Functionalization of silica nanoparticles for nucleic acid ... · nucleic acids, such as plasmid...

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

http://www.thenanoresearch.com on May. 31, 2018

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Nano Research

https://doi.org/10.1007/s12274-018-2116-7

Functionalization of silica nanoparticles for

nucleic acid delivery

Rimpei Kamegawa1, Mitsuru Naito2, Kanjiro

Miyata1,*

1Department of Materials Engineering, Graduate

School of Engineering, The University of Tokyo,

Tokyo, Japan.

2Center for Disease Biology and Integrative

Medicine, Graduate School of Medicine, The

University of Tokyo, Tokyo, Japan

The present review describes the functionalization and performances of

silica nanoparticles for nucleic acid delivery. Their functionalities include

loading and programmed release of nucleic acids, active targeting,

endosome escape, and biocompatibility.

Author 1–3, http://www.bmm.t.u-tokyo.ac.jp/english/index.html

Functionalization of silica nanoparticles for nucleic

acid delivery

Rimpei Kamegawa1, Mitsuru Naito

2, Kanjiro Miyata

1 ()

1Department of Materials Engineering, Graduate School of Engineering, The University of Tokyo, Tokyo 113-8656,

Japan. 2Center for Disease Biology and Integrative Medicine, Graduate School of Medicine, The University of Tokyo, Tokyo

113-0033, Japan

Received: day month year

Revised: day month year

Accepted: day month year

(automatically inserted by

the publisher)

© Tsinghua University Press

and Springer‐Verlag Berlin

Heidelberg 2014

KEYWORDS

silica nanoparticle,

mesoporous silica, silica

coating, nucleic acid, drug

delivery

ABSTRACT

Silica nanoparticles (SiNPs) have been widely engineered for biomedical

applications, such as bioimaging and drug delivery, because of their high

tunability, which allows them to achieve the desired function. The present

review describes the functionalization and performances of SiNPs for nucleic

acid delivery. Nucleic acids, including plasmid DNA (pDNA) and small

interfering RNA (siRNA), have been highlighted as the next generation of

molecular target drugs for the treatment of intractable diseases. However, their

low bioavailability requires delivery systems that avoid nuclease attack and

kidney filtration to ensure the efficient access to the target cell cytoplasm or

nucleus. First, the biological significances of nucleic acids and the functions

required for their successful delivery are described. Then, the SiNPs designed

for nucleic acid delivery are introduced according to their functionalities,

including nucleic acid loading and releasing, cellular uptake, endosomal

escape, and biocompatibility. Additionally, the codelivery potential of SiNPs is

described. Finally, current challenges and future directions of SiNPs for

advanced nucleic acid delivery are proposed.

1 Introduction

Recently, silica nanoparticles (SiNPs) have been

vigorously engineered for applications in bioimaging

and drug delivery because their structures and

surface properties can be precisely controlled. As a

promising example for bioimaging applications, a

cancer‐targeted ultra‐small SiNP has been developed

and is currently being evaluated in phase I/II clinical

trials in the USA (ClinicalTrials.gov identifiers:

NCT02106598) [1, 2]. This nanoparticle has a

diameter of 7 nm, contains cyanine 5 dyes in the core,

and is surface‐modified with oligo(ethylene glycol)

(molecular weight (MW): 500 Da) terminated with

radio‐labeled cyclic arginine–glycine–aspartic acid–

tyrosine (124I‐cRGDY) peptide ligands for

fluorescence‐ and radio‐based dual cancer imaging.

Before initiating dose escalation phase I clinical trials,

a single dose of the targeted SiNP was administered

intravenously to five patients with metastasic

Nano Research

DOI (automatically inserted by the publisher)

Address correspondence to Miyata K. [email protected]

Review Article

| www.editorialmanager.com/nare/default.asp

2 Nano Res.

melanoma to investigate the biodistribution and

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

syndrome (ClinicalTrials.gov identifier:

NCT03108664) and homozygous familial

hypercholesterolemia (ClinicalTrials.gov identifier:

NCT02963311), which involve the silencing of the

transient receptor potential cation channel subfamily

V member 1 and proprotein convertase

subtilisin/kexin type 9, respectively. Many programs

tested for the treatment of liver diseases and cancers

have utilized delivery systems, such as ligand

conjugates and lipid nanoparticles (LNPs), taking


4 Nano Res.

advantage of their targeting abilities, as described

below. Notably, the phase III clinical trial of Patisiran,

which is an LNP formulation comprising cationic and

PEGylated lipids, has been completed for the

treatment of transthyretin‐mediated familial

amyloidosis, and its application for new drug

approval was submitted to the US FDA in 2017 [26].

Micro RNA (miRNA), a naturally occurring

noncoding or untranslated RNA molecule whose

structure resembles that of siRNA, can also suppress

the gene expressions in a sequence‐specific manner

through RISC formation [24]. In this mechanism,

miRNA hybridizes a series of partially

complementary mRNAs, which contrasts with the

mechanism involving siRNA that occurs through the

complete matching with the target mRNA. The lack

or downregulation of specific miRNA often causes

the cells to become abnormal, even cancerous, and

thus, miRNA replacement can be envisaged as a

potential therapeutic modality.

Another example can be found in the

single‐stranded DNA, i.e., ASO and its analogs,

which exhibit approximately 20 bases. It was

originally reported in 1978 that a chemically

modified ASO was capable of hybridizing

complementary mRNA to induce the mRNA

degradation by RNase H, which recognizes

DNA/RNA hybrids as a substrate [27]. This silencing

functionality of ASO has been recently expanded for

the silencing of miRNA as anti‐miRNAs.

Furthermore, additional “RNase H‐free” mechanisms,

i.e., exon skipping and exon inclusion, have been

described more recently. When bound to premature

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


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.


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


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

blocking of the entrance of the pores during the

initial loading process. To further increase the

occupied ratio, one interesting study developed a

dendritic SiNP (DSN) featuring center‐radial pores

[15]. This DSN had a quite high loading capacity of

pDNA (80 µg pDNA/mg DSN) [51] and siRNA (158

µg siRNA/mg DSN) [52]. The occupied ratios in DSN

were calculated to be ~40% for siRNA, which is

approximately 10 times larger than that of LMSN.

3.1.2 Hydrogen bonding. Small nucleic acids can be

loaded into small pores of MSNs through hydrogen

bonding using chaotropic salt solutions containing

guanidine hydrochloride at acidic pH [53, 54]. Thus,

a high concentration of guanidine hydrochloride (e.g.,

0.67 M) effectively shields the negative charges of the

nucleic acids and MSNs, thereby attenuating their

electrostatic repulsion. The guanidine ions capture

water molecules strongly, which leads to the

dehydration of nucleic acids and MSNs. Moreover,

most of the silanol groups of MSNs are protonated

because of the acidic condition. These circumstances

prompt the formation of hydrogen bonding between

nucleic acid phosphates and MSN silanols [55].

Following this procedure, siRNA was effectively

incorporated into the pore structure of MSNs under

the chaotropic condition with 66.7% of ethanol,

which facilitated the dehydration of siRNA [54]. A

high loading capacity of 30 µg siRNA/mg MSN was

achieved. In contrast, a negligible amount of siRNA

was adsorbed on the outer surface of nonporous

SiNPs under the same condition probably because of

the energetically unfavorable adsorption of the rigid

cylindrical architecture of siRNA on the spherical

surface with a diameter of 140 nm.

3.1.3 Silica coating. Although small nucleic acids can

be completely enclosed into the pores of MSNs, there

is a substantial difficulty in the complete enclosing of

large nucleic acids, such as pDNA. As a result of this

incomplete adsorption, such large nucleic acid

payloads may not be sufficiently protected from

nuclease attack. In addition, the binding of large

nucleic acids onto SiNPs may induce the formation of

intermolecular aggregates. To overcome these

bottlenecks, a facile approach has been developed for

the loading of pDNA into SiNPs by means of the

silica coating of cationic nanoparticles containing

nucleic acids (Fig. 4) [56–59]. To this end, polyion

complexes (PICs) were first prepared by mixing

nucleic acids and cationic polymers in aqueous

media. Then, the PICs were coated with a silica layer

by simple mixing of the PIC solution with a silicate

solution. The pDNA (or siRNA) was thereby

completely encapsulated within the silica layer.

When the concentration of silicate was relatively low

(<5 mM), large aggregates were formed because of

the neutralization of the charge by the reaction of

positively charged PICs and anionic silicates. In

contrast, monodispersed silica‐coated PICs (sPICs)

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).


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


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


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,

leading to adverse side effects. To overcome this

drawback, cationic nanoparticles developed for

administrations have been covered with

biocompatible macromolecules. To this end, PEG is

the most commonly used macromolecule because of

its neutral, hydrophilic, and flexible nature.

Generally, PEGs with an MW of 2,000–5,000 are

conjugated to the primary amines installed on the

SiNP via NHS ester coupling reaction [73, 79].

Notably, vascular endothelial growth factor‐targeted

siRNA‐loaded MSN modified with KALA‐installed

PEG chains suppressed the tumor growth by a

volume ratio of one sixth as compared with the case

of saline in orthotopic ovarian tumor‐bearing mice

when total 17.5 nmol of siRNA was intravenously

injected [73]. PEG can also be incorporated on the

silica surface through the electrostatic interaction

between the anionic silica surface and the cationic

segments of the block copolymers of PEG to prepare

mPIC by simply mixing SiNPs with the block

copolymers [57, 59]. It is worth noting that mPIC

without any ligands showed ~50% gene silencing for

vascular endothelial growth factor in a subcutaneous

human renal cell carcinoma (OS‐RC‐2) model via

intravenous injection (1.25 mg/kg mouse) [57].

Another interesting approach to install PEG is SiNPs

coated with a PEGylated lipid bilayer, termed

“protocells” [74]. In turn, the protocell is prepared by

mixing siRNA‐loaded LMSN and PEGylated

liposome, followed by centrifugation [74, 80, 81].

The size of PEGylated SiNPs was demonstrated to

remain unaltered in cell culture medium containing

10% fetal bovine serum, whereas large aggregates

were formed from non‐PEGylated cationic SiNPs [82].

It should be noted that PEGylation also contributes to

avoiding the entrapment of nanoparticles by the

reticuloendothelial system (RES) in the liver and

spleen during the circulation. When fluorescently

labeled cationic MSNs with a diameter of 80 nm were

administered intravenously, a large amount was

found to accumulate in the liver, which was followed

by biliary excretion 60 min after injection [83]. This

result suggests that the cationic MSNs underwent

opsonization by serum proteins, resulting in RES

entrapment. In contrast, PEGylated SiNPs smaller

than 120 nm showed a longer blood circulation

property, with their half‐life being 180 min, which

allowed for significant tumor accumulation in a

human squamous carcinoma xenograft model [3, 84,

85], avoiding RES entrapment. These reports

demonstrate that PEGylation is an essential approach

for the systemic delivery of drugs and nucleic acids

by nanoparticles.

Although the PEGylation of nanoparticles greatly

contributes in reducing nonspecific interactions with

biological components, the uptake by target cells is

concurrently compromised, generating the so‐called

PEG dilemma. In this context, the ligand‐mediated

active targeting strategy has been applied to enhance

the selective uptake by target cells. Thus, surface

modification with a ligand molecule, which

specifically binds to a receptor overexpressed on the

target cell membrane, enables the increase in specific

cellular uptake of SiNPs. For example, folic acid (FA)

has been used for the active targeting of various

cancers, including breast cancer cells, which

overexpress the folate receptor on their surface [86].

To introduce FAs on SiNPs, MSNs were first

functionalized with thiol groups (MSN‐SH) by a

modified Stöber method using


11 Nano Res.

(3‐mercaptopropyl)trimethoxysilane (MPTMS) [44].

The MSN‐SH was then reacted with 2,2′‐dipyridyl

disulfide to obtain a pyridyl disulfide‐capped MSN

(MSN‐SS‐Py), which was subsequently modified

with 3‐mercaptopropionic acid to obtain a

carboxylated MSN (MSN‐COOH). This MSN‐COOH

was then conjugated covalently with PLL‐grafted PEI

(PEI‐PLL) copolymers (MSN‐PP) via amidation.

Ultimately, the MSN‐PP was conjugated with

folate‐PEG‐COOH via

1‐ethyl‐3‐(3‐dimethylaminopropyl)carbodiimide

(EDC)/NHS coupling between the carboxyl groups of

folate‐PEG‐COOH and the amino groups of PEI‐PLL

(MSN‐PPPFA) (Fig. 5(a)). The as‐prepared

MSN‐PPPFA was efficiently uptaken by cultured

breast cancer (MDA‐MB‐231) cells, twice as much as

the MSN‐PP. Importantly, the uptake amount of

MSN‐PPPFA was decreased to the same level as that

of MSN‐PP in the presence of free FA as a competitor.

These results indicate that the enhanced uptake of

MSN‐PPPFA was ascribed to the FA ligands on the

surface and their binding to the FA receptor on the

breast cancer cells. Finally, approximately 80% of

B‐cell lymphoma 2 (Bcl‐2) gene was silenced in the

cultured breast cancer cells when 75 nM

Bcl‐2‐targeted siRNA was transfected with

MSNs‐PPPFA/siRNA at a weight ratio of 20/1. In

other studies, the cRGD peptide was used as a cancer

targeting ligand, which binds to v3 and v5

integrins overexpressed on the surface of various

cancer cells and cancer‐related angiogenic

endothelial cells [87, 88]. Our group also developed a

cRGD peptide‐installed SiNP based on the

aforementioned mPICs [59]. The mPICs were, in turn,

prepared by the surface coating of sPICs with a block

copolymer of cRGD‐installed PEG and PAsp(TEP).

Interestingly, the number of cRGD ligands per

particle was regulated from 150 to 530, depending on

the silicate concentration used in the preparation of

sPICs (Fig. 5(b)). The most efficient cellular uptake of

siRNA payloads was obtained for the mPIC having a

larger number of cRGD ligands, which suggests that

the multivalent binding between cRGD ligands and

the integrin receptors plays a critical role in the

enhancement of the cellular uptake. Ultimately, the

optimized cRGD‐installed mPICs achieved 3‐fold

higher uptake efficiency in cultured cervical cancer

(HeLa) cells compared with control mPICs without

cRGD. Ultimately, cRGD‐installed mPIC induced 30 %

gene knockdown in cultured HeLa cells at 100 nM

siRNA. In another study, a synthetic peptide 94

(SP94), which is composed of 12 amino acid residues

and which was identified by phage display, was

installed in the aforementioned protocell for

targeting hepatocellular carcinoma cells [74, 89] (Fig.

5(c)). After conjugation of

succinimidyl‐[(N‐maleimidopropionamido)‐tetracosa

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


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


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


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


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

to siloxane by PEGylation may retard the

decomposition of SiNPs, decreasing the rate of

excretion from the body that is associated with

accumulative adverse effects. Therefore, SiNP‐based

delivery systems must be provided with enhanced

biodegradability. To this end, doping with inorganic

ions, such as calcium and manganese, and installing

environment‐responsive functional groups into the

silica matrices are promising approaches to accelerate

the degradation of SiNPs in response to the

intracellular environment.

Although this review mainly focused on

cancer‐targeted drug and nucleic acid delivery, it is

worth mentioning that SiNPs are also useful for

targeting other diseased sites, e.g., epithelial and

immune cells through oral route for the treatment of

inflammatory bowel disease (IBD). This disease is

often developed in young people, and an effective

treatment still remains to be established [111].

Recently, several genes have been identified as

potential therapeutic targets [111]. Indeed, some

studies have already reported the oral delivery of

small drugs by utilizing SiNPs [112]. One example is

a 140 nm‐sized MSN covalently modified with

5‐aminosalicylic acid (5‐ASA), which is an

anti‐inflammatory drug for the treatment of IBD


17 Nano Res.

[113]. This MSN was found to promote the burst

release of 5‐ASA in the presence of

pancreatin‐containing digestive enzymes, such as

trypsin and chymotrypsin, whereas no release was

observed in phosphate buffer (pH 6.8). Additionally,

MSN was bound to inflamed tissues six times larger

than normal tissues when orally administered in

2,4,6‐trinitrobenzenesulfonic acid‐induced colitis

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).

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