this regard is that solid
tumors are also equipped with leaky vasculatures,
which allows the preferential accumulation in the
tumor tissues of nanoparticles of several tens of
nanometers in size. This phenomenon has been
termed enhanced permeability and retention (EPR)
effect [4, 5]. Most of the cancer‐targeted nanoparticles
rely on this effect. However, recent studies on cancer
pathophysiology suggest a considerable
heterogeneity related to the EPR effect. For instance,
the size threshold for tumor accumulation of
nanoparticles is reported to vary significantly
according to the tumor types, stages, and positions,
even in the same patient [5]. In addition to the size
effect, the nanoparticles (or delivery systems) can be
further functionalized to ensure the preferential
accumulation in the target organ/tissue by the
so‐called active targeting strategy [32]. Generally,
cells overexpress a specific protein or sugar chain on
their surface. Then, nanoparticles functionalized with
targeting ligands that have high affinity for such
protein or sugar chain bind preferentially to the
target cellular surface. Interestingly, the affinity or
avidity of nanoparticle delivery systems for target
cells can be amplified by the functionalization of such
nanoparticles with multiple ligands at high density,
which affords multivalent binding to the target
cellular surface [32].
After reaching a target cell, the delivery systems
need to be uptaken by the cells. Cellular uptake of
nanoparticles or macromolecules generally occurs via
the endocytosis pathway. Positively charged
nanoparticles can facilitate adsorptive endocytosis
through electrostatic interactions with the negatively
charged cellular surface. However, strongly cationic
nanoparticles are nonspecifically bound to negatively
charged proteins and sugar chains, leading to
unspecific cellular uptake or unexpected adverse side
effects, such as agglomeration in the blood. To
circumvent this problem, the attenuation of the
cationic surface charges in the nanoparticles via
modification with neutral and hydrophilic polymer
chains, such as poly(ethylene glycol) (PEG), has been
extensively explored [33]. The endocytosed
nanoparticles are sequestered by the endosome and
delivered to the perinuclear region along with the
microtubule [34]. Then, acidified late endosome fuses
with lysosome, and the endosomal contents,
including nanoparticles, undergo enzymatic
degradation. To avoid the lysosomal degradation, the
delivery systems must escape from the endosome to
the cytoplasm. Ultimately, the delivery systems have
to release the nucleic acid payload in the cytoplasm
(mRNA, siRNA, miRNA, and ASO) or the nucleus
(pDNA and ASO) so that the biological or therapeutic
functions of the nucleic acids can be realized.
| www.editorialmanager.com/nare/default.asp
6 Nano Res.
Figure 3. Schematic illustration of nucleic acid delivery. The delivery systems need to avoid nuclease degradation and renal excretion, extravasate from the bloodstream to tissues, be internalized by target cells avoiding the uptake by nontarget cells, escape from endosome, and ultimately release the nucleic acid payload.
3 Silica nanoparticles for nucleic acid
delivery
3.1 Loading of nucleic acids
3.1.1 Electrostatic interactions. One of the most
widely tested methods for loading nucleic acids to
SiNPs is the electrostatic interaction between
negatively charged nucleic acids and positively
modified SiNPs. SiNPs fabricated by Stöber or
reverse microemulsion methods have a negatively
charged surface derived from deprotonated silanol
groups in aqueous solution. Thus, their modification
with positively charged moieties, such as primary
amines, is required for the nucleic acid loading. One
of the simplest approaches to this end is surface
coating via electrostatic interactions with cationic
polymers, such as polyethylenimine (PEI) [35, 36]
and poly(L‐lysine) (PLL) [37]. In particular, PEI has
been widely used for the positive charging of SiNPs
because it simultaneously provides them with
endosome escapability based on the proton sponge
effects [38], as will be described in Section 3.3.
Although SiNPs are readily coated with cationic
polymers through electrostatic interactions, the
bound polymers may be detached from SiNPs in the
body fluid, which contains abundant competitive
charged macromolecules. To avoid this unwanted
detachment, covalent modifications have also been
performed by silane coupling of SiNPs with
(3‐aminopropyl)trimethoxy silane (APTMS) or
(3‐aminopropyl)triethoxy silane (APTES) [39–42] or
by amidation reactions of carboxylated SiNPs with
amine‐containing oligomers/polymers [43–45]. An
interesting alternative approach is the calcium ion
doping into SiNPs, which allows the coupling
between nucleic acid phosphates and calcium ions
embedded in the silica matrix [46]. It should be noted
that the nucleic acids adsorbed on SiNPs become
significantly tolerant to enzymatic degradation,
thereby enhancing the cellular uptake efficiency [35,
37, 39–42, 45, 46].
Since nucleic‐acid‐based drugs are incorporated into
SiNPs through electrostatic interaction, the loading
amount of drugs is mainly determined by their
surface area and shape, besides the strength of such
interaction. Thus, it seems reasonable to assume that
SiNPs with larger surface area should incorporate
drugs more efficiently. To increase the surface area,
various porous architectures have been designed [16,
17]. MSNs with ~3 nm‐sized pores are suitable for
incorporating low molecular weight drugs, such as
DOX with an MW of approximately 500 Da.
However, the loading of nucleic acids to such MSNs
is substantially hindered by the significantly larger
size of nucleic acids, e.g., ~3.3 MDa for pDNA and 13
kDa for siRNA. Therefore, larger pore sizes are
required for an effective nucleic acid loading. It has
been reported that 250 nm‐sized MSNs with large
pores of 23 nm (LMSN) exhibit an appreciably higher
loading capacity for pDNA (25 µg pDNA/mg LMSN)
than those featuring 2 nm‐sized small pores (SMSN)
(<8.3 µg pDNA/mg SMSN) [47]. Additionally,
LMSNs modified with PEG were shown to be
capable of eliciting a higher loading capacity for
siRNA (16.6 µg siRNA/mg LMSN), whereas
negligible siRNA loading was observed for the
SMSN counterparts presumably because of their
small pore size as well as the PEGylation impeding
siRNA binding to the outer surface [48]. Assuming
that the siRNA surface is considered as a rectangle
with a length of 6 nm and a width of 3 nm [49, 50],
the ratio of the surface area occupied with siRNA to
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
7 Nano Res.
the total surface area was calculated to be only ~4%,
even for LMSN. This substantially small value
suggests that siRNA was not enclosed in the deep
part of the uniform pores most likely because of the
were successfully prepared at relatively high silicate
concentrations (>5 mM) [56]. The formation of a silica
layer was verified by the change of the charge from
positive to negative, i.e., +20 to −20 mV in zeta
potential, and the increase in size of the order of 5–10
nm in thickness after the silica coating, as well as by
elemental analysis [58, 59]. The silica coating was
shown to contribute appreciably to the stabilization
Figure 4. Silica coating of PICs containing nucleic acids. Schematic illustration of the preparation of sPIC (a) and a TEM image of sPIC (b). Reproduced with permission from ref. [57]. Copyright American Chemical Society, 2012 (b).
| www.editorialmanager.com/nare/default.asp
8 Nano Res.
of PICs against the dissociation induced by counter
polyanions. Interestingly, the formed silica layer (or
silica hydrogel) was gradually dissolved under dilute
conditions within a day, according to the equilibrium
shift to silicic acids [56, 57]. This dissolution allowed
for the efficient release of the nucleic acid payloads to
take place after sPICs were internalized into cultured
cells.
3.2 Releasing of nucleic acids
Nucleic acid delivery systems must effectively
release the nucleic acid payload after reaching the
target site while avoiding premature release. These
apparently conflicting functions have been integrated
into delivery systems by utilizing
environment‐responsive chemical reactions. Since the
cytoplasm is a reducing environment compared with
the extracellular milieu because of the presence of a
high concentration of glutathione (GSH), the most
abundant reducing agent in the body, disulfide
bonds are utilized as intracellularly cleavable bonds.
In a series of studies, the SiNP surface was modified
with amine compounds via disulfide bonds [43–45,
51, 60]. For example, an amine‐functionalized MSN
(MSN‐NH2) was prepared by co‐condensation of
tetraethoxy orthosilicate (TEOS) and 3‐aminopropyl
triethoxysilane (APTMS) [43]. The MSN‐NH2 was
then reacted with succinic anhydride and finally
modified with cystamine. The as‐prepared MSN
(MSN‐Linker‐Cys) with disulfide bonds was mixed
with oligo DNA for its electrostatic adsorption, and
the DNA‐loaded MSN‐Linker‐Cys was conjugated
with N‐hydroxysuccinimide (NHS)‐terminated PEG
(MW: 2,000). Although the obtained nanoparticle
released only 10% of the adsorbed DNA at pH 7.4 in
the absence of GSH, it released ~100% after three‐day
incubation at 10 mM GSH at the same pH.
Endosomal acidic pH (~5.5) can also be utilized as
an intracellular signal for triggering the release of
nucleic acid payloads from delivery systems. Thus,
an acidic‐pH‐responsive film‐coated magnetic MSN
was developed for siRNA delivery [64, 65]. The
acidic‐pH‐responsive film was prepared on the
magnetic MSNs incorporating siRNA within the pore
by their mixture with tannic acids and aluminum
ions, which formed a chelate complex on the
magnetic MSN surface that was destabilized at acidic
pH presumably because of the protonation of some
of the phenolic hydroxyl groups in the tannic acids
[63]. This chelate coating suppressed the release of
siRNA payloads and enhanced their tolerance
against nuclease degradation at physiological pH. On
the other hand, approximately 90% of the siRNA
payloads were released from the system when
incubated for 48 h at pH 5. Ultimately, the acidic
pH‐responsive film‐coated magnetic MSN showed 60%
reduction in viability of cultured osteosarcoma cells
under magnetic field by delivering polo‐like kinase
1‐targeted siRNA (75 nM), whereas a control MSN
loaded with enhanced green fluorescence protein
(EGFP)‐targeted siRNA did not show any
cytotoxicity under the same condition. This result is
consistent with the efficient siRNA release from the
MSN in the cells. It should be noted that
acidic‐pH‐responsive drug release strategies have
been extensively investigated for low MW drugs [20].
For example, the DOX release from MSNs was
significantly accelerated by surface coating with
cationic polymers, such as PEI or chitosan, bearing
low pKa amines; the high protonation of these
polymers under acidic conditions causes an increase
of the electrostatic repulsion with each other, thereby
promoting the decoating and the concomitant drug
release [52, 64]. The dissolution of ion‐doped SiNPs
facilitated at acidic pH can also be utilized to trigger
the release of payloads [65, 66].
External stimuli, such as light and heat, can also be
used to promote the release of nucleic acids from
SiNP‐based delivery systems. Surface modification of
SiNP with cationic moieties via UV light‐cleavable
linker enables the release of nucleic acids responding
to UV light, similar to the disulfide bond responding
to reductive environments. Also, dehybridization of
oligonucleotide duplex by heating permits the
release of single‐stranded oligonucleotide when one
end of the oligonucleotide duplex was conjugated to
SiNP. In this regard, direct UV irradiation and
heating to the body may have a difficulty in in vivo
therapeutic applications because they more likely
exert adverse side effects, including tissue damages.
Additionally, UV light is appreciably absorbed by the
skin and cannot reach deep tissues [67]. To overcome
this drawback in UV irradiation and heating,
upconversion nanoparticles (UCNPs) and Au
nanorods (AuNRs) have been utilized to locally
generate UV light and heat, respectively, by
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
9 Nano Res.
irradiation of near‐infrared (NIR) light. Of note, NIR
light is more applicable to human body than the
other light stimuli because it can penetrate tissues
more deeply with the minimum adverse effect [67].
Indeed, UCNPs or AuNRs have been encapsulated
into SiNPs for the NIR‐triggered nucleic acid release.
In a previous study, an UCNP‐encapsulated SiNP
was modified with a cationic moiety via an
o‐nitrobenzyl photolabile linker, followed by loading
of siRNA on the surface through electrostatic
interactions [68]. The irradiation of NIR light at 980
nm for 2 hours resulted in a decrease in the
absorbance at 321 nm derived from the photolabile
linker, indicating that the photolabile linker was
cleaved to lose the UV absorbance. When
EGFP‐targeted siRNA was loaded into the
UCNP‐encapsulated SiNP and transfected to
cultured EGFP‐expressing HeLa cells, the
fluorescence intensity derived from EGFP was
significantly decreased to approximately 20% after
NIR irradiation, compared to the case without the
irradiation. This result demonstrates that the
UCNP‐encapsulated SiNP facilitated the release of
siRNA in the cells in response to NIR irradiation. In
another study, AuNRs were incorporated into SiNP
for the NIR light‐induced release of single‐stranded
oligonucleotides [69]. The oligonucleotide duplexes
with a thiol group at the one‐strand end were
conjugated on AuNR‐incorporated mesoporus SiNP
(AuMS) via Michael addition reaction with
maleimide groups on the silica surface. The release of
fluorescently labeled single‐stranded
oligonucleotides from AuMS was clearly observed
after NIR irradiation. Similarly, dendronized
semiconducting polymers are also reported as a
photothermal conversion material for the
photo‐responsive gene delivery [70].
3.3 Cellular uptake and intracellular trafficking
Nucleic acids must be efficiently uptaken by target
cells and should escape from the endosome in order
to avoid the lysosomal digestion. For enhanced
cellular uptake and endosomal escape, SiNPs have
been surface‐coated with cationic molecules. The
cellular membrane is negatively charged, and
therefore, it is attached by positively charged
nanoparticles. This attachment promotes the
adsorptive endocytosis of nanoparticles. Positively
charged nanoparticles can also bind to the endosomal
membrane to destabilize the membrane integrity and
induce the endosomal escape. In this regard, cationic
molecules or polymers with low pKa amines are
known to favor the endosomal escape [38]. Low pKa
amines, which are not protonated in the extracellular
milieu (pH 7.4), can be protonated in the acidic
endosomal compartment. This amine protonation
induces the influx of protons and chloride ions into
the endosome, which results in an elevation of the
osmotic pressure and the consequent destabilization
of the endosomal membrane. This mechanism is
called the proton sponge effect [38]. It should be
noted that another mechanism is also proposed for
the membrane destabilization, which relies on the
positive charge density of cationic materials. In the
latter mechanism, the amine protonation in the
endosome provides the cationic component with
higher positive charge density, leading to the
stronger binding to the negatively charged
endosomal membrane and, thereby, inducing the
membrane destabilization more effectively [71].
PEI is one of the most widely tested cationic
polymers for enhanced endosomal escape and has
been used for the surface coating of SiNPs [35, 36, 44,
72]. The MW of PEI has been shown to affect
significantly both transfection efficiency and
cytotoxicity of PEI‐coated SiNPs [35]. Thus, PEI with
large MWs of about 25 kDa afforded high
transfection efficiency presumably because of the
enhancement of the endosomal escape, whereas
severe cytotoxicity was also elicited. In contrast,
lower cytotoxicity in cultured pancreatic (PANC‐1
and BxPC3) and liver (HEPA‐1) cancer cells was
observed for PEI with modestly smaller MWs of ~10
kDa, whose transfection efficiency was still
significant. To further decrease the cytotoxicity
derived from cationic materials, a variety of cationic
polymers have been developed for the
functionalization of SiNPs [57, 59, 72–74]. For
example, our group developed polyaspartamide
derivatives bearing two‐ or four‐repeated
aminoethylene units in the side chains (termed
PAsp(DET) or PAsp(TEP), respectively), which
exhibited a large difference in the degree of
protonation between pH 7.4 and pH 5.5. This result is
consistent with a high proton sponge capacity and
positive charge density in the acidic endosomal
| www.editorialmanager.com/nare/default.asp
10 Nano Res.
compartment that favors endosome disruption [75,
76]. These derivatives were utilized to provide the
aforementioned sPICs with endosome escapability by
surface‐coating [57, 59]. The obtained multilayered
PICs (mPICs) induced efficient gene silencing in vitro
and in vivo by delivering siRNA. Similarly, fusogenic
peptides, such as KALA and H5WYG, were utilized
for the fabrication of endosome‐escapable SiNPs,
which afforded a high RNAi efficiency in vitro and in
vivo [72–74]. It is worthy of note that the KALA
peptide consists of 30 amino acids containing a
repeated sequence of lysine, alanine, leucine, and
alanine, and the H5WYG peptide has a
histidyl‐residue‐rich sequence [77, 78]. These
peptides undergo conformational changes at acidic
pH, inducing the membrane destabilization in a
similar manner to that of the influenza‐virus‐derived
fusogenic peptide.
Cationic nanoparticles enable the adsorptive
endocytosis through electrostatic interactions with
the negatively charged cellular surface. However,
these nanoparticles bind nonspecifically to anionic
proteins and glycosaminoglycans, as well as to blood
cells and endothelial cells, under in vivo conditions.
These nonspecific adsorptions may induce secondary
aggregation in blood capillaries and local tissues,
ethylene glycol] to the primary amines in the lipid
bilayer of protocells, SP94 and aforementioned
H5WYG peptides with C‐terminal cysteine residues
were conjugated to the PEG linker (~6 SP94/protocell
and ~240 H5WYG/protocell). The peptide‐installed
protocells induced a 90% decrease in target protein
expression by delivering siRNA to the hepatocellular
carcinoma cells, whereas no gene silencing was
elicited in normal hepatocytes, demonstrating the
high specificity of the SP94 peptide for hepatocellular
carcinoma targeting.
Figure 5. Schematic illustrations of SiNP functionalization. PEGylation with targeting ligands via amidation (a), electrostatic binding (b), or fusion with PEGylated liposome (c). Ligand density can be controlled by changing the silicate concentration in the preparation of sPICs (b). Endosome escapability can be imparted to the protocells by conjugating endosomolytic peptides on the surface (c). Adapted with permission from ref. [44]. Copyright American Chemical Society, 2016 (a); and ref. [74], Copyright American Chemical Society, 2012 (c).
3.4 Toxicity of silica nanoparticles
The cytotoxicity of SiNPs is affected by various
parameters, such as size, surface charge, surface
chemistry, shape, stability, and concentration of the
nanoparticles, which renders the precise
understanding of the cytotoxic mechanism of SiNPs
difficult. Nevertheless, recent studies have suggested
that the silanol groups present on the SiNP surface
are one of the major causes of the adverse side effects
of SiNPs. In particular, these silanol groups are
reported to be responsible for the hemolysis of
erythrocytes [90, 91] and also for the generation of
the radical oxygen species (ROS), leading to the
oxidative stress in endosomes and the production of
proinflammatory cytokines [92, 93]. One possible
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
13 Nano Res.
cause for the ROS production is the activation of
dihydronicotinamide‐adenine dinucleotide
phosphate oxidase by the assembly of cellular
membrane lipids with the silanol groups on SiNPs
[93]. For this reason, chemical modification of the
surface silanol groups has been proved as an
effective way to reduce the inherent toxicity of SiNPs.
For example, the conversion of the silanol groups to
carboxylic or sulfonic groups by silane coupling
reaction led to a dramatic reduction of the hemolytic
activity of the carboxylated or sulfonated SiNPs to 2%
of that of the nonmodified parent SiNPs [90]. It is
worthy of note that all of the tested SiNPs displayed
similar negative zeta potentials (–50 mV), which
indicates that the hemolytic activity of SiNPs cannot
be simply correlated with the surface charge density.
In a similar manner, carboxylated or sulfonated
SiNPs of 1 µm in diameter generated a lower level of
ROS compared with the nonmodified control when
they were incubated with a human acute monocytic
leukemia cell [93]. Another promising approach for
reducing the toxicity of SiNPs is their PEGylation.
PEGylated MSNs have been shown to reduce
significantly the hemolytic activity (less than 1% of
deionized water) even at high concentration (500
µg/mL MSN), compared with non‐PEGylated control
MSNs (15% of deionized water), through the
inhibition of the contact of erythrocytes with the
silanol groups on the silica surface [94].
The accumulation toxicity of SiNPs is a critical
concern when considering their use under in vivo
conditions [19]. To avoid these adverse effects, the
metabolic pathway of SiNPs should be adequately
considered in material design. A clinical trial using
SiNPs with a diameter of 7 nm demonstrated that
such ultra‐small nanoparticles could be significantly
eliminated from the body through renal excretion
over three days after systemic administration [2].
However, this ultra‐small size limits the loading
capacity of the nanoparticles for macromolecular
drugs such as nucleic acids. In this regard, the
development of biodegradable SiNPs is envisaged as
a plausible approach for accelerating the elimination.
The silanol groups have been shown to play a crucial
role in triggering the degradation of SiNPs, mainly
through the formation of hydrogen bonding with
water molecules, which reduces the energy barrier of
the proton transfer from a water molecule to a
siloxane bond [95]. On the other hand, the
PEGylation of SiNPs through silane coupling, which
is associated with the conversion of silanol to
siloxane, was reported to lower significantly the
degradation rate of SiNPs [96–99]. Thus, the content
of silanol groups seems to generate a conflicting
situation regarding the toxicity of SiNPs; although an
increase in the silanol content promotes a faster
degradation of SiNPs, it may also facilitate the
hemolytic activity and ROS generation. A plausible
solution for this conflict is the functionalization of the
silica matrices, rather than the surface, to achieve
biodegradability.
A number of approaches for the preparation of
SiNPs with enhanced biodegradability have been
reported [19]. One of these approaches is the doping
of SiNPs with inorganic ions. Thus, manganese ions
were doped into MSNs by hydrothermal treatment of
MSNs in the presence of MnSO4·H2O at 180 °C for 12
h [65]. The Mn‐doped MSNs had Mn–O bonds
sensitive to acidic and reductive conditions within
the silica matrix and were completely degraded 48 h
after incubation with 10 mM GSH at pH 5. It is
worthy of note that the environment‐responsive
cleavage of the Mn–O bonds and subsequent
manganese extraction contribute to the acceleration
of the hydrolysis of the siloxane bonds. In addition,
when the Mn‐doped MSNs were incubated for three
days with cultured cancer cells and then subjected to
transmission electron microscopic (TEM) imaging, no
nanoparticle structures were observed in the cells
(Fig. 6). Furthermore, upon intravenous
administration of PEGylated Mn‐doped MSNs into
mice, 70% and 10% of the Si content were eliminated
from the body as urine and feces, respectively, 48 h
after administration, which was in sharp contrast
with the elimination rates of 15% and 25% observed
for the parent MSN without Mn doping. In another
study, a hybrid nanocomposite MSN/HAP with silica
and hydroxyapatite was prepared in the presence of
CaCl2 and Na2HPO4·12H2O [66]. The Fourier
transform infrared and X‐ray photoelectron
spectroscopy data indicated the presence of the Si–
O–Ca–O–Si structure in the silica matrix. Although
the MSN/HAP incubated for one week at pH 7.4
hardly released calcium ions (<5 mg/L), virtually all
of the calcium ions were released after incubation for
8 h at pH 5 (~230 mg/L). The corresponding TEM
| www.editorialmanager.com/nare/default.asp
14 Nano Res.
image displayed that the MSN/HAPs with a diameter
of 90 nm were degraded to pieces smaller than 20 nm
after 12 h of incubation at pH 5. This significant
degradability was mainly due to the acid‐triggered
degradation of hydroxyapatite in the silica matrix
and the concomitant removal of Ca2+ ions from the
structure of Si–O–Ca. When fluorescently labeled
MSN/HAPs and MSNs were intravenously
administrated into mice, the fluorescence intensity in
the urine 48 h after injection was five‐fold higher for
MSN/HAP compared with MSN. These results
demonstrate the higher degradability of the hybrid
nanocomposite leading to a more rapid renal
clearance.
Another approach for providing SiNPs with
biodegradability is the installation of cleavable bonds
that respond to a specific biological environment.
Disulfide bonds have been installed into silica
matrices to fulfill the reductive
environment‐responsive degradation of SiNPs [100–
102]. For example, disulfide‐installed SiNPs were
fabricated by the Stöber method using TEOS with
disulfide‐bridged silane compounds, e.g.,
bis(triethoxysilylpropyl) disulfide [100, 101]. TEM
images displayed that a disulfide‐installed MSN of 90
nm in diameter was degraded to smaller pieces (~10
nm) after incubation for seven days under a
reductive condition mimicking the cell interior. This
disulfide‐installed MSN was also observed to
degrade in cultured C6 glioma cells after 48 h of
incubation [101]. In a different study,
disulfide‐installed SiNPs were fabricated by the
Stöber method using APTMS with
dithiobis‐succinimidyl propionate (DTSP) [102].
Notably, these disulfide‐installed SiNPs started to
decompose in 10 mM dithiothreitol (DTT) solution
within 2.5 h, whereas their degradation was not
observed in a solution without DTT. It should also be
noted that the degradation profile of these
disulfide‐installed SiNPs was further accelerated by
noncovalent drug loading through the formation of
weakly condensed silica network [103]. A
DOX‐loaded disulfide‐installed SiNP (DS‐DOX) was
fabricated by the Stöber method with TEOS, DOX,
and the disulfide‐bridged silane compound BTOCD,
which was, in turn, prepared by reacting
3‐(triethoxysilyl)propyl isocyanate with cystamine
dihydrochloride [104]. The corresponding TEM
images displayed the enhanced degradability of
DS‐DOX after incubation for four days with 10 mM
DTT, compared with a control disulfide‐installed
SiNP without DOX (DS). When intravenously
administrated into mice, 50% of the Si content of
DS‐DOX was eliminated from the body 48 h after
administration. This value was much higher than
those of DS (~20%) and a control SiNP without
DS/DOX (~10%).
Figure 6. TEM images displaying the morphological change of Mn-doped MSNs at pH 5 after 6 h (a), 12 h (b), and 48 h (c) of incubation with 10 mM GSH or 1 d (d), 2 d (e), and 3 d (f) after internalization by cultured cells. Adapted with permission from ref. [65]. Copyright American Chemical Society, 2016.
4 Codelivery with low molecular weight
drugs
The codelivery of different types of drugs is a
promising therapeutic modality that benefits from
synergistic therapeutic effects. SiNPs have shown a
strong potential for the codelivery of small drugs
with nucleic acids. Small drugs can be embedded in
the silica matrix or entrapped in the porous structure
simultaneously with nucleic acids [45, 103, 105–107].
For example, a codelivery system of DOX with
pDNA was developed [103], in which DOX was
embedded in BTOCD‐modified SiNPs with redox
responsiveness (as described in Section 3.4), and
pDNA was electrostatically incorporated onto the
nanoparticle surface (Fig. 7). Interestingly, the
embedding of DOX facilitated the decomposition of
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
15 Nano Res.
the BTOCD‐modified SiNPs because of the weakly
condensed silica network (as described in Section 3.4),
which resulted in the efficient release of DOX and
pDNA after being internalized into cells. This system
enabled both high gene expression efficiency and
cytotoxicity derived from pDNA and DOX,
respectively, in cultured cancer cells. Furthermore,
the codelivery of DOX and pDNA encoding p53
achieved a higher anticancer effect in C6 glioma
tumor‐bearing mice compared with the single
delivery of either DOX or p53‐coding pDNA with the
same system.
Figure 7. Schematic illustration of redox responsive codelivery system of DOX and pDNA. DOX is embedded in disulfide-containing silica network. Adapted with permission from ref. [103]. Copyright Wiley-VCH, 2017.
One of the critical issues in cancer chemotherapy is
the drug resistance of cancer cells. Drug resistance
occurs through the expression of proteins linked to
the resistance, such as drug efflux pump and
antiapoptotic proteins [108]. A plausible way to
overcome the cancer cell resistance is the codelivery
of anticancer drugs with siRNA that induces the gene
silencing of the resistant proteins. The first example
that utilized this combination incorporated both
DOX and siRNA targeting B‐cell lymphoma 2 (Bcl‐2)
into MSNs [105]. It is worth noting that the Bcl‐2
protein plays a critical role in the antiapoptotic
mechanism, where overexpressed Bcl‐2 proteins
effectively block a common pathway of cell death
induced by cytotoxic drugs [108]. Thus, the silencing
of the Bcl‐2 gene enhances the cytotoxic effect of
DOX by suppressing the antiapoptotic mechanism.
DOX was incorporated into the 3 nm‐sized pores,
and the drug‐loaded MSNs (200 nm in diameter)
were shielded with a polyamidoamine dendrimer for
minimizing premature drug release, which was
followed by siRNA loading on the MSN surface. The
DOX/siRNA‐coloaded delivery systems showed a
superior cytotoxic effect (IC50: 17 nM) in cultured
multidrug‐resistant human ovarian cancer
(A2780/DOX) cells compared with the single delivery
of DOX (IC50: 2.2 µM). A different study also
reported the synergistic effect of DOX and siRNA in
a multidrug‐resistant breast cancer (MCF‐7/MDR)
xenograft model [106]. This study fabricated a
PEG‐PEI block copolymer‐modified MSN with a
diameter of 50 nm, which incorporated DOX and
siRNA encoding the drug efflux pump
P‐glycoprotein (P‐gp). Note that the P‐gp‐encoded
siRNA was selected using a high‐throughput
screening in a multidrug‐resistant cancer cell line to
obtain an enhanced synergetic effect. As a result,
MSN elicited a significantly enhanced antitumor
effect by 80% of tumor inhibition (It) in an
MCF‐7/MDR xenograft model, as compared with
DOX‐loaded MSN (It = 62%) and DOX and scrambled
siRNA‐coloaded MSN (59%).
Recent study has aimed at increasing the
drug‐loading capacity of SiNPs for enhanced
anticancer effect. A significant increase in
drug‐loading capacity was achieved by modifying
MSNs with cycrodextrin (CD)‐grafted PEI (PEI‐CD)
[107]. The hydrophobic cavity of CD was utilized for
additional loading of DOX, which afforded a 2.5‐fold
improvement of the DOX‐loading capacity compared
with the PEI‐modified MSN without CD. DOX and
siRNA were efficiently released from the
PEI‐CD‐modified MSN under acidic condition
probably because of the electrostatic repulsion
between DOX and protonated PEI [109]. This system
greatly suppressed the tumor growth after
intravenous administration in an orthotopic
MDA‐MB‐231 breast cancer model.
Besides the codelivery with anticancer drugs, the
codelivery of nucleic acids with probes has
promising applications in theranostics [110]. For
example, a trans‐cyclooctene (TCO) and
dibenzocyclooctyne (DBCO)‐bifunctionalized PEG
was first conjugated with an azide‐modified
compound composed of aspartic acid–glutamic acid–
valine–aspartic acid (DEVD) peptide and
| www.editorialmanager.com/nare/default.asp
16 Nano Res.
2‐acetyl‐6‐amino‐napthalene (AAN), which is a
caspase 3‐responsive fluorescent probe, to obtain a
TCO–DEVD–AAN conjugate. The amide bond
between DEVD and AAN is cleaved by caspase‐3,
which is, in turn, activated during the apoptosis
process, and the cleaved AAN emits fluorescence as a
molecular sensor for apoptosis. After a
tetrazine‐modified MSN was loaded with a small
molecular inhibitor (sm‐21) in the pore and with
ASOs on the surface, TCO–DEVD–AAN was
conjugated to the MSN by inverse electron demand
Diels–Alder reaction. It is worthy of note that both
payloads, sm‐21 and ASO, were selected for inducing
the apoptosis of cancer cells by the inhibition of the
miRNA‐21 activity. Thus, the sm‐21/ASO‐coloaded
MSN significantly inhibited the miRNA‐21 activity in
cultured HeLa cells. The fluorescence was first
detected from the cultured HeLa cells after 6 h of
incubation with sm‐21/ASO‐coloaded MSNs, and the
fluorescence intensity continued to increase for 18 h.
This study demonstrates that cancer cell apoptosis
can be continuously monitored using this codelivery
system.
5 Conclusions and future perspectives
In this review, the potential of SiNPs as nucleic acid
delivery systems has been highlighted. MSNs with
large pores (~20 nm) have a higher
nucleic‐acid‐loading capacity than conventional
MSNs with small pores (~2 nm). This loading
capacity is especially high for dendritic MSN with
large center‐radial pores. The high loading capacity
allows us to achieve a high drug weight/nanoparticle
weight ratio and decreases the dose amount of
nanoparticle components, thereby reducing the
potential adverse side effects. The silica coating of
PICs is also a promising way to incorporate large
nucleic acids, such as pDNA and mRNA, into
silica‐based delivery systems. The sPIC system
enables the complete encapsulation of large nucleic
acids within the silica layer, protecting them
effectively from enzymatic degradation. Additionally,
the SiNP surface can be readily modified with
functional materials in order to achieve the required
biocompatibility and functionalities for preferential
binding to the target cells (active targeting),
endosomal escape, and selective release of nucleic
acid payloads in the cells. Furthermore, the
codelivery of nucleic acids with anticancer drugs by
SiNPs is highly effective for the treatment of
drug‐resistant cancers, which are one of the most
critical concerns in cancer chemotherapies. In
particular, the drug‐resistant gene silencing by
siRNA has been shown to enhance significantly the
anticancer effect of anticancer drugs.
Despite the aforementioned advantages of
SiNP‐based delivery, there are several critical issues
that must be addressed for its pharmaceutical
implementation. In this regard, acute toxicities of
SiNPs, such as hemolysis, have been correlated with
their outer surface area or the amount of silanol
groups on the outer surface. The PEGylation of the
SiNP surface has been demonstrated to significantly
reduce the hemolytic activity by disturbing the
contact between the surface silanol groups and
erythrocytes through steric hindrance. Meanwhile,
the surface silanol groups have a major role in the
hydrolysis of SiNPs. The conversion of silanol groups
model mice. Ultimately, MSN elicited a significant
therapeutic effect, which was comparable to that
obtained from four times higher doses of free 5‐ASA.
To perform the delivery of nucleic acids through the
oral route, the delivery systems must protect the
nucleic acid payloads under the fairly harsh
conditions of the gastrointestinal tract, which
contains various digestive enzymes. Moreover, the
delivery systems need to be tolerant to the strong
acid environment in the stomach. The high stability
of SiNPs, especially at acidic pH, should render them
applicable for the oral delivery of nucleic acids.
Acknowledgments
This work was financially supported by Center of
Innovation (COI) program from Japan Science and
Technology Agency (JST), Grants‐in‐Aid for Scientific
Research (KAKENHI Grant Numbers 17H02098)
from Ministry of Education Culture, Sports, Science
and Technology (MEXT) and Japan Society for the
Promotion of Science through Program for Leading
Graduate Schools (MERIT). This work was also
partially supported by the Project for Cancer
Research and Therapeutics Evolution (P‐CREATE)
and Basic Science and Platform Technology Program
for Innovative Biological Medicine from Japan
Agency for Medical Research and Development
(AMED).
References
[1] Benezra, M.; Penate-Medina, O.; Zanzonico, P. B.; Schaer,
D.; Ow, H.; Burns, A.; DeStanchina, E.; Longo, V.; Herz, E.; Iyer, S.; et al. Multimodal silica nanoparticles are effective cancer-targeted probes in a model of human melanoma. J. Clin. Invest. 2011, 121, 2768–2780.
[2] Phillips, E.; Penate-Medina, O.; Zanzonico, P. B.; Carvajal, R. D.; Mohan, P.; Ye, Y.; Humm, J.; Gönen, M.; Kalaigian, H.; Schöder, H.; et al. Clinical translation of an ultrasmall inorganic optical-PET imaging nanoparticle probe. Sci.
Transl. Med. 2014, 6, 260ra149.
[3] Meng, H.; Xue, M.; Xia, T.; Ji, Z.; Tarn, D. Y.; Zink, J. I.; Nel, A. E. Use of size and a copolymer design feature to improve the biodistribution and the enhanced permeability and retention effect of doxorubicin-loaded mesoporous silica nanoparticles in a murine xenograft tumor model. ACS Nano 2011, 5, 4131–4144.
[4] Matsumura, Y.; Maeda, H. A new concept for macromolecular therapeutics in cancer chemotherapy: mechanism of tumoritropic accumulation of proteins and the antitumor agent smancs. Cancer Res. 1986, 46, 6387–6392.
[5] Maeda, H. Toward a full understanding of the EPR effect in primary and metastatic tumors as well as issues related to its heterogeneity. Adv. Drug. Deliv. Rev. 2015, 91, 3-6.
[6] Stöber, W.; Fink, A.: Bohn, E. Controlled growth of monodisperse silica spheres in the micron size range. J. Colloid Interface Sci. 1968, 26, 62–69.
[7] Osseo-Asare, K.; Arriagada, F. J. Preparation of SiO2 nanoparticles in a non-ionic reverse micellar system. Colloids Surf. 1990, 50, 321–339.
[8] Yokoi, T.; Sakamoto, Y.; Terasaki, O.; Kubota, Y.; Okubo, T.; Tatsumi, T. Periodic arrangement of silica nanospheres assisted by amino acids J. Am. Chem. Soc. 2006, 128, 13664–13665.
[9] Hartlen, K. D.; Athanasopoulos, A. P. T.; Kitaev, V. Facile preparation of highly monodisperse small silica spheres (15 to >200 nm) suitable for colloidal templating and formation of ordered arrays. Langmuir 2008, 24, 1714–1720.
[10] Wang, H.; Schaefer, K.; Moeller, M. In situ immobilization of gold nanoparticle dimers in silica nanoshell by microemulsion coalescence. J. Phys. Chem. C 2008, 112, 3175–3178.
[11] Stjerndahl, M.; Andersson, M.; Hall, H. E.; Pajerowski, D. M.; Meisel, M. W.; Duran, R. S. Superparamagnetic Fe3O4/SiO2 nanocomposites: enabling the tuning of both the iron oxide load and the size of the nanoparticles. Langmuir 2008, 24, 3532–3536.
[12] Kuijk, A.; van Blaaderen, A.; Imhof, A. Synthesis of monodisperse, rodlike silica colloids with tunable aspect ratio. J. Am. Chem. Soc. 2011, 133, 2346–2349.
[13] Chen, M.; Wu, L.; Zhou, S.; You, B. A method for the fabrication of monodisperse hollow silica spheres. Adv. Mater. 2006, 18, 801–806.
[14] Lu, F.; Wu, S. -H.; Hung, Y.; Mou, C. -Y. Size effect on cell uptake in well-suspended, uniform mesoporous silica nanoparticles. Small 2009, 5, 1408–1413.
[15] Du, X.; Shi, B.; Liang, J.; Bi, J.; Dai, S.; Qiao, S. Z. Developing functionalized dendrimer-like silica nanoparticles with hierarchical pores as advanced delivery nanocarriers. Adv. Mater. 2013, 25, 5981–5985.
[17] Du, X.; Qiao, S. Z. Dendritic silica particles with center-radial pore channels: promising platforms for catalysis and biomedical applications. Small 2015, 11, 392–413.
Nanoparticles for bioimaging. Adv. Colloid Interface Sci. 2006, 123–126, 471–485.
[19] Croissant, J. G.; Fatieiev, Y.; Khashab N. M. Degradability and clearance of silicon, organosilica, silsesquioxane, silica mixed oxide, and mesoporous silica nanoparticles. Adv. Mater. 2017, 29, 1604634.
[20] Moreira, A. F.; Dias, D. R.; Correia, I. J. Stimuli-responsive mesoporous silica nanoparticles for cancer therapy: A review. Micropor. Mesopor. Mat. 2016, 236, 141–157.
[21] Tang, L.; Cheng, J. Nonporous silica nanoparticles for nanomedicine application. Nano Today 2013, 8, 290–312.
[22] Yin, H.; Kanasty, R. L.; Eltoukhy, A. A.; Vegas, A. J.; Dorkin, J. R.; Anderson, D. G. Non-viral vectors for gene-based therapy. Nat. Rev. Genet. 2014, 15, 541–555.
[23] Kibbe, M. R.; Hirsch, A. T.; Mendelsohn, F. O.; Davies, M. G.; Pham, H.; Saucedo, J.; Marston, W.; Pyun, W. -B; Min, S. -K; Peterson, B. G. et al. Safety and efficacy of plasmid DNA expressing two isoforms of hepatocyte growth factor in patients with critical limb ischemia. Gene Ther. 2016, 23, 306–312.
[24] Wilson, R. C.; Doudna. J. A. Molecular mechanisms of RNA interference. Annu. Rev. Biophys. 2013, 42, 217–239.
[25] Fire, A.; Xu, S.; Montgomery, M. K.; Kostas, S. A.; Driver, S. E.; Mello, C. C. Potent and specific genetic interference by double-stranded RNA in Caenorhabditis elegans. Nature 1998, 391, 806–811.
[26] Suhr, O. B.; Coelho, T.; Buades, J.; Pouget, J.; Conceicao, I.; Berk, J.; Schmidt, H.; Waddington-Cruz, M.; Campistol, J. M.; Bettencourt, B. R.; et al. Efficacy and safety of patisiran for familial amyloidotic polyneuropathy: a phase II multi-dose study. Orphanet J. Rare Dis. 2015, 10, 109.
[27] Stephenson, M. L.; Zamecnik, P. C. Inhibition of Rous sarcoma viral RNA translation by a specific oligodeoxyribonucleotide. Proc. Natl. Acad. Sci. USA 1978, 75, 285–288.
[28] Sharma, V. K.; Sharma, R. K.; Singh, S. K. Antisense oligonucleotides: modifications and clinical trials. Med. Chem. Comm. 2014, 5, 1454–1471.
[29] Seymour, L. W.; Duncan, R.; Strohalm, J.;, Kopeček, J. Effect of molecular weight (Mw) of N-(2-hydroxypropyl)methacrylamide copolymers on body distribution and rate of excretion after subcutaneous, intraperitoneal, and intravenous administration to rats. J. Biomed. Mater. Res. 1987, 21, 1341–1358.
[30] Choi, H. S.; Liu, W.; Misra, P.; Tanaka, E.; Zimmer, J. P.; Ipe, B. I.; Bawendi, M. G.; Frangioni, J. V. Renal clearance of quantum dots. Nat. Biotechnol. 2007, 25 1165–1170.
[31] Schroeder, A.; Heller, D. A.; Winslow, M. M.; Dahlman, J. E.; Pratt, G. W.; Langer, R.; Jacks, T.; Anderson, D. G. Treating metastatic cancer with nanotechnology. Nat. Rev. Cancer 2012, 12, 39–50.
[32] Albanese, A.; Tang, P. S.; Chan, W. C. W. The effect of nanoparticle size, shape, and surface chemistry on biological systems. Annu. Rev. Biomed. Eng. 2012, 14, 1–16.
[33] Harris, J. M.; Chess, R. B. Effect of pegylation on pharmaceuticals. Nat. Rev. Drug. Discov. 2003, 2, 214–221.
[34] Matteoni, R.; Kreis, T. E.; Translocation and clustering of endosomes and lysosomes depends on microtubules. J. Cell Biol. 1987, 105, 1253–1265.
[35] Xia, T.; Kovochich, M.; Liong, M.; Meng, H.; Kabehie, S.; George, S.; Zink, J. I.; Nel, A. E. Polyethyleneimine coating enhances the cellular uptake of mesoporous silica nanoparticles and allows safe delivery of siRNA and DNA constructs. ACS Nano 2009, 3, 3273–3286.
[36] Fuller, J. E.; Zugates, G. T.; Ferreira, L. S.; Ow, H. S.; Nguyen, N. N.; Wiesner, U. B.; Langer, R. S. Intracellular delivery of core–shell fluorescent silica nanoparticles. Biomaterials 2008, 29, 1526–1532.
[37] Zhu, S. -G.; Xiang, J. -J.; Li, X. -L.; Shen, S. -R.; Lu, H.; Zhou, J.; Xiong, W.; Zhang, B. -C.; Nie, X. -M.; Zhou, M.; et al. Poly(L-lysine)-modified silica nanoparticles for the delivery of antisense oligonucleotides. Biotechnol. Appl. Biochem. 2004, 39, 179–187.
[38] Behr, J. -P. The proton sponge: A trick to enter cells the viruses did not exploit. Chimia. 1997, 51, 34–36.
[39] Roy, I.; Ohulchanskyy, T. Y.; Bharali, D. J.; Pudavar, H. E.; Mistretta, R. A.; Kaur, N.; Prasad, P. N. Optical tracking of organically modified silica nanoparticles as DNA carriers: A nonviral, nanomedicine approach for gene delivery. Proc. Natl. Acad. Sci. U. S. A. 2005, 102, 279–284.
[40] Bharali, D. J.; Klejbor, I.; Stachowiak, E. K.; Dutta, P.; Roy, I.; Kaur, N.; Bergey, E. J.; Prasad, P. N.; Stachowiak, M. K. Organically modified silica nanoparticles: A nonviral vector for in vivo gene delivery and expression in the brain. Proc. Natl. Acad. Sci. U. S. A. 2005, 102, 11539–11544.
[41] Huang, X.; Tao, Z.; Praskavich, J. C.; Goswami, A.; Al-Sharab, J. F.; Minko, T.; Polshettiwar, V.; Asefa, T. Dendritic silica nanomaterials (KCC-1) with fibrous pore structure possess high DNA adsorption capacity and effectively deliver genes in vitro. Langmuir 2014, 30, 10886–10898.
[42] Zheng, H.; Wen, S.; Zhang, Y.; Sun, Z. Organosilane and polyethylene glycol functionalized magnetic mesoporous silica nanoparticles as carriers for CpG immunotherapy in vitro and in vivo. PLoS ONE 2015, 10, e0140265.
[43] Zhang, J.; Niemelä, M.; Westermarck, J.; Rosenholm, J. M. Mesoporous silica nanoparticles with redox-responsive surface linkers for charge-reversible loading and release of short oligonucleotides. Dalton Trans. 2014, 43, 4115–4126.
[44] Zhou, X.; Chen, L.; Nie, W.; Wang, W.; Qin, M.; Mo, X.; Wang, H.; He, C. Dual-responsive mesoporous silica nanoparticles mediated codelivery of doxorubicin and Bcl-2 siRNA for targeted treatment of breast cancer. J. Phys. Chem. C 2016, 120, 22375–22387.
[45] Wu, M.; Meng, Q.; Chen, Y.; Zhang, L.; Li, M.; Cai, X.; Li, Y.; Yu, P.; Zhang, L.; Shi, J. Large pore-sized hollow mesoporous organosilica for redox-responsive gene delivery and synergistic cancer chemotherapy. Adv. Mater. 2016, 28, 1963–1969.
[46] Yu, M.; Xue, Y.; Ma, P. X.; Mao, C.; Lei, B. Intrinsic ultrahigh drug/miRNA loading capacity of biodegradable bioactive glass nanoparticles toward highly efficient pharmaceutical delivery. ACS Appl. Mater. Interfaces 2017, 9, 8460–8470.
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
19 Nano Res.
[47] Kim, M. -H.; Na, H. -K.; Kim, Y. -K.; Ryoo, S. -R.; Cho, H. S.; Lee, K. E.; Jeon, H.; Ryoo, R.; Min, D. -H. Facile synthesis of monodispersed mesoporous silica nanoparticles with ultralarge pores and their application in gene delivery. ACS Nano 2011, 5, 3568–3576.
[48] Na, H. -K.; Kim, M. -H.; Park, K.; Ryoo, S. -R.; Lee, K. E.; Jeon, H.; Ryoo, R.; Hyeon, C.; Min, D. -H. Efficient functional delivery of siRNA using mesoporous silica nanoparticles with ultralarge pores. Small 2012, 8, 1752–1761.
[49] Desigaux, L.; Sainlos, M.; Lambert, O.; Chevre, R.; Letrou-Bonneval, E.; Vigneron, J. -P.; Lehn, P.; Lehn, J. -M.; Pitard, B. Self-assembled lamellar complexes of siRNA with lipidic aminoglycoside derivatives promote efficient siRNA delivery and interference. Proc. Natl. Acad. Sci. U.S.A. 2007, 104, 16534−16539.
[50] Svintradze, D. V.; Mrevlishvili, G. M. Fiber molecular model of atelecollagen-small interfering RNA (siRNA) complex. Int. J. Biol. Macromol. 2005, 37, 283−286.
[51] Li, Y.; Hei, M.; Xu, Y.; Qian, X.; Zhu, W. Ammonium salt modified mesoporous silica nanoparticles for dual intracellular-responsive gene delivery. Int. J. Pharm. 2016, 511, 689–702.
[52] Du, X.; Xiong, L.; Dai, S.; Kleitz, F.; Qiao, S. Z. Intracellular microenvironment-responsive dendrimer-like mesoporous nanohybrids for traceable, effective, and safe gene delivery. Adv. Funct. Mater. 2014, 24, 7627–7637.
[53] Li, X.; Zhang, J.; Gn, H. Adsorption and desorption behaviors of DNA with magnetic mesoporous silica nanoparticles. Langmuir 2011, 27, 6099–6106.
[54] Li, X.; Xie, Q. R.; Zhang, J.; Xia, W.; Gu, H. The packaging of siRNA within the mesoporous structure of silica nanoparticles. Biomaterials 2011, 32, 9546– 9556.
[55] Melzak, K. A.; Sherwood, C. S.; Turner, R. F. B.; Haynes, C. A. Driving forces for DNA adsorption to silica in perchlorate solutions. J. Colloid Interface Sci. 1996, 181, 635–644.
[57] Suma, T.; Miyata, K.; Anraku, Y.; Watanabe, S.; Christie, R. J.; Takemoto, H.; Shioyama, M.; Gouda, N.; Ishii, T.; Nishiyama, N.; et al. Smart multilayered assembly for biocompatible siRNA delivery featuring dissolvable silica, endosome-disrupting polycation, and detachable PEG. ACS Nano 2012, 6, 6693–6705.
[58] Gouda, N.; Miyata, K.; Christie, R. J.; Suma, T.; Kishimura, A.; Fukushima, S.; Nomoto, T.; Liu, X.; Nishiyama, N.; Kataoka, K. Silica nanogelling of environment-responsive PEGylated polyplexes for enhanced stability and intracellular delivery of siRNA. Biomaterials 2013, 34, 562–570.
[59] Naito, M.; Azuma, R.; Takemoto, H.; Hori, M.; Yoshinaga, N.; Osawa, S.; Kamegawa, R.; Kim, H. J.; Ishii, T.; Nishiyama, N.; et al. Multilayered polyion complexes with dissolvable silica layer covered by controlling densities of cRGD-conjugated PEG chains for cancer-targeted siRNA delivery. J. Biomater. Sci. Polym. Ed. 2017, 28, 1109–1123.
[60] Lin, D.; Cheng, Q.; Jiang, Q.; Huang, Y.; Yang, Z.; Han, S.; Zhao, Y.; Guo S.; Liang, Z.; and Dong, A. Intracellular cleavable poly(2-dimethylaminoethyl methacrylate) functionalized mesoporous silica nanoparticles for efficient siRNA delivery in vitro and in vivo. Nanoscale 2013, 5, 4291–4301.
[61] Xiong, L.; Bi, J.; Tang, Y.; Qiao, S. -Z. Magnetic core–shell silica nanoparticles with large radial mesopores for siRNA delivery. Small 2016, 12, 4735–4742.
[62] Ping, Y.; Guo, J.; Ejima, H.; Chen, X.; Richardson, J. J.; Sun, H.; Caruso, F. pH-responsive capsules engineered from metal-phenolic networks for anticancer drug delivery. Small 2015, 11, 2032–2036.
[63] Ejima, H.; Ricardson, J. J.; Liang, K.; Best, J. P.; van Koeverden, M. P.; Such, G. K; Cui, J.; Caruso, F. One-step assembly of coordination complexes for versatile film and particle engineering. Science 2013, 341, 154–157.
[64] Feng, W.; Nie, W.; He, C.; Zhou, X.; Chen, L.; Qiu, K.; Wang, W.; Yin, Z. Effect of pH-responsive alginate/chitosan multilayers coating on delivery efficiency, cellular uptake and biodistribution of mesoporous silica nanoparticles based nanocarriers. ACS Appl. Mater. Interfaces 2014, 6, 8447−8460.
[65] Yu, L.; Chen, Y.; Wu, M.; Cai, X.; Yao, H.; Zhang, L.; Chen, H.; Shi, J. “Manganese extraction” strategy enables tumor-sensitive biodegradability and theranostics of nanoparticles. J. Am. Chem. Soc. 2016, 138, 9881−9894.
[66] Hao, X.; Hu, X.; Zhang, C.; Chen, S.; Li, Z.; Yang, X.; Liu, H.; Jia, G.; Liu, D.; Ge, K.; et al. Hybrid mesoporous silica-based drug carrier nanostructures with improved degradability by hydroxyapatite. ACS Nano 2015, 9, 9614–9625.
[67] Suyver, J. F.; Aebischer, A.; Biner, D.; Gerner, P.; Grimm, J.; Heer, S.; Krämer, K. W.; Reinhard, C.; Güdel, H. U. Novel materials doped with trivalent lanthanides and transition metal ions showing near-infrared to visible photon upconversion. Opt. Mater. 2005, 27, 1111–1130.
[68] Yang, Y.; Liu, F.; Liu, X.; Xing, B. NIR light controlled photorelease of siRNA and its targeted intracellular delivery based on upconversion nanoparticles. Nanoscale 2013, 5, 231–238.
[69] Chang, Y. -T.; Liao, P. -Y.; Sheu, H. -S.; Tseng, Y. -J.; Cheng, F. -Y.; Yeh, C. -S. Near-infrared light-responsive intracellular drug and siRNA release using Au nanoensembles with oligonucleotide-capped silica shell. Adv. Mater. 2012, 24, 3309–3314.
[71] Fischer, D.; Li, Y.; Ahlemeyer, B.; Krieglstein, J.; Kissel, T. In vitro cytotoxicity testing of polycations: influence of polymer structure on cell viability and hemolysis. Biomaterials 2003, 24, 1121–1131.
[72] Li, X.; Chen, Y.; Wang, M.; Ma, Y.; Xia, W.; Gu, H. A mesoporous silica nanoparticle–PEI–fusogenic peptide system for siRNA delivery in cancer therapy. Biomaterials 2013, 34, 1391–1401.
H.; Di, W. Highly effective antiangiogenesis via magnetic mesoporous silica-based siRNA vehicle targeting the VEGF gene for orthotopic ovarian cancer therapy. Inter. J. Nanomedicine 2015, 10, 2579–2594.
[74] Ashley, C. E.; Carnes, E. C.; Epler, K. E.; Padilla, D. P.; Phillips, G. K.; Castillo, R. E.; Wilkinson, D. C.; Wilkinson, B. S.; Burgard, C. A.; Kalinich, R. M.; et al. Delivery of small interfering RNA by peptide-targeted mesoporous silica nanoparticle-supported lipid bilayers. ACS Nano 2012, 6, 2174–2188.
[75] Miyata, K.; Oba, M.; Nakanishi, M.; Fukushima, S.; Yamasaki, Y.; Koyama, H,; Nishiyama, N.; Kataoka, K. Polyplexes from poly(aspartamide) bearing 1,2-diaminoethane side chains induce pH-selective, endosomal membrane destabilization with amplified transfection and negligible cytotoxicity. J. Am. Chem. Soc. 2008, 130, 16287–16294.
[76] Suma, T.; Miyata, K.; Ishii, T.; Uchida, S; Uchida, H.; Itaka, K.; Nishiyama, N.; Kataoka, K. Enhanced stability and gene silencing ability of siRNA-loaded polyion complexes formulated from polyaspartamide derivatives with a repetitive array of amino groups in the side chain. Biomaterials 2012, 33, 2770–2779.
[77] Lee, H,; Jeong, J. H.; Park, T. G. A new gene delivery formulation of polyethylenimine/DNA complexes coated with PEG conjugated fusogenic peptide. J. Control. Release 2001, 76, 183–192.
[78] Moore, N. M.; Sheppard, C. L.; Barbour, T. R.; Sakiyama-Elbert, S. E. The effect of endosomal escape peptides on in vitro gene delivery of polyethylene glycol-based vehicles. J. Gene Med. 2008, 10, 1134–1149.
[79] Ngamcherdtrakul, W.; Morry, J.; Gu, S.; Castro, D. J.; Goodyear, S. M.; Sangvanich, T.; Reda, M. M.; Lee, R.; Mihelic, S. A.; Beckman, B. L.; et al. Cationic polymer modified mesoporous silica nanoparticles for targeted siRNA delivery to HER2+ breast cancer. Adv. Funct. Mater. 2015, 25, 2646–2659.
[80] Liu, J.; Jiang, X.; Ashley, C.; Brinker, J. Electrostatically mediated liposome fusion and lipid exchange with a nanoparticle-supported bilayer for control of surface charge, drug containment, and delivery. J. Am. Chem. Soc. 2009, 131, 7567–7569.
[81] Ashley, C. E.; Carnes, E. C.; Phillips, G. K.; Padilla, D.; Durfee, P. N.; Brown, P. A.; Hanna, T. N.; Liu, J.; Phillips, B.; Carter, M. B.; et al. The targeted delivery of multicomponent cargos to cancer cells by nanoporous particle-supported lipid bilayers. Nat. Mater. 2011, 10, 389–397.
[82] Graf, C.; Gao, Q.; Schütz, I.; Noufele, C. N.; Ruan, W.; Posselt, U.; Korotianskiy, E.; Nordmeyer, D.; Rancan, F.; Hadam, S.; et al. Surface functionalization of silica nanoparticles supports colloidal stability in physiological media and facilitates internalization in cells. Langmuir 2012, 28, 7598−7613.
[83] Souris, J. S.; Lee, C. -H.; Cheng, S. -H.; Chen, C. -T.; Yang. C. -S.; Ho, J. A.; Mou, C. -Y.; Lo, L. -W. Surface charge-mediated rapid hepatobiliary excretion of mesoporous silica nanoparticles. Biomaterials 2010, 31, 5564–5574.
[84] He, Q.; Zhang, Z.; Gao, F.; Li, Y.; Shi, J. In vivo
biodistribution and urinary excretion of mesoporous silica nanoparticles: Effects of particle size and PEGylation. Small 2011, 7, 271–280.
[85] He, X.; Nie, H.; Wang, K.; Tan, W.; Wu, X.; Zhang, P. In vivo study of biodistribution and urinary excretion of surface-modified silica nanoparticles. Anal. Chem. 2008, 80, 9597–9603.
[86] Sudimack, J.; Lee, R. J. Targeted drug delivery via the folate receptor. Adv. Drug Deliv. Rev. 2000, 41, 147–162.
[87] Arap, W.; Pasqualini, R.; Ruoslahti, E. Cancer treatment by targeted drug delivery to tumor vasculature in a mouse model. Science 1998, 279, 377–380.
[88] Schiffelers, R. M.; Ansari, A.; Xu, J.; Zhou, Q.; Tang, Q.; Storm, G.; Molema, G.; Lu. P. Y.; Scaria, P. V.; Woodle, M. C. Cancer siRNA therapy by tumor selective delivery with ligand-targeted sterically stabilized nanoparticle. Nucleic Acids Res. 2004, 32, e149.
[89] Lo, A.; Lin, C. -T.; Wu, H. -C. Hepatocellular carcinoma cell-specific peptide ligand for targeted drug delivery. Mol. Cancer Ther. 2008, 7, 579–589.
[90] Slowing, I. I.; Wu, C. -W.; Vivero-Escoto, J. L.; Lin, V. S. -Y. Mesoporous silica nanoparticles for reducing hemolytic activity towards mammalian red blood cells. Small 2009, 5, 57–62.
[91] Yu, T.; Malugin, A.; Ghandehari, H. Impact of silica nanoparticle design on cellular toxicity and hemolytic activity. ACS Nano 2011, 5, 5717–5728.
[92] Murugadoss, S.; Lison, D.; Godderis, L.; Van Den Brule, S.; Mast, J.; Brassinne, F.; Sebaihi, N.; Hoet, P. H. Toxicology of silica nanoparticles: an update. Arch. Toxixol. 2017, 91, 2967–3010.
[93] Morishige, T; Yoshioka, Y.; Inakura, H.; Tanabe, A.; Yao, X.; Narimatsu, S.; Monobe, Y.; Imazawa, T.; Tsunoda, S.; Tsutsumi, Y.; et al. The effect of surface modification of amorphous silica particles on NLRP3 inflammasome mediated IL-1β production, ROS production and endosomal rupture. Biomaterials 2010, 31, 6833–6842.
[94] He, Q.; Zhang, J.; Shi, J.; Zhu, Z.; Zhang, L.; Bu, W.; Guo, L.; Chen, Y. The effect of PEGylation of mesoporous silica nanoparticles on nonspecific binding of serum proteins and cellular responses. Biomaterials 2010, 31, 1085–1092,
[95] Cypryk, M.; Apeloig, Y. Mechanism of the acid-catalyzed Si-O bond cleavage in siloxanes and siloxanols. A theoretical study. Organometallics 2002, 21, 2165–2175.
[96] Cauda, V.; Schlossbauer, A.; Bein, T. Bio-degradation study of colloidal mesoporous silica nanoparticles: Effect of surface functionalization with organo-silanes and poly(ethylene glycol). Micropor. Mesopor. Mat. 2010, 132, 60–71.
[97] Lin, Y. -S.; Abadeer, N.; Hurley, K. R.; Haynes, C. L. Ultrastable, redispersible, small, and highly organomodified mesoporous silica nanotherapeutics. J. Am. Chem. Soc. 2011, 133, 20444–20457.
[98] Cauda, V.; Argyo, C.; Bein, T. Impact of different PEGylation patterns on the long-term bio-stability of colloidal mesoporous silica nanoparticles. J. Mater. Chem. 2010, 20, 8693–8699.
[99] Lin, Y. -S. Abadeer, N.; Haynes, C. L. Stability of small mesoporous silica nanoparticles in biological media. Chem.
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research
21 Nano Res.
Commun. 2011, 47, 532–534.
[100] Wang, D.; Xu, Z.; Chen, Z.; Liu, X.; Hou, C.; Zhang, X.; Zhang, H. Fabrication of single-hole glutathione-responsive degradable hollow silica nanoparticles for drug delivery. ACS Appl. Mater. Interfaces 2014, 6, 12600–12608.
[101] Maggini, L.; Cabrera, I.; Ruiz-Carretero, A.; Prasetyanto, E. A.; Robinet, E.; Cola, L. D. Breakable mesoporous silica nanoparticles for targeted drug delivery. Nanoscale, 2016, 8, 7240–7247.
[102] Huang, X. -C.; Wu, L. -B.; Hsu, J. -F.; Shigeto, S.; Hsu, H. -Y.; Biothiol-triggered, self-disassembled silica nanobeads for intracellular drug delivery. Acta Biomater. 2015, 23, 263–270.
[103] Zhang, Q.; Shen, C.; Zhao, N.; Xu, F. -J. Redox-responsive and drug-embedded silica nanoparticles with unique self-destruction features for efficient gene/drug codelivery. Adv. Funct. Mater. 2017, 27, 1606229.
[104] Xu, Z.; Zang, K.; Liu, X.; Zang, H. A new strategy to prepare glutathione responsive silica nanoparticles. RSC Advances 2013, 3, 17700–17702.
[105] Chen, A. M.; Zhang, M.; Wei, D.; Stueber, D.; Taratula, O.; Minko, T.; He, H. Co-delivery of doxorubicin and Bcl-2 siRNA by mesoporous silica nanoparticles enhances the efficacy of chemotherapy in multidrug-resistant cancer cells. Small 2009, 5, 2673–2677.
[106] Meng, H.; Mai, W. X.; Zhang, H.; Xue, M.; Xia, T.; Lin, S.; Wang, X.; Zhao, Y.; Ji, Z.; Zink, J. I.; et al. Codelivery of an optimal drug/siRNA combination using mesoporous silica nanoparticles to overcome drug resistance in breast cancer in vitro and in vivo. ACS Nano 2013, 7, 994–1005.
[107] Shen, J.; Liu, H.; Mu, C.; Wolfram, J.; Zhang, W.; Kim, H. -C.; Zhu, G.; Hu, Z.; Ji, L. -N.; Liu, X.; et al. Multi-step encapsulation of chemotherapy and gene silencing agents in functionalized mesoporous silica nanoparticles. Nanoscale 2017, 9, 5329–5341.
[108] Holohan, C.; Van Schaeybroeck, S..; Longley, D. B.; Johnston, P. G. Cancer drug resistance: an evolving paradigm. Nat. Rev. Cancer 2013, 13, 714–726.
[109] Skovsgaard, T. Transport and binding of daunorubicin, adriamycin, and rubidazone in ehrlich ascites tumour cells. Biochem. Pharmacol. 1977, 26, 215–222.
[110] Yu, C.; Qian, L.; Ge, J.; Fu, J.; Yuan, P.; Yao, S. C. L.; Yao, S. Q. Cell-penetrating poly(disulfide) assisted intracellular delivery of mesoporous silica nanoparticles for inhibition of miR-21 function and detection of subsequent therapeutic effects. Angew. Chem. Int. Ed. 2016, 55, 9272−9276.
[111] Zhanga, S.; Langer, R.; Traverso, G. Nanoparticulate drug delivery systems targeting inflammation for treatment of inflammatory bowel disease. Nano Today 2017, 16, 82–96.
[112] Florek, J.; Caillard, R.; Kleitz, F. Evaluation of mesoporous silica nanoparticles for oral drug delivery – current status and perspective of MSNs drug carriers. Nanoscale 2017, 9, 15252–15277.
[113] Moulari, B.; Pertuit, D.; Pellequer, Y.; Lamprecht, A. The targeting of surface modified silica nanoparticles to inflamed tissue in experimental colitis. Biomaterials 2008,
29, 4554–4560.
www.theNanoResearch.com∣www.Springer.com/journal/12274 | Nano Research