1
Polysaccharide-based nucleic acid nanoformulations
Koen Raemdoncka, Thomas F. Martensa,b, Kevin Braeckmansa,b, Jo Demeestera, Stefaan C. De
Smedta,*
a Laboratory
of General Biochemistry and Physical Pharmacy, Faculty of Pharmaceutical Sciences,
Ghent University, Harelbekestraat 72, B-9000 Ghent, Belgium
b
Center for Nano-and Biophotonics, Faculty of Pharmaceutical Sciences, Ghent University,
Harelbekestraat 72, B-9000 Ghent, Belgium
*Corresponding author during submission: Dr. Koen Raemdonck
koen.raemdonck@ugent.be
Tel: +32 9 2648095
Fax: +32 9 2648189
Faculty of Pharmaceutical Sciences, Ghent University
Harelbekestraat 72, B-9000 Ghent, Belgium
Corresponding author after submission: Prof. Dr. Stefaan C. De Smedt
stefaan.desmedt@ugent.be
Tel: +32 9 2648076
Fax: +32 9 2648189
Faculty of Pharmaceutical Sciences, Ghent University
Harelbekestraat 72, B-9000 Ghent, Belgium
2
Abstract
Therapeutic application of nucleic acids requires their encapsulation in nanosized carriers that enable
safe and efficient intracellular delivery. Before the desired site of action is reached, drug-loaded
nanoparticles (nanomedicines) encounter numerous extra-and intracellular barriers. Judicious
nanocarrier design is highly needed to stimulate nucleic acid delivery across these barriers and
maximize the therapeutic benefit. Natural polysaccharides are widely used for biomedical and
pharmaceutical applications due to their inherent biocompatibility. At present, there is a growing
interest in applying these biopolymers for the development of nanomedicines. This review highlights
various polysaccharides and their derivatives, currently employed in the design of nucleic acid
nanocarriers. In particular, recent progress made in polysaccharide-assisted nucleic acid delivery is
summarized and the specific benefits that polysaccharides might offer to improve the delivery
process are critically discussed.
Keywords
siRNA
plasmid DNA
Dextran
Chitosan
Hyaluronic acid
Cyclodextrin
Nanomedicine
Gene therapy
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Table of Contents
1.
2.
3.
Introduction ..................................................................................................................................... 5
1.1.
Nucleic acid nanotherapeutics ................................................................................................ 5
1.2.
Polysaccharides in nucleic acid delivery .................................................................................. 9
Dextran-based nucleic acid nanotherapeutics .............................................................................. 11
2.1.
Introduction to dextran ......................................................................................................... 11
2.2.
Dextran-based matrices ........................................................................................................ 12
2.3.
Dextran as an additive in nucleic acid nanotherapeutics ...................................................... 16
2.3.1.
Polyelectrolyte complexes with dextran sulphate ........................................................ 16
2.3.2.
Polyelectrolyte complexes with modified dextran........................................................ 17
2.3.3.
Dextran coated nucleic acid nanoparticles ................................................................... 19
Chitosan-based nucleic acid nanotherapeutics............................................................................. 20
3.1.
Introduction to chitosan ........................................................................................................ 20
3.2.
Chitosan-based matrices ....................................................................................................... 21
3.3.
Chitosan as a polycation in nucleic acid nanotherapeutics................................................... 23
3.3.1.
Polyelectrolyte complexes with native chitosan ........................................................... 23
3.3.2.
Polyelectrolyte complexes with modified chitosan ...................................................... 24
3.3.3.
Hybrid chitosan-based polyelectrolyte complexes ....................................................... 31
3.4.
4.
Nanoparticles coated with chitosan ...................................................................................... 31
Hyaluronic acid-based nucleic acid nanotherapeutics .................................................................. 33
4.1.
Introduction to hyaluronic acid ............................................................................................. 33
4.2.
Hyaluronic acid-based matrices ............................................................................................ 36
4.3.
Hyaluronic acid as an additive in nucleic acid nanotherapeutics.......................................... 37
4.3.1.
Hyaluronic acid polyelectrolyte complexes ................................................................... 37
4.3.2.
Hyaluronic acid-based conjugates ................................................................................. 38
4.3.3.
Hyaluronic acid core-shell particles............................................................................... 44
4.4.
Influence of hyaluronic acid molecular weight ..................................................................... 46
4
5.
6.
7.
Cyclodextrin-based nucleic acid nanotherapeutics....................................................................... 47
5.1.
Introduction to cyclodextrins ................................................................................................ 47
5.2.
Cyclodextrin polymer nanoparticles ..................................................................................... 48
5.2.1.
Polymers with cyclodextrin backbone........................................................................... 49
5.2.2.
Cyclodextrin modifications of pre-existing polymers .................................................... 52
Other polysaccharides in nucleic acid nanotherapeutics .............................................................. 54
6.1.
-Glucans ............................................................................................................................... 54
6.2.
Alginate.................................................................................................................................. 56
6.3.
Arabinogalactan, pullulan and pectin.................................................................................... 58
Conclusions and future perspectives ............................................................................................ 59
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1. Introduction
1.1. Nucleic acid nanotherapeutics
Macromolecular single and double-stranded nucleic acids, such as antisense oligonucleotides (AsON),
small interfering RNA (siRNA), and plasmid DNA (pDNA) show great therapeutic promise in the
treatment of a wide variety of pathologies including cancer and inflammatory, neurological,
cardiovascular or metabolic disorders [1,2]. This is clearly reflected by the many clinical trials
launched in the past two decades evaluating the safety and efficacy of nucleic acid therapeutics for
these different pathologies [1,3]. With gene therapy, the underlying genetic cause of a pathology
could be treated instead of standard symptom management. Where a dominant disease-causing
gene is overexpressed, sequence-specific gene silencing therapy via AsON or siRNA can block the
expression of the encoded protein on the post-transcriptional level. On the other hand, if a genetic
mutation causes deficient production of a vital protein, gene supplementation via pDNA could
potentially restore functional protein expression [4].
To attain a therapeutic phenotype, the aforementioned nucleic acid drugs require delivery into the
target cell, albeit at a distinct intracellular location. AsON and siRNA function at the posttranscriptional level through sequence-specific binding with their target messenger RNA (mRNA) in
the cytosol, opposite to pDNA that has to reach the cell nucleus to access the transcriptional
machinery [2]. Several physical delivery methods, such as electroporation, hydrodynamic injection
and ultrasound-assisted delivery, have shown to be successful in transferring naked nucleic acids to
target cells. However, these methods are generally quite impractical or limited to local delivery in
easily accessible tissues such as skin and muscle [4,5]. Another well-known strategy to assist nucleic
acid therapeutics in reaching their intracellular site of action, is formulating them in nanosized
carriers. Although these can be exploited for systemic delivery, drug loaded nanoparticles (NPs) still
encounter numerous obstacles en route to their intracellular target upon in vivo administration
(Figure 1) [2,6]. In general, NPs should protect the nucleic acid payload from enzymatic degradation,
improve its biodistribution toward the diseased tissue and allow delivery into the desired
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intracellular compartment of the target cells [2]. Typically, nucleic acid nanocarriers can be divided in
viral and nonviral vectors. Viruses are by far the most potent gene therapy vectors since they have
naturally evolved to acquire optimal nucleic acid delivery capabilities, also explaining their prominent
use in clinical trials [3,7]. Nonetheless, viral vectors have a limited payload capacity and their use in
viral gene therapy entails inherent safety risks [6,7]. Therefore, gene therapy research has
increasingly focused on optimizing the delivery of nucleic acids with synthetic nonviral delivery
agents such as (cationic) liposomal formulations and polymers [2]. Despite a lower transfection
efficiency compared to viral vectors, their low immunogenicity, versatility and ease of manufacturing
make nonviral nanomedicines promising alternatives for gene therapy [7].
The size, geometrical shape and surface chemistry of nanomedicines predominantly dictate their in
vivo behavior at the nano-bio interface [8,9]. Following intravenous delivery, NPs should be large
enough to avoid fast renal clearance via glomerular filtration, but still be able to cross the capillary
endothelium (extravasation) in order to reach the target tissue interstitium. For instance, passive
tumor targeting of drug-loaded NPs can be assisted by the so-called enhanced permeability and
retention (EPR) effect. Defective angiogenesis in rapidly growing tumors often entails the formation
of leaky blood vessels through which NPs can easily extravasate. The impaired lymphatic drainage
further contributes to the accumulation of the nanocarriers in the tumor extracellular matrix [10].
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Figure 1. Overview of the main biological barriers for nucleic acid therapy following intravenous
injection of nucleic acid nanoparticles (NPs). These should be large enough to circumvent renal
clearance, but small enough to be able to extravasate and accumulate in the target tissue. Moreover,
the nucleic acid NP should evade uptake by the mononuclear phagocyte system (MPS) and has to
protect the encapsulated nucleic acid against degradation. Once the carrier has reached the target
cell, it needs to deliver the therapeutic RNA or DNA to the cytoplasm. This generally involves cellular
uptake through endocytosis, followed by escape from the endosome and carrier disassembly in the
cytosol. When nuclear entry is required, the nuclear-envelope embedded nuclear pore complex
poses an additional barrier. ECM: extracellular matrix, NPC: nuclear pore complex. Adapted with
permission from ref [6]. Copyright  Elsevier.
Passive tissue accumulation is also correlated with blood circulation time. Thus NPs are envisioned
that remain sufficiently stable in the bloodstream and are able to evade fast removal from the
systemic circulation by the mononuclear phagocyte system (MPS). Decorating the NP surface with
hydrophilic polymers, such as poly(ethylene glycol) (PEG), is an established strategy to improve
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colloidal stability in the extracellular environment and to reduce recognition by specialized MPS
phagocytes [2].
In addition, ample control over the NP fate at the subcellular level following cellular uptake has
become increasingly important, aiming to amplify the percentage of the endocytosed nucleic acid
drugs that is effectively delivered into the cell cytoplasm or the nucleoplasm [11]. Cellular
internalization of nanomedicines can be mediated by adsorptive or receptor-mediated endocytosis.
While the former involves non-specific hydrophobic and/or ionic interactions with the cell-surface
proteoglycans, the latter requires modification of the NP surface with targeting ligands that can
specifically recognize and bind membrane-anchored receptor proteins. It is generally assumed that
nucleic acid nanotherapeutics are internalized via an endocytic uptake mechanism [2,11].
Mammalian cells have been shown to display distinct endocytic entry portals and there is emerging
evidence that the transfection efficiency is strongly linked to the uptake mechanism which on its turn
is influenced by the nanocarriers physicochemical properties. Following internalization, NPs typically
will accumulate in endolysosomes containing various acid hydrolases. To avoid nucleic acid
degradation, the nanocarrier should preferably escape from the endosomes into the cell cytoplasm
and release the encapsulated or complexed nucleic acid. Furthermore, when delivery into the cell
nucleus is required (e.g. for transgene expression), the nuclear envelope poses another important
barrier. Plasmid DNA or nucleic acid NPs generally are too bulky for translocation across the nuclear
envelope-embedded nuclear pore complex (NPC), that only allows passive diffusion of molecules
with a hydrodynamic diameter below 10 nm [11]. For this reason, the transgene expression obtained
via non-viral gene delivery systems is often dependent on cell division, during which the nuclear
envelope is disassembled [12].
Rational nanocarrier design and a fundamental insight into the underlying extra-and intracellular
barriers are needed to potentiate nucleic acid delivery across these barriers and maximize the
therapeutic benefit. Both natural and synthetic materials are currently investigated as building blocks
in biomedical nanotechnology. This review focuses on polysaccharides and their derivatives in
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nanocarrier design and details recent important contributions toward improved AsON, siRNA and
pDNA delivery. In particular, this review will highlight specific advantages ascribed to both
established and less well-known polysaccharides used as influential constituents in nucleic acid
nanocarriers.
1.2. Polysaccharides in nucleic acid delivery
Polysaccharides are defined as polymeric carbohydrate structures composed of repeating
monosaccharide units adjoined by glycosidic bonds. They form an important class of naturally
occurring biopolymers that can be obtained from various natural and abundant resources, such as
algae (e.g. alginate), plants (e.g. pectins, cellulose, cyclodextrins), microorganisms (e.g. dextran,
pullulan) and animals (e.g. chitosan, hyaluronan) [13]. Polysaccharides are often quite heterogeneous
in structure and chemical composition, distinguishing neutral or charged, linear or branched and low
or high molecular weight (Mw) polymers with varying hydrophilicity (Figure 2). In addition, the
chemical microenvironment in which these polymers are dissolved may greatly influence their
physicochemical identity and hence also their interaction with therapeutic nucleic acids. This
diversity in polysaccharide structure therefore also entails the formation of nucleic acid NPs with
varying biophysical properties [7], affecting the efficiency with which the various existing extra-and
intracellular barriers in nucleic acid delivery can be overcome [1,2,6]. In general, owing to their
natural origin, polysaccharides are often described as biodegradable, biocompatible and nonimmunogenic, which are attractive properties for pharmaceutical and biomedical applications. For
this reason, polysaccharides are also often employed in hybrid nanostructures to reduce the toxicity
of (synthetic) materials (e.g. poly(ethylene imine) or PEI). Various protocols exist for the production
of nucleic acid loaded polysaccharide NPs, for which the reader is referred to other published reviews
on the topic [13,14].
The in vivo biodistribution of nanocarriers can be markedly influenced by the incorporation of
specific polysaccharides. Many polysaccharides, e.g. chitosan, alginate and hyaluronic acid (HA), are
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excellent bioadhesive materials warranting their application in tissue engineering and mucosal drug
delivery [14]. Alternatively, the presence of a dense layer of polysaccharides in a brush-like
configuration on the nanocarrier surface (e.g. dextran and chitosan) may also lengthen their
circulation time in the bloodstream by reduced MPS recognition, as a result of steric shielding and
mitigated complement activation [14-16]. In addition, the ability to recognize particular
carbohydrate-binding cell-surface receptors confers the use of polysaccharides (e.g. hyaluronic acid)
or oligosaccharide motifs (e.g. galactosan side chains in pectins) as targeting moieties for nucleic acid
delivery to specific cell types [17,18]. Instead of evading MPS phagocytes, treatment of inflammatory
disorders often requires specific targeting to phagocytic cell types (macrophages, dendritic cells),
which is the key advantage of -glucan containing NPs that recognize the ‘dendritic-cell-associated Ctype lectin-1’ or dectin-1 receptor frequently expressed on antigen presenting cells (APCs).
Another advantage linked to the polysaccharide structure is the ease of chemical modification due to
the availability of various functional groups (e.g. hydroxyls, amines, carboxylic acids) on the glycosidic
units. Most often, modifications are introduced with the aim to overcome specific hurdles such as
insufficient nucleic acid binding, fast MPS clearance and/or endosomal escape. Frequently
researchers evaluated modification with PEI to overcome endosomal sequestration. It is believed
that the endosomal buffering by the amine groups in PEI result in osmotic rupture of the endosomes
and release of the endosomal content into the cytosol (often referred to as the proton sponge effect)
[2]. Cyclodextrins are of particular interest toward further derivatization because of their intrinsic
ability to form stable inclusion complexes with hydrophobic ‘guest’ molecules (see 5.1.). All of the
virtues listed above conjointly make polysaccharides promising and attractive biomaterials in many
drug delivery applications.
2. Dextran-based nucleic acid nanotherapeutics
2.1. Introduction to dextran
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Dextran is a highly water-soluble polysaccharide of bacterial origin, produced by Lactobacillus,
Leuconostoc and Streptococcus species. It is composed predominantly of -1,6-linked glucopyranose
units with low degree of 1,3-branching (Figure 2) [19]. Native dextran is characterized by a high Mw
and polydispersity, both of which can be tailored by controlled hydrolysis and subsequent
fractionation [19].
Figure 2. Chemical structures of polysaccharides frequently incorporated in nanocarriers for nucleic
acid delivery. Dextran consists of repeating -1,6-linked D-glucose units with low degree of -1,3branching. Chitosan is a positively charged polysaccharide with repeating D-glucosamine and Nacetyl-D-glucosamine units that are linked via -(1,4) glycosidic bonds. Hyaluronan is an anionic
polysaccharide composed of D-glucuronic acid and N-acetyl-D-glucosamine disaccharide building
blocks, linked via alternating -1,4 and -1,3-glycosidic bonds. Alginate is composed of alternating
blocks of β-(1,4)-D-mannuronic acid and α-(1,4)-L-guluronic acid. Pullulan consists of -1,6-linked
maltotriose units. Schizophyllan has a main chain of 1,3--D-linked glucose units with 1,6--Dglucosyl side groups every third glucose. For details on the more complex chemical structure of
pectin and arabinogalactan, the reader is referred to references as mentioned in section 6.3.
Dextran chains may also be produced using dextran sucrase for the transfer of D-glucopyranose from
sucrose to acceptor molecules [20]. Importantly, low Mw dextran in particular already entails a long
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history of clinical applications in humans, e.g. as plasma volume expander, to enhance peripheral
blood flow or as rheological excipient in artificial tears [19,21]. The combined advantages of its
hydrophilic character, biocompatibility, biodegradability and ease of chemical derivatization all relate
to its suitability as drug delivery biopolymer [22].
2.2. Dextran-based matrices
Different types of dextran-based NPs have been described in the literature with the aim to enhance
intracellular nucleic acid delivery. The group of Stefaan De Smedt aimed to design controlled-release
carriers for siRNA delivery based on dextran [23,24]. Cationic dextran hydrogel NPs (dextran
nanogels, dex-NGs) were successfully prepared by an inverse emulsion photopolymerization of
dextran hydroxyethyl methacrylate (dex-HEMA) in mineral oil. To incorporate cationic charges in the
nanoscopic hydrogel network, cationic methacrylate monomers were copolymerized with dex-HEMA.
An important feature of these hydrogel particles is their biodegradability under physiological
conditions, owing to the hydrolysable carbonate ester linking the HEMA moieties to the dextran
backbone [25]. Preformed dex-NGs could encapsulate siRNA based on electrostatic interaction with a
maximal loading exceeding 50 pmol of siRNA per µg of lyophilized dex-NGs [25]. Although an
excellent reporter gene knockdown was obtained in hepatoma cells, confocal fluorescence
microscopy revealed that a substantial fraction of the endocytosed siRNA loaded dex-NGs (siDEXNGs) accumulates in acidified organelles, likely endolysosomes. This prompted the authors to
investigate the impact of photochemical internalization (PCI) on the obtained RNAi effect [26]. PCI is
an established endosomolytic method that involves the use of amphiphilic photosensitizers (PS) that
accumulate in the membranes of endocytic vesicles [27]. In this study, mesotetraphenylporphine
carrying two sulfonate groups on adjacent phenyl rings (TPPS2a) was used. Upon illumination with a
specific light source, excitation of the PS compound induces the formation of reactive oxygen species
(ROS), primarily singlet oxygen (Figure 3A). This highly reactive intermediate can damage cellular
components, but the effect is mainly confined to the local production site owing to its short range of
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action and short lifetime [27]. This localized effect will therefore selectively disrupt the endosomal
membranes, releasing the entrapped macromolecules or NPs into the cytosol. Since its discovery in
1999 [28], PCI has been successfully applied to stimulate the cytosolic delivery of several types of
macromolecules (peptides, proteins, nucleic acids), incorporated in non-viral carrier systems [29]. It
was found that application of PCI, even several days post-transfection, was able to significantly
improve and lengthen the gene knockdown obtained with siDEX-NGs (Figure 3B and 3C). These data
indicate that PCI is able to liberate a fraction of the siRNA or siDEX-NGs that remain trapped in
intracellular organelles. Applying vesicular compartments as drug depots can thus be regarded as a
potential strategy to prolong the therapeutic response when endosomal escape is effect-limiting
[26]. Nevertheless, dex-NGs loaded with anti-tumor necrosis factor  (anti-TNF- siRNA also
mediated high gene silencing and minor toxicity or off-target transcriptional changes in LPS activated
macrophages, without assistance of any endosomolytic tool [30]. In this report, siDEX-NGs
outperformed other polymeric nanocarriers such as trimethylated chitosan (TMC), G4/G7 poly(amido
amine) (PAMAM) dendrimers and poly(DL-lactic-co-glycolic acid) (PLGA) NPs [30].
Naeye et al. extended this work by decorating siDEX-NGs with a hydrophilic PEG shell to develop a
drug delivery platform capable of intravenous siRNA administration [31]. Although PEGylating the
nanogels did not prevent the partial dissociation of the encapsulated siRNA in human plasma,
fluorescence single particle tracking (fSPT) measurements revealed that PEGylation was required to
prevent their aggregation. Because intravenous application not only implies the contact with plasma
constituents but also with millions of blood cells, the interactions between blood cells and
(PEGylated) siDEX-NGs were investigated [32].
None of the nanogel formulations caused significant erythrocyte lysis, but positively charged
nanogels were shown to induce platelet aggregation. Flow cytometry data confirmed that nanogels
hardly bind to erythrocytes while a clear charge dependent interaction with platelets and leukocytes
was observed, with positively charged nanogel again demonstrating significant cellular binding [32].
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Figure 3. Influence of photochemical internalization (PCI) on in vitro siRNA delivery via biodegradable
cationic dextran nanogels (dex-NGs). (A) Schematic graph clarifying the endosomolytic mechanism of
PCI with the amphiphilic photosensitizer TPPS2a. (B) Confocal micrographs visualize the improved
delivery of green fluorescently labeled siRNA (siGLO) into Huh-7 hepatoma cells. (C) The application
of PCI at a later time-point after transfection with siDEX-NGs enables prolonged gene silencing in
contrast to a cationic lipoplex formulation (Lipofectamine™ RNAiMAX). Reprinted and adapted with
permission from ref [26]. Copyright  Elsevier.
Tripathi et al. reported on a different type of cationic modification of dextran NPs. To this end,
dextran was first crosslinked with 1,4-butanediol diglycidyl ether. Partial oxidation of dextran
hydroxyl groups to form aldehydes enabled grafting of branched PEI (bPEI). The resulting dextran-gbPEI nanocomposites were able to electrostatically bind pDNA and displayed endosomal buffering,
likely contributing to endosomal escape. The gene expression profile was evaluated in vitro and in
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vivo and showed maximal transgene expression in the spleen of Balb/c mice, possible due to capture
by splenic MPS phagocytes [33].
In an attempt to bypass chemotherapeutic failure due to multidrug resistance (MDR) in
osteosarcoma, Susa and coworkers formulated siRNA targeting the ATP binding cassette transporter
B1 (ABCB1, MDR1) gene in lipid-modified dextran-based polymeric NPs [34]. Overexpression of
MDR1 may be the predominant factor that confers resistance to a broad spectrum of
chemotherapeutics in many types of cancer [35]. Dextran was chosen as the primary NP building
block by virtue of its biocompatibility and biodegradability. The dextran particles were modified with
stearyl moieties and equipped with a PEG coat. The authors show that RNAi-mediated MDR1
downregulation by dextran carriers in drug-resistant osteosarcoma cell lines reversed the MDR
phenotype and re-sensitized these cell lines to the cytotoxic effect of doxorubicin.
The group of Jean Fréchet reported on a tunable and modular particulate system, constructed from
acetal-modified dextran (Ac-DEX) and designed for protein antigen and vaccine adjuvant delivery to
antigen presenting cells (APCs) [36,37]. The hydrophobic acetal groups make the modified Ac-DEX
soluble in organic solvents but insoluble in water, allowing the feasible production of Ac-DEX based
micro-and nanoparticles for protein antigen encapsulation via standard emulsification methods [36].
These acid-degradable particles were further optimized for gene delivery in both phagocytic and nonphagocytic cells by incorporating small amounts of a degradable cationic polymer (poly--amino
ester or PBAE) in Ac-DEX particles [38]. For the intracellular delivery of siRNA via this polymeric
platform, spermine-modified dextran was applied for the preparation of Ac-DEX, instead of including
PBAEs during particle production (Figure 4). The authors argue that the cationic nature of spermine
enables a better complexation of the negatively charged siRNA and improved interaction with the
target cell membrane. Spermine is a tetravalent organic amine that is present in mammalian cells in
millimolar concentrations and is considered non-cytotoxic [39]. Following endocytic uptake, acidcatalyzed hydrolysis of the particles will occur in the endolysosomal compartments releasing the
siRNA. The siRNA is thought to reach the cell cytoplasm as a result of endosomal bursting by virtue of
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an amine-induced proton sponge effect. This is further potentiated by intraluminal osmotic pressure
build-up through endosomal accumulation of spermine-Ac-DEX degradation products [40].
Figure 4. Dextran modified with hydrophobic acetal and cationic spermine moieties (spermine-AcDEX) can be used to form NPs for siRNA encapsulation and intracellular delivery. The chemical
structure of spermine-Ac-DEX and a representative scanning electron micrograph of the resulting
siRNA-loaded NPs are shown. Reprinted with permission from ref [40]. Copyright  American
Chemical Society.
2.3. Dextran as an additive in nucleic acid nanotherapeutics
2.3.1. Polyelectrolyte complexes with dextran sulphate
Alternatively, the negatively charged dextran sulphate (DS) can be used to prepare nanosized
polyelectrolyte complexes (PECs) together with cationic natural or synthetic (bio)polymers, such as
chitosan (see below) [41] or PEI [42,43]. PECs result from direct interaction of oppositely charged
polyelectrolytes in solution and within PEC NPs polycations are the driving force for complexation of
negatively charged nucleic acids. PEC structure and stability are primarily influenced by the polymer
characteristics (Mw, flexibility, charge density) and the chemical environment (pH, ionic strength,
temperature) [14]. The rationale behind the use of (polysaccharide) polyanions as an additive in PEC
drug formulations is to create more stable nanocomplexes, to minimize polycation-induced toxicity
and/or to stimulate cellular interactions via carbohydrate-binding receptors [14,42,44-46]. PECs have
often been used for protein and small molecule delivery [14,41]. However, Cho et al. reported on
PECs composed of DS and the positively charged poly-L-arginine (PLR) polypeptide to encapsulate
siRNA targeting the epidermal growth factor receptor (EGFR) [44]. Given the small and rigid nature of
17
an siRNA duplex, this may result in less stable and larger complexes with polycations (polyplexes)
compared to pDNA polyplexes. The incorporation of polyanions, such as DS or HA, has shown to
result in more compact siRNA polyplexes [45]. In vitro gene silencing and in vivo tumor growth
inhibition in head and neck cancer cells was demonstrated by Cho et al. with their optimized PEC
formulation [44]. Intratumoral injection was employed in the in vivo xenograft mouse model instead
of intravenous injection, since the authors recognize that improved PEC stability in the extracellular
environment is critical before systemic administration can be considered [44]. In another report,
pDNA delivery in human corneal cells was achieved by using hybrid NPs built from cationized gelatin,
modified with the low Mw oligoamine spermine, and the polysaccharide polyanions DS or
chondroitin sulphate. The particles were further stabilized by ionic crosslinking of the cationized
gelatin with pentasodium tripolyphosphate (TPP), thus obtaining particles with a more hydrogel-like
structure. It was found that the hybrid particles modestly improved the in vitro toxicity profile when
compared with the particles prepared in the absence of polyanions, without interfering with their
transfection efficiency. Again, the in vivo colloidal stability and gene delivery performance in relevant
animal models remain to be elucidated for this gene delivery system [46].
2.3.2. Polyelectrolyte complexes with modified dextran
Many examples can be found in the literature describing drug delivery assisted by chemically
modified dextran conjugates [20]. In the context of nucleic acid delivery, cationic modifications of the
dextran backbone with natural or synthetic oligo- or polyamines are preferred to allow electrostatic
interaction and formation of stable polyplexes with therapeutic nucleic acids. Many research groups
investigated conjugates of dextran and PEI in an attempt to reduce the conspicuous toxicity of the
latter polymer and improve the stability of its polyplexes in the presence of serum through a
PEGylation-like shielding effect. Early reports mainly focused on high Mw linear or branched PEI,
however with limited success in confining PEI-mediated toxicity while maintaining gene transfer
activity or polyplex stability [47-49]. The cytotoxicity of PEI is dependent on the polymer architecture
18
and increases with increasing Mw [50,51]. It has been demonstrated that reduced toxicity and
consistent biological activity can coincide when assembling small non-toxic PEI units into larger
biodegradable polymeric structures which is a prerequisite for efficient nucleic acid complexation
into NPs [52,53]. For this reason, other groups investigated the grafting of low Mw
oligoethyleneimine (OEI) moieties onto (biodegradable) dextran chains. In general, this approach
resulted in a more satisfying outcome with regard to combining decreased cytotoxicity with
maintained or even increased NP integrity and intracellular gene transfer ability [54-57].
Instead of employing small synthetic cationic units, several reports describe the use of natural
occurring oligoamines in the design of cationic dextran conjugates. It is hypothesized that the use of
biogenic amines would increase the biocompatibility of the resulting polycation. Azzam et al.
synthesized a library of over 300 different polycations by grafting spermine oligoamines to a dextran
or arabinogalactan backbone. Only the dextran-spermine polycations of defined Mw were able to
efficiently transfect different cell lines in vitro [58]. The obtained dextran-spermine conjugates were
further modified with PEG [59] or hydrophobic oleate residues [60] to gain colloidal stability in the
extracellular environment and improve intracellular gene delivery, respectively. It was demonstrated
that the micellar form of oleate-modified dextran-spermine achieved better transfection efficiencies
in vitro at high serum concentrations (50%) in contrast to unmodified dextran-spermine [61]. In a
recent comparable approach, Yang et al. reported on the grafting of another biogenic oligoamine, i.e.
agmatine, to low Mw dextran chains. Further hydrophobic modification of the dextran hydroxyls was
performed with lauric acid, thereby resulting in a polymeric construct composed exclusively from
biocompatible endogenous molecules. The authors showed that both agmatine and fatty acid
conjugation contributed to the high reporter gene expression levels obtained in COS-7 and HEK 293
cells [62]. Agmatine is produced by decarboxylation of L-arginine via the mitochondrial enzyme
arginine decarboxylase and is widely distributed in mammalian tissues. Its involvement in diverse
physiological processes and its low in vivo toxicity both translate in a variety of potential clinical
applications of exogenous agmatine [63]. Thus, agmatine-modified dextran cannot merely be
19
regarded as an inert excipient since it should be taken into consideration that its degradation
products could entail unforeseen pharmacological effects. In another report, Thomas et al. described
the modification of dextran with protamine, which is a naturally occurring cationic oligopeptide
isolated from salmon sperm [64]. Protamines are highly specialized sperm chromosomal proteins
responsible for the compaction of DNA. They carry arginine rich motifs and peptide sequences that
may act as a nuclear localization signal (NLS), both contributing to a more favorable intracellular
pDNA delivery [65,66]. Dextran-protamine conjugates mediated high transfection efficiency in vitro
without significant cytotoxicity and appeared to be haemocompatible, providing opportunities
toward in vivo application [64]. Several other examples of cationized dextrans have been described in
the literature which will not be discussed in detail here [67-69].
2.3.3. Dextran coated nucleic acid nanoparticles
Polysaccharide coating of NPs has been considered as a biomimetic alternative to PEG in order to
provide them with so-called ‘stealth’ properties and improve their in vivo performance [16]. For
many types of polysaccharides (e.g. dextran, chitosan, pullulan, hyaluronic acid) such surface
coatings have been successfully demonstrated via various methodologies. Native dextran, charged
dextran (dextran sulphate or diethyl amino ethyl (DEAE)-dextran) and hydrophobically modified
dextran were evaluated to coat polymeric particles as well as liposomal vesicles. The deposition of a
dextran coat could modulate protein adsorption, improve colloidal stability and mitigate
complement activation [14-16]. In a recent report, layer-by-layer (LbL) polyelectrolyte coated
quantum dots (QDs), with a terminal polysaccharide outer layer (dextran sulphate (DS) or hyaluronic
acid (HA)) were evaluated for their in vivo biodistribution in mice. Both DS and HA terminated
particles exhibited significantly prolonged circulation times (elimination half-lives of 3.2 h and 8.4 h,
respectively) when compared with uncoated QDs. However, the particles capped with a DS outer
layer also showed increased liver accumulation, possibly due to the interaction with receptors (HARE
20
receptor, also see 4.1.) on the liver sinusoidal endothelial cells (LSEC) [70] and/or phagocytic uptake
by liver macrophages (Kupffer cells) triggered by the high negative charge of DS [70,71].
To date, only few reports are available in the literature describing dextran coated NPs in the context
of nucleic acid delivery. Early reports describe dextran-g-poly(L-lysine) coated poly(lactic acid) (PLA)
NPs and DEAE-dextran modified poly(alkylcyanoacrylate) NPs for pDNA and AsON delivery
respectively [72-74]. More recently, Delgado et al. published on the design of solid lipid nanoparticles
(SLNs), modified with dextran and protamine, for gene delivery [66,75]. An aqueous solution of
dextran was incubated with protamine and pDNA prior to mixing with preformed SLNs. In the
resulting gene delivery vector, dextran-protamine-DNA complexes are electrostatically adsorbed on
the SLN surface. The particles displayed significant transgene expression in vivo following ocular
administration in rats [66] and intravenous injection in mice [75] and gene transfer was more
successful when protamine and dextran were incorporated in the formulation. As mentioned above
(2.3.2), the arginine-rich polycationic protamine aids in DNA complexation and may enhance DNA
entry into the nucleus [65,66,75]. The presence of dextran on the particle surface is important
toward improved in vivo biocompatibility. In addition, dextran seemed to influence the cellular
uptake mechanism, leading to improved transfection efficiency by promoting clathrin-mediated
endocytosis [66,75].
3. Chitosan-based nucleic acid nanotherapeutics
3.1. Introduction to chitosan
Chitosan is obtained by partial deacetylation of chitin, which is the natural main structural
component of the crustaceans exoskeleton and the cell wall of fungi [76,77]. It is a linear and
positively charged polysaccharide with repeating D-glucosamine and N-acetyl-D-glucosamine units
that are linked via -(1,4) glycosidic bonds (Figure 2) [77]. Because of its cationic nature, chitosan is a
very popular candidate among natural polysaccharides for nucleic acid complexation and delivery
[13]. Additional advantages described for chitosan, in line with the other polysaccharides mentioned
21
in this review, are its low cytotoxicity, low immunogenicity and its biodegradability [13,78-80]. A
more specific beneficial feature that can be ascribed to chitosan and chitosan derivatives are their
mucoadhesive and permeation-enhancing properties, explaining the frequent use in mucosal drug
delivery and tissue engineering [81-84]. Irrespective of these advantages, the low water-solubility of
chitosan at physiological pH is an important limitation for its clinical use. Deacetylation of chitosan
exposes D-glucosamine primary amines with a pKa ~6.5, explaining why chitosan is only soluble in an
acidic aqueous environment [85]. Therefore, also for chitosan an important amount of effort has
been put forward to chemically modify this polysaccharide, e.g. by quaternization of the Dglucosamine moieties [86], in order to make it more amenable for drug delivery (see below). The
nucleic acid delivery performance of native chitosan and its derivatives is highly dependent on a wide
variety of formulation-related factors, e.g. chitosan Mw, degree of deacetylation, type of salt and NP
preparation technique. The influence of these parameters has recently been reviewed in detail by
Mao et al. [78]. For a more in depth view on chitosan-based nucleic acid delivery systems, we also
would like to refer the reader to a review by Buschmann et al., published in this same theme-issue of
Advanced Drug Delivery Reviews. Given the enormous wealth of literature reports describing
chitosan nanotherapeutics, this review will mainly focus on the most recent advances made with this
type of polysaccharide for various nanoarchitectures and various (chemical) modifications. Several
preparation techniques can be harnessed for the production of nanosized particles from chitosan
[84], among which covalent or ionic chitosan crosslinking (chitosan hydrogel NPs) and self-assembly
with nucleic acids (chitosan polyplexes) are most often reported in the context of nucleic acid
delivery.
3.2. Chitosan-based matrices
Therapeutic nucleic acids can be encapsulated in chitosan hydrogel nanoparticles (nanogels) in which
chitosan polymeric chains are interconnected via covalent or ionic crosslinkages. Although physical
crosslinking methods generally lead to 3D polymer networks of inferior mechanical strength, ionic
22
crosslinking is the preferred method as it avoids the use of cytotoxic crosslinking agents (e.g.
glutaraldehyde) and can occur under relatively mild conditions [13,14]. It involves the formation of
ionic inter-and intramolecular linkages between chitosan protonated amines with the aid of small
multivalent ionic molecules. As mentioned before, TPP has been most widely used as polyanion
crosslinker for chitosan [13,77,84]. Chitosan nanogels can be formed by merely mixing an alkaline
TPP phase with an acidic chitosan solution, paying proper attention to the chitosan:TPP ratio to yield
stable solidified NPs [77]. Chitosan nanogels may consist solely of chitosan or can be prepared in the
presence of other hydrophilic macromolecular compounds. In this way, ionic gelation has been
frequently applied to further stabilize chitosan-based polyelectrolyte complexes with various
polyanions and/or bioactive nucleic acids. The resulting chitosan NPs are thus held together by
cooperative electrostatic interactions with both nucleic acids and TPP, giving rise to more compact
nanostructures [77]. Moreover, it has also been demonstrated that the condensed crosslinked
polymer matrix allows a more sustained gene expression, ascribed to a more extended release time
of the pDNA [87,88].
The group of Maria Alonso pioneered TPP induced ionic gelation of chitosan to form nanoscopic drug
delivery carriers [77]. Csaba et al. formulated pDNA and short dsDNA oligonucleotides into
chitosan/TPP crosslinked NPs prepared with chitosan of varying Mw. Low Mw chitosan (10 kDa)
provided more compact nanocarriers (~100 nm) opposed to high Mw chitosan (125 kDa) on account
of the lower viscosity of the former polymer dispersion, in line with other reports in the literature
[89]. Importantly, also the efficiency of transfection seemed highly dependent on the chitosan Mw,
with low Mw chitosan/TPP NPs showing superior gene transfer in vitro. In addition, low Mw
chitosan/TPP particles displayed a marked transgene expression following intratracheal
administration in mice, albeit comparable with the corresponding low Mw chitosan polyplexes in the
absence of TPP [87]. This study provided the basis for further optimization of chitosan/TPP NPs for
siRNA and pDNA delivery. To further enhance the transfection efficiency, colloidal stability and the
toxicity profile of the original chitosan/TPP formulation, PEG moieties were grafted to the chitosan
23
backbone and HA was included. To prepare the hybrid NPs, a preformed HA/TPP mixture was added
to a chitosan-g-PEG solution under magnetic stirring. The resulting HA/chitosan-g-PEG/TPP nanogels
demonstrated efficient transgene expression and moderate gene silencing in HEK 293T cells, albeit
that gene suppression values were similar to the commercial lipofectamine 2000 formulation [90].
Katas and Alpar directly compared the physicochemical properties and in vitro RNAi activity of
chitosan-siRNA NPs prepared via three different methods: complex coacervation, ionic gelation in the
presence of siRNA and adsorption of siRNA onto the surface of preformed chitosan nanogels. They
found that siRNA entrapment in chitosan/TPP nanogels showed better reporter gene suppression
compared to the other preparation methods, likely due to the improved siRNA loading [91], thereby
corroborating earlier results from Maria Alonso’s group obtained with oligonucleotides [77].
In a combined effort to improve both nanocarrier-mediated cellular delivery and transgene
expression efficiency at the level of the pDNA itself, Gaspar et al. formulated compact supercoiled
(sc) pDNA topoisoforms in chitosan NPs by ionotropic gelation [92]. It is known that sc pDNA
outperforms its relaxed open circular (oc) and linearized counterparts with regard to transfection
efficiency [93], thus warranting further optimization of pDNA topology in gene transfer applications.
The authors succeeded in purifying sc pDNA via high throughput arginine affinity chromatography
and subsequently encapsulated the recovered pDNA into TPP induced chitosan nanogels.
Importantly, the mild reaction conditions for the preparation of the chitosan NPs only induced minor
topoisoform conversion, in contrast to what is reported in the literature for other NP manufacturing
techniques [92].
3.3. Chitosan as a polycation in nucleic acid nanotherapeutics
3.3.1. Polyelectrolyte complexes with native chitosan
The preparation of chitosan PECs is based on the process of complex coacervation in which
oppositely charged polyelectrolytes are mixed together. As highlighted earlier, chitosan has been
studied elaborately for nucleic acid complexation and condensation into polyplex NPs, where the
24
negatively charged therapeutic nucleic acid functions as counter polyion in the complex coacervate.
The most straightforward chitosan polyplex formulation is obtained by merely mixing native chitosan
with the nucleic acid therapeutics, preferably at acidic pH [94-96]. Klausner et al. applied ultrapure
chitosan oligomers for ocular gene delivery. Low Mw chitosan with a high degree of deacetylation
achieved in vivo demonstrated reporter transgene expression levels superior to PEI after intrastromal
injections in rat corneas [97].
However, the chitosan polymer chains should have a sufficient charge density at physiological pH to
ensure good complexation and stability of the resulting polyplex in the extracellular environment,
which is often markedly hampered by the partial protonation of the glucosamine groups [81,98].
Intracellular nucleic acid delivery is further restricted by limited endosomal escape, likely due to its
weak buffering capacity, making native chitosan less efficient in comparison with other cationic
polymers such as PEI [80,99]. To impart improved nucleic acid delivery properties to chitosan, many
alterations to its native structure have been proposed. The repeating glucosidic units in chitosan
contain two hydroxyl groups and one primary amine that are regarded as potential reactive sites for
chemical modifications, depending on the desired biomedical application [81].
3.3.2. Polyelectrolyte complexes with modified chitosan
To increase the aqueous solubility of chitosan and the colloidal stability of chitosan PECs at
physiological pH, quaternization of the primary amine was investigated by many groups. In the
context of nucleic acid therapy N,N,N-trimethyl chitosan chloride (TMC) is most often reported [100].
To further optimize the delivery properties of TMC, Verheul et al. reported on the preparation of
nanoparticles based on thiolated TMC [101]. The introduction of thiol groups enables the formation
of intra- and intermolecular disulfide bonds that can provide additional NP stability in high salt
conditions or in the presence of competing anions [101,102]. The presence of free thiol groups also
enables further functionalization with e.g. thiolated PEG or HA [102] and may lead to enhanced
muco-adhesive properties by formation of disulfide bonds with extracellular mucin glycoproteins.
25
Moreover, the latter interaction could stimulate cellular internalization and therefore also gene
transfer, which was demonstrated by Lee et al. in various cell lines in vitro and in vivo in mice [88].
To further address the issue of forming siRNA polyplexes with inferior colloidal stability relative to
pDNA, the concept of polymeric siRNAs was recently proposed as an alternative strategy towards
more stable siRNA polyplexes [103]. It is expected that increasing the Mw and total number of
anionic charges by virtue of siRNA polymerization would result in more stable and condensed
nanosized complexes with polycations due to increased multivalent electrostatic interactions. Lee
and coworkers demonstrated the polymerization of thiolated siRNA (poly-siRNA) via reducible
disulfide linkages that can be selectively cleaved again in the reductive intracellular environment
following cytosolic delivery in the target cell [104]. However, when poly-siRNA was incubated with
cationic thiolated glycol chitosan (TGC) in HEPES buffer (pH 8), unstable complexes were formed as a
result of the weak charge interactions. Nevertheless, the electrostatic complexation enabled
additional intra-and intermolecular disulfide bonding within the TGC polymer chains, leading to
better condensed nanostructures (Figure 5). The TGC polyplexes displayed improved resistance
against polyanion destabilization and showed superior tumor accumulation over PEI polyplexes
following systemic injection in tumor-bearing mice. Moreover, significant vascular endothelial
growth factor (VEGF) silencing and concomitant reduction in tumor growth was observed for the TGC
polyplexes [104].
Recognizing the key role of intracellular vector unpackaging and nuclear entry in determining the
efficacy of transgene expression, Zhao et al. focused on the modification of chitosan with short
phosphorylatable peptides to overcome these impediments [105]. As mentioned before,
uncomplexed pDNA and nucleic acid NPs are too large to reach the nucleoplasm, given that the
nuclear-envelope embedded NPC only allow passive diffusion of molecules with a hydrodynamic
diameter below 10 nm [11].
26
Figure 5. (A) Production of polymeric siRNA (poly-siRNA) via disulfide groups. (B) Thiolated glycol
chitosan (TGC) can form polyplex structures with poly-siRNA via electrostatic interaction. Subsequent
crosslinking of TGC thiol groups leads to more condensed NPs. The intracellular release of
monomeric siRNA is mediated by glutathione (GSH) stimulated reduction of the disulfide bonds.
Reprinted with permission from ref [104]. Copyright  Wiley-VCH.
NLS containing proteins are recognized by nuclear transport receptors like importins and transportins
that actively enhance protein translocation across the NPC. Anchoring NLS peptides to pDNA and
even cationic polymers is believed to aid in nuclear delivery of transgenes, albeit with limited success
[106,107]. Therefore, a peptide was selected carrying a N-terminal Simian virus 40 large T antigen
(SV40 T-ag) derived NLS sequence and a serine-containing motif that could be selectively
phosphorylated by kinases present in nuclear cell lysates. The C-terminal end of the selected peptide
consisted of two additional arginines to increase the isoelectric point and enable the formation of
stably condensed DNA polyplexes at physiological pH. Following nuclear translocation, it was
hypothesized that the introduction of anionic phosphate groups by virtue of nuclear kinase activity
on the oligopeptide, would entail specific intranuclear polyplex dissociation. The peptide modified
chitosan polyplexes showed modest improvement in gene transfer efficiency when compared to
lipofection [105]. In this report the authors mainly focused on intracellular polyplex stability and the
nuclear import of polyplex and/or transgene, but unfortunately did not consider the importance of
the endosomal barrier in nucleic acid delivery. This issue is addressed by many others, given the fact
27
that the often observed low transfection efficiency of chitosan is thought to be related to its less
efficient escape from endosomes [78]. It is indeed known that the endosomal buffering capacity of
chitosan is significantly lower opposed to PEI [80,99]. However, in this theme issue Buschmann et al.
also describe that when comparing PEI with chitosan based on molar ionisable amines instead of
mass, fully deacetylated chitosan actually may have a larger endolysosomal buffering capacity than
PEI. In addition, some reports claim several other mechanisms by which chitosan could overcome the
endosomal barrier, e.g. its direct charge-dependent membrane destabilizing effect or indirect as a
result of endosomal enzymatic degradation. In the latter case, accumulation of chitosan degradation
products would evoke an increase in osmotic pressure eventually leading to perturbation of the
endolysosomal membrane [78].
Nevertheless, to bestow chitosan with improved endosomal destabilizing properties, both polymers
and small molecules have been grafted to the chitosan backbone, mainly with the aim to obtain a
cationic polymer capable of buffering the lumen of endolysosomes. In line with chemical
modifications carried out on other polysaccharides already described in this review, several authors
reported on the modification of chitosan or chitosan derivatives with PEI [108-113]. However, this
may again raise concerns regarding the cytotoxicity of the resulting conjugates. As an alternative,
Ghosn et al. used a 1-step carbodiimide reaction to modify the glucosamine functional groups of
chitosan with imidazol-4-acetic acid (IAA), thus introducing secondary and tertiary amines and
concurrently also endosomal buffering capacity [114]. A similar approach was used by Pego and
coworkers [99,115]. In vitro transgene expression and gene silencing using chitosan-IAA polyplexes
with pDNA and siRNA respectively was significantly improved over unmodified chitosan [116]. To
prevent aggregation of the polyplexes in vivo, PEG chains with a Mw of 5000 Da were anchored to
the surface of the polyplexes via a succinimidyl ester. PEGylated chitosan-IAA polyplexes mediated
substantial gene silencing in lung and/or liver following intranasal and intravenous administration
[114]. Chang et al. modified chitosan with pending histidine amino acids of which the imidazole ring
confers endosomal buffering capacity [117], a modification also recently exploited for dextran [67].
28
Likewise, urocanic acid that also bears an imidazol ring, was coupled to chitosan to increase its
transfection efficiency [118].
In an attempt to mimic the cellular translocation mediated by cell penetrating peptides (CPPs),
guanidinylated chitosan (GCS) was synthesized for gene and siRNA delivery [119,120]. CPPs have
demonstrated the ability to deliver peptides and nucleic acids in to cells. The most extensively
studied CPPs are characterized by their high content of positively charged arginines and lysines,
which are believed to play a major role in the cellular binding and intracellular delivery of
biomacromolecules. Although there is no consensus yet on their specific mechanism of cellular
internalization, it is believed that CPPs can assist bulky macromolecular payloads and even NPs in
crossing the cellular barriers [121]. The guanidinium group of arginine amino acid residues seems
imperative for cellular translocation, thereby providing a rationale for GCS evaluation as nucleic acid
nanocarrier [122]. It was demonstrated by Zhai et al. that GCS could improve DNA condensation at
neutral pH and improved cellular uptake when compared to native chitosan, both likely mediated by
the strong basic character of the guanidinium group that replaces the glucosamine primary amine
[119]. Unfortunately, no data on the cellular uptake mechanism and/or intracellular trafficking of the
guanidinylated polyplexes was provided. Correspondingly, GCS was used to complex siRNA in a
follow-up study investigating the potential of the GCS carrier toward pulmonary gene silencing via
intratracheal administration. To further improve pulmonary targeting, the 2-adrenergic receptor
(2-AR) agonist salbutamol was ligated to the GCS polymer. Evidence suggests that 2-AR agonists
could be used as a targeting ligand to ameliorate in vivo receptor-mediated delivery of
nanotherapeutics in the lung [123]. The salbutamol functionalized GCS particles achieved increased
gene silencing over non-targeted GCS polyplexes in the lungs of EGFP transgenic mice [120].
Several other examples of modified or native chitosan, equipped with appending cellular targeting
moieties, can be highlighted [110]. For instance, mannosylated chitosan-g-PEI polyplexes could be
targeted to APCs bearing mannose lectin receptors, as demonstrated in vitro in a macrophage cell
line [124]. Galactosylated chitosan was synthesized toward targeted gene transfer in hepatocytes via
29
the asialoglycoprotein receptor [125,126]. In another report, folate was applied as a targeting ligand
in the context of pulmonary tumor targeting [127]. Recently, CD7-specific single-chain antibodies
were conjugated to chitosan for targeted siRNA delivery to T-cells [128]. The group of Anil Sood
succeeded in developing a cyclic Arg-Gly-Asp (RGD) peptide anchored chitosan NP for siRNA delivery
(Figure 6) [129]. The peptide was coupled to the chitosan backbone by thiolation of glucosamine
residues using the crosslinking reagent N-succinimidyl-3-(2-pyridyldithio)-propionate (SPDP). The
RGD peptide specifically binds to v 3 integrin receptors often overexpressed in tumors. In addition,
the v3 integrin receptor is also selectively expressed by endothelial cells of the tumor vasculature,
making it an attractive binding site for anti-angiogenic therapy [130]. The NPs showed distinct tumor
(vascular) targeting in vivo in distinct orthotopic murine models of ovarian carcinoma. Tumor
selective delivery of siRNA, targeting genes involved in ovarian cancer progression, led to marked
inhibition of tumor growth [129].
The group of Ana Pego further built on their imidazol-modified chitosan to design a nanocarrier
capable of targeting the peripheral nervous system (PNS) [131]. In addition to imidazol grafting,
remaining glucosamine moieties were reacted with 2-iminothiolane. Thiolation was pursued to
enable grafting of the targeting ligand via a bifunctional PEG spacer. The non-toxic carboxylic
terminal fragment of tetanus toxine (HC fragment) was chosen as targeting moiety because of its
ability to mediate neuron cell-specific internalization. Although specific targeting was demonstrated
in a neuroblastoma cell line and a primary dorsal root ganglia (DRG) neuron culture model, the
percentage of transfected primary neurons was markedly reduced when compared to the HC
functionalized PEI-based vector[131,132]. It remains to be clarified if the targeted chitosan-based
formulation will be able to target the PNS in vivo and if sufficient expression levels can be achieved to
attain a therapeutic benefit in peripheral neuropathies [131].
30
Figure 6. In vivo gene silencing of periostin (POSTN), involved in cell survival and invasion, via siRNA
loaded chitosan nanoparticles (CH-NP) or integrin-targeted RGD-CH-NP. (A) The cyclic RGD peptide is
coupled to the chitosan backbone via disulfide coupling chemistry. (B and C) A single intravenous
injection of POSTN siRNA/RGD-CH-NP (0.15 mg/kg) resulted in significantly improved POSTN silencing
compared with the non-targeted formulation in an SKOV3ip1 (v3 integrin positive) murine model
of ovarian cancer, as assessed by both western blotting (B) and immunohistochemistry (C). Reprinted
and modified with permission from ref [129]. Copyright  American Association for Cancer Research.
31
3.3.3. Hybrid chitosan-based polyelectrolyte complexes
Chitosan PECs have also been constructed containing other polyanions (e.g. polysaccharides and
polypeptides) in addition to the nucleic acids [13,133-135]. For instance, Liao et al. demonstrated
enhanced gene silencing with ternary CS/siRNA/poly--glutamic acid complexes [134]. In another
example, hybrid NPs of chitosan and HA could instigate gene transfer into primary chondrocytes for
the treatment of osteoarthritis. The presence of HA in the particles is believed to enhance receptormediated uptake via interaction with the CD44 cell-surface receptor (see 4.1), that is highly
expressed on osteoarthritic chondrocytes [136]. As briefly mentioned in 2.3.1., the polyanions may
also bring about the added advantage of improved colloidal stability as a result of multivalent ionic
interactions, which is of particular interest for siRNA containing formulations [137].
3.4. Nanoparticles coated with chitosan
In addition to the nanocarriers described above, in which chitosan mainly governs NP formation and
nucleic acid complexation, chitosan may function as a coating material for surface modification of
various types of NPs. For instance, many groups reported on the modification of biodegradable and
biocompatible poly(D,L-lactic-co-glycolic acid) (PLGA) NPs with chitosan [138]. Because of its cationic
nature, chitosan easily binds with the negatively charged PLGA giving rise to core-shell NPs endowed
with a positive zeta-potential. Mostly, chitosan is applied during the emulsion evaporation method
used to prepare PLGA particles. It is thus postulated that the chitosan polymer chains can be found
both within the PLGA polymer matrix as on its surface. This surface modification can strongly
enhance the particles affinity for negatively charged cellular membranes and trigger subsequent
endocytic uptake. Moreover, because of chitosan affinity for mucosal surfaces and its permeation
enhancing properties, NPs equipped with a chitosan coat are attractive for mucosal delivery
[14,16,77]. The group of Claus-Michael Lehr developed chitosan-modified PLGA NPs to deliver
therapeutic oligonucleotides, adsorbed onto the surface of preformed chitosan-coated NPs, to lung
cancer cell lines [139,140]. Yuan et al. evaluated chitosan-PLGA nanoparticles for siRNA delivery into
32
HEK 293T cells. Since the siRNA was pre-complexed with chitosan prior to emulsification, this
formulation strategy will likely encapsulate a fraction of the siRNA in the PLGA core [141]. Tahara and
coworkers optimized an emulsion solvent diffusion method without sonication to prepare chitosancoated PLGA NPs. The encapsulation efficiency of siRNA in the PLGA matrix was amplified by its precomplexation with cationic lipids (1,2-dioleoyl-3-trimethylammonium-propane, DOTAP) and its
addition to the PLGA containing organic phase before emulsification in a chitosan supplemented
aqueous stabilizer solution [142].
Next to PLGA, other types of solid hydrophobic NPs have been used as template for chitosan coating.
Poly(isohexylcyanoacrylate) (PIHCA) with chitosan shell were used to complex siRNA targeting the
Ras homologous A (RhoA) protein that has been shown to promote cell proliferation and metastasis
[143]. Repeated intravenous injections every 3 days during a 30 day period of the anti-RhoA siRNA
formulation in nude mice inhibited xenografted breast cancer proliferation >90% at an siRNA dose of
0.15 mg/kg. Of equal importance, this treatment regimen was shown to be without in vivo toxicity
when compared with the untreated anesthetized animals as judged by body weight progression,
histological assessment of vital organs and quantification of biochemical markers for hepatic, renal
and pancreatic function [143].
More recently, de Martimprey and coworkers applied chitosan-coated poly(isobutylcyanoacrylate)
(PIBCA) NPs for in vivo siRNA delivery via intratumoral injection to treat papillary thyroid carcinoma
[144]. To enable intravenous delivery of this nanoformulation and passive EPR-mediated tumor
targeting, the authors succeeded in downsizing the PIBCA and PIHCA particles below 100 nm by
adding pluronic as a surfactant during NP synthesis. Multiple intravenous injections with a
cumulative siRNA dose of 5 mg/kg led to almost complete blockage of tumor growth in the same
murine tumor model [145]. The NPs were synthesized via redox radical emulsion polymerization as
opposed to earlier reports by the same group in which an anionic emulsion polymerization was
employed [146]. The former method has the advantage that the polysaccharide chains adopt a
33
distinct configuration on the NP surface that can bypass complement activation thus contributing to
lower recognition by the MPS [147].
Following the development of chitosan-coated liposomes (frequently termed chitosomes) for
mucosal delivery of macromolecules (e.g. oral delivery of insulin), some reports also addressed the
usefulness of these hybrid vesicular NPs for nucleic acid delivery [148,149]. For instance, the
pRc/CMV-HBs(S) plasmid, encoding the S region of the hepatitis B antigen, was encapsulated in
neutral liposomes consisting of 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE), egg
phosphatidylcholine (egg PC) and cholesterol [150]. Glycol chitosan was adsorbed onto the surface of
the liposomes and subsequently administered intranasally to mice to evaluate their suitability for
non-invasive immunization via the nasal mucosa. The glycol chitosan coated liposomes exhibited
humoral, mucosal and cellular immune responses that were superior to immunization with naked
pDNA [150].
4. Hyaluronic acid-based nucleic acid nanotherapeutics
4.1. Introduction to hyaluronic acid
Hyaluronic acid (HA) is a nonsulfated glycosaminoglycan (GAG), composed of alternating disaccharide
units of N-acetyl-D-glucosamine (GlcNAc) and D-glucuronic acid (GlcA), linked by alternating β-1,4
glycosidic bonds and β-1,3 glucuronidic bonds (Figure 2) [151]. Despite its plain appearance, its
functions and interactions are extremely diverse depending on its size. In mammalian organisms,
native HA is usually found as an unbranched, high Mw (several million Daltons) linear polymer with
exceptional physicochemical properties. In this high Mw form, HA is one of the main constituents of
the extracellular matrix (ECM) and is known to have a mechanical and structural role in the synovial
fluid, the vitreous humour of the eye and in connective tissues of e. g. the umbilical cord and dermis
[152]. Aside from these extracellular functions conveyed by the physicochemical properties of HA,
so-called HA binding hyaladherins extend the functionality of HA to the modulation of cellular fate by
receptor-mediated intracellular signaling. For example, HA has been suggested to influence cell
34
signaling to the cyclin D1 pathway [153], regulate the immune response [154] and influence both cell
proliferation and migration [155,156]. Some hyaladherins play important roles in the degradation
and uptake pathways of HA, mostly cell receptors in contact with the extracellular environment. To
date, already several of these HA-specific cell receptors have been identified, such as cluster
determinant 44 (CD44), receptor for hyaluronate-mediated motility (RHAMM), HA receptor for
endocytosis (HARE) and lymphatic vessel endothelial hyaluronan receptor-1 (LYVE-1), all with specific
functions. More detailed information about these various HA binding proteins and their respective
functions can be found in an extensive review recently published by Jiang et al.[157].
Nowadays, HA is easily extracted from rooster combs or produced by microbial fermentation [158].
In the medical field, non-immunogenic, highly viscous HA solutions are commercially available as a
surgical aid in ophthalmology and as viscosupplementation for synovial fluids in patients with
osteoarthritis [152]. The ubiquitous distribution of HA throughout the body and it’s intrinsic
properties such as hydrophilicity, biocompatibility, biodegradability and non-immunogenicity all
endorse the use of HA for biomedical needs. Aside from these advantages, several other
characteristics have encouraged the use of HA in drug delivery. The muco-adhesive properties of HA
were exploited to increase the residence time of small molecule therapeutics on ocular mucosa or in
topical wound healing [159,160], whereas the hydrophilic nature of HA was employed to reduce
unspecific interactions with proteins in the blood stream and prolong in vivo circulation time [161].
Nearly ten years ago, it was documented that the hydrophilic nature of HA could also serve as a
cryoprotectant in the preparation of unilamellar liposomal formulations [162]. When using liposomes
for drug delivery, the unilamellar structure is of utmost importance. Upon rehydration after
lyophilization, the emulsions usually revert back to multilamellar liposomes, yet it was found that the
covalent attachment of HA on the unilamellar liposomes could preserve their structure after
rehydration. Finally, one of the most promising advantages of HA is based on the previously
mentioned interactions with HA-specific cell receptors such as CD44, rationalizing HA as a targeting
strategy to CD44-(over) expressing tissues, such as tumors and diseased livers [163-165]. This
35
targeting strategy was elegantly documented in vitro by Surace et al. in HA receptor-expressing cells
compared to control cells (Figure 7) [166]. Taken together, it is widely acknowledged that HA can
provide several advantages to drug delivery, and in the next section we would like to highlight the
use of HA specifically for nucleic acid delivery vectors.
Figure 7: In vitro targeting properties of HA towards the CD44 receptor, as demonstrated by Surace
et al. Comparative transfection efficiency in MCF-7 (A) and MDA-MB231 (B) cells treated with
DE:DOPE lipoplexes (black) and lipoplexes containing 0.10 mg/mL of HA-DOPE conjugate (white) in
the presence of increasing amounts of anti-CD44 antibody Hermes-1 and anti-ErbB2 antibody used as
control. It is noticed that lipoplexes containing the HA-DOPE conjugate show an increased
transfection efficiency in HA receptor-containing MDA-MB231 cells compared to MCF-7 cells. This
increase in transfection efficiency is downregulated by increasing amounts of anti-CD44 antibody,
indicating a CD44 receptor-specific uptake in MDA-MB231 cells. Reproduced with permission from
ref [166].
36
4.2. Hyaluronic acid-based matrices
Since HA is advantageous in terms of biodegradability and non-immunogenicity, and given its use in
drug delivery [152,158], it would make sense to propose this biopolymer as a nucleic acid delivery
vector. Unfortunately, due to the negative charge of both HA and nucleic acids, nucleic acid NPs
based solely on HA are not that common. Regardless, some groups have reported on the ability of HA
matrices to incorporate therapeutic nucleic acids. HA hydrogels and microspheres were evaluated
for the sustained release of pDNA encoding for platelet derived growth factor (PDGF) [167] and βgalactosidase [168], respectively. The HA matrices were formed by an adipic dihydrazide (ADH)
crosslinking reaction, where pDNA and high Mw HA are mixed together, after which ADH and 1ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) are added to ensure crosslinking. Another group
employed a film of ADH-crosslinked HA as a physical barrier to prevent post-operative peritoneal
adhesions, where the film was also able to release previously incorporated pDNA encoding for
hyaluronan synthase 2 in the peritoneal cavity over a prolonged period of time [169]. In all three
studies, the sustained release of the pDNA was confirmed in vitro by degradation of the matrices in a
hyaluronidase solution. The released pDNA was shown to be intact and bioactive by in vitro
transfection assays using commercially available transfection agents. Consequently, aside from the in
vivo data by Yun et al. [168], it must be emphasized that in these studies only the bioactivity of the
released pDNA was assessed. This way, the macroscopic HA matrices are only evaluated in terms of
their ability to protect therapeutic DNA, not to ensure its uptake in the target cells, one of the most
difficult barriers to overcome in nucleic acid delivery. To do so, some groups have proposed the use
of HA hydrogels not only for their biocompatible features, but also for their known interactions with
HA-specific cell receptors. In contrast with the larger hydrogels discussed previously, Lee et al.
reported using an emulsion method and ultrasonication to fabricate nano-sized biodegradable thiolcrosslinked HA nanogels, with anti-GFP-siRNA physically entrapped during the emulsification process
[170]. The disulfide-linkages in the hydrogels were incorporated to ensure cargo release in the
reductive intracellular environment, as was demonstrated by the release of intact siRNA in a GSH-
37
concentration dependent fashion. Compared to cationic polymers such as PEI and PLL, the HA
hydrogels showed much lower cytotoxicity and higher gene silencing capabilities in the presence of
serum. Additionally, the purpose of HA to guarantee specific uptake of the hydrogels in HA receptorpositive cells was verified by a competitive binding assay in HCT-116 cells, where the co-incubation of
HA/siRNA-hydrogels with free HA drastically decreased the gene silencing effect, indicative of the
specific uptake of the hydrogels by HA receptor-mediated endocytosis [170] .
4.3. Hyaluronic acid as an additive in nucleic acid nanotherapeutics
4.3.1. Hyaluronic acid polyelectrolyte complexes
As mentioned before, HA on its own is not considered a good nucleic acid vector due to its negative
charge. Therefore, HA is mostly used as an additive to an existing nucleic acid carrier to endow the
latter with added advantages. As it is known that most nucleic acid vectors have a positive charge,
anionic HA can be electrostatically complexed with these cationic polymers and the nucleic acids to
form PECs by supramolecular self-assembly. By mixing low Mw HA with biodegradable PLR, PECs
capable of complexing siRNA were easily formed [171]. The authors proposed to incorporate HA in
the siRNA-vector to grant additional HA receptor-specific targeting capabilities. Besides an increased
colloidal stability in physiological conditions, the proposed targeting advantage was confirmed by an
in vitro uptake profile and in vivo gene silencing effect which was positively correlated with CD44
receptor density. Crosslinked hyaluronic acid-chitosan oligomer (HA-CSO) PECs were synthesized by
de la Fuente et al. to improve the gene delivery potential in the cornea by prolonging residence time
of gene delivery NPs on the ocular mucosa due to HA’s muco-adhesive nature [172]. The particles in
this study were prepared by an ionotropic gelation technique, where CSOs were mixed with a
solution containing HA, TPP and the desired pDNA. Furthermore, the authors propose that HA might
influence the intracellular trafficking of the gene carriers towards the perinuclear region of the cell,
given previously published data by Tammi et al. [173] and Evanko et al. [174]. This premise was
further studied by Contreras-Ruiz et al. by visualizing the uptake of the same crosslinked HA-CSO
38
hydrogels using live cell fluorescence microscopy [175]. It was shown that the NPs were effectively
taken up in an HA receptor-specific process. Further experiments with chemical inhibitors of
endocytic processes demonstrated that the particles were not taken up by clathrin-mediated
endocytosis, but rather by caveolin-mediated endocytosis, which was confirmed by colocalization
studies of NPs with caveolin [175]. As it has previously been reported that caveolin-dependent
endocytosis bypasses the lysosomal degradation pathway and delivers its cargo near the Golgi
apparatus or in the perinuclear region [176], the authors postulate the idea of HA as a targeting
ligand not only for HA-specific cell receptors, but also for intracellular compartments (Figure 8).
4.3.2. Hyaluronic acid-based conjugates
For a more stable and permanent availability of HA on the nucleic acid vector, some groups prefer to
covalently attach an HA-molecule to the building blocks of the NP before the self-assembly. Several
HA-conjugates have already been described in the literature. For example, Takei et al. conjugated
low Mw HA to PLL by a reductive amination reaction using sodium cyanoborohydride (NaBH3CN) as a
reductant [177]. The resulting copolymer was able to efficiently complex pDNA due to the cationic
PLL. The HA bestowed the carrier with in vivo HARE receptor-specific targeting properties, as
demonstrated by the enrichment of particles at liver sinusoidal endothelial cells (LSECs) after
intravenous injection in animal models. Usually though, the carboxyl-groups on the HA are activated
with EDC in acid buffer to form a highly reactive O-acylisourea, which can easily react in a basic
medium with a free amine to form a covalent amide-linkage. These free amine groups are frequently
encountered on most known non-viral polymeric nucleic acid vectors, such as PEI. By conjugating PEI
on a HA backbone via amide-bonds, Jiang et al. envisioned a non-viral nucleic acid carrier with the
target-specific and biocompatible advantages of HA, whilst retaining the ability to escape the
endosomes via the buffering effect of the amine-containing PEI [178].
39
Figure 8: Internalization pathway for hyaluronic acid (HA) – chitosan oligomer (CSO) NPs, as proposed
by Contreras-Ruiz et al.. The HA on the particles interacts with CD44 receptors in the plasma
membrane, triggering the caveolae-mediated internalization of the NPs. From the caveosome, NPs
are sorted to the endoplasmic reticulum (ER) and then to the nucleus, bypassing the lysosomal
degradation pathway. CD44 receptors and caveolin are then recycled and carried back to the
membrane through the Golgi network. The authors therefore envision HA as a targeting ligand not
only for the CD44 receptor, but also for the caveolae-mediated endocytic pathway. Reproduced with
permission from ref [175].
40
These conjugates were able to efficiently complex siRNA by means of electrostatic interactions with
the amine groups of PEI. The authors demonstrated the HA receptor-specific binding in vitro with
preferred uptake and enhanced silencing effects in LYVE-1-expressing cells (B16F1 cells) compared to
HEK 293 cells, which do not express HA-specific receptors.
To determine how the chemical conjugation of PEI to HA would affect its receptor-specificity, the
authors investigated in a follow-up study the receptor-mediated uptake of QD-modified HA in
different mol %, and showed an increased accumulation of QD-HA conjugates in the liver up to 35
mol % modification, and a non-targeted systemic distribution with higher mol % modifications
[178,179]. Less mol % modification with PEI resulted in the inability of the conjugates to form
complexes with siRNA due to the shortage of positive groups, which was demonstrated in their first
study [178]. To complement the in vitro targeting studies of the HA coated NPs, Jiang et al. further
investigated the particles in terms of in vivo targeting. After injection in xenografted C57BL/6 mice,
the authors note that the HA coated particles preferably accumulate in tissues with HA receptorexpressing cells, such as the liver, kidneys and tumors, in agreement with in vivo data on the QD-HA
conjugates [179]. The siVEGF-PEI/HA particles were significantly better at suppressing tumor growth
compared to uncoated siVEGF/PEI complexes [180]. Aside from the main advantage of HA receptor
specificity, the authors also noted that the HA PEI conjugates showed less cytotoxicity in vitro
compared to PEI alone, yet still more cytotoxic than HA alone.
To further optimize the nucleic acid carrier, the group proposed to use low Mw PEI crosslinked with
cystamine bisacrylamide (CBA). This disulfide-linked PEI is known to be less cytotoxic because of its
biodegradability in reductive environments, and Park et al. examined if an HA/SS-PEI conjugate
would still show the same targeting abilities. These optimized particles were employed in in vivo
studies for tumor therapy with anti-VEGF-siRNA and to treat liver cirrhosis with anti-transforming
growth factor (TGF)-β siRNA targeted to the liver [181,182]. In a first study, the authors
demonstrated an enhanced therapeutic effect of the siVEGF/SS-PEI/HA-nanoparticles after
intratumoral injection compared to siVEGF/SS-PEI-particles, indicating the preservation of tumor-
41
specific targeting (Figure 9) [182]. In a follow-up study, a therapeutic effect was noticed after
intravenous injection, which they attribute in part to the maintained targeting effects towards
cirrhotic livers, yet also to the decreased unspecific interactions with blood proteins and longer
circulation time on the other.
Figure 9: Anti-tumoral therapeutic effect of anti-VEGF-siRNA (siVEGF)/(PEI-SS)-b-HA complex in
female balb/c mice where CT-26 colon cancer cells were injected for tumor inoculation and growth.
(A) Tumor volume change with increasing time after intratumoral injection of a control of 5% glucose
solution, siVEGF/PEI-SS, non-specific Luc siRNA (siLuc)/(PEI-SS)-b-HA, and siVEGF/(PEI-SS)-b-HA
complexes. The treatments were performed three times after 8, 11, and 14 days. The results
represent mean ± SD (n = 3). (B) Photo-images of dissected tumor tissues after 20 days. Both in (A)
and (B), the increased therapeutic effect of HA-containing particles is clearly seen. Reproduced with
permission from [182]. Copyright  Elsevier.
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Figure 10. In vivo biodistribution of Cy5.5-labeled particles consisting of 5β-cholanic acid
(CA)/hyaluronic acid (HA)-conjugates. 2, 6 and 10 denote the degree of substitution (DS), defined as
the number of CA molecules per 100 sugar residues of HA. (A) In vivo non-invasive fluorescence
imaging of HA-NPs in tumor-bearing mice. Time-dependent whole body images of athymic nude mice
bearing SCC7 tumors after intravenous injection of HA-NPs. Arrows indicate the sites of tumors. (B)
Ex vivo fluorescence images of normal organs and tumors collected at two days post-injection of HANPs. (C) Quantification of the ex vivo tumor targeting characteristics of HA-NPs in tumor-bearing
mice. Error bars represent standard deviation (n = 3). Reproduced with permission from [183].
Copyright  Elsevier.
This enhanced biodistribution of HA-coated particles was correspondingly visualized in a study by
Choi et al., where HA-nanoparticles were intravenously injected in tumor-bearing mice and analyzed
at different time-points (Figure 10) [183]. In contrast to the HA/polycation-conjugates used by Park
et al., the HA particles in the study of Choi and coworkers had a hydrophobic segment incorporated,
such as 5β-cholanic acid, to further ensure the spontaneous self-assembly of core/shell-structured
nanoparticles. This has already been used in a drug delivery setting, where the authors hoped the
HA-shell would confer colloidal stability and mobility in the vitreous humour, the extracellular matrix
43
in the center of the eye [184]. Here, hydrophobic small molecule drugs could easily be loaded in the
hydrophobic core, but in the context of nucleic acid delivery the loading of hydrophilic nucleic acids
could pose a problem. Nevertheless, there have been reports of hydrophobation of nucleic acids with
CTAB [185], enabling its complexation in the hydrophobic core, although the possible toxic effects of
CTAB [146] was not further investigated in this study. A more elegant approach was proposed by
Shen et al., who added a cationic spermine segment to the HA block copolymers for the
complexation of siRNA [186]. In their study, the authors wanted to elucidate the uptake mechanism
of HA coated particles, and found that they were mostly endocytosed by caveolae-dependent
endocytosis. These findings nicely concur with the proposed hypothesis of Contreras-Ruiz and
colleagues to employ HA as a targeting ligand towards the caveosome [175]. Nevertheless, a
different study by Zaki et al. reported on the clathrin-dependent uptake mechanisms of both
uncoated and HA-coated nucleic acid NPs [187]. The differences between the findings of Zaki et al.
on the one hand [187], and Shen et al. [186] and Contreras-Ruiz et al. [175] on the other, could be
explained in part by the different cell types used or the size of the particles. Indeed, despite the fact
that the findings of Contreras-Ruiz et al. [175] agree with those of Shen et al. [186], the latter authors
attributed the caveolae-dependence to the size of the particles rather than claiming HA can be used
as a targeting ligand to an intracellular compartment. It has been reported that particle size can
influence the uptake pathway where larger particles around 200 – 500 nm in size would preferably
be endocytosed in a caveolae-dependent manner [188]. In short, the intracellular targeting of NPs
can’t solely be attributed to the presence of HA and regarding the previously mentioned studies,
some other remarks should also be taken in consideration. When interpreting data from chemical
endocytosis inhibitor experiments, care should always be taken. Aside from the risk of eliciting
cytotoxicity, the effects of these inhibitors are usually poorly characterized, non-specific and cell
type-dependent [11,189], all of which might influence the findings of the previously mentioned
studies.
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4.3.3. Hyaluronic acid core-shell particles
The conflicting results between caveolae-dependent endocytosis [175,186] and clathrin-dependent
endocytosis [187] could also be attributed to a different particle structure. Instead of incorporating
HA in a PEC or a covalently attached micelle-coat, Zaki et al. employed an electrostatic “post-coating”
strategy, where HA is added to a preformed crosslinked chitosan/TPP core to ensure HA availability
by means of an outer HA-shell. Indeed, this electrostatic complexation could prove beneficial, as it
had been suggested by Kim et al. that chemical modification could disturb the functionalities of
native HA [171]. Moreover, because the negative charge density of HA is not high enough to displace
the electrostatic bonds between the nucleic acids and the polymer, HA would be unable to displace
the nucleic acids or even disturb supramolecular self-assembled PECs [190]. Electrostatic coating of
preformed cationic PEI/pDNA polyplexes was documented by Hornof et al. [191], who justified a low
Mw HA coating by facilitating uptake in human corneal epithelial (HCE) cells via CD44-mediated
endocytosis. Wang et al. employed the same strategy to improve stability in physiological conditions
and specifically target HA receptor-expressing cells [192]. Again, HA was added to preformed
PEI/pDNA binary polyplexes, resulting in negatively charged ternary complexes which displayed
increased stability in physiological salt media compared to the uncoated polyplexes. Furthermore,
shielding of the polyplexes with HA resulted in a significant increase in transfection efficiency in HA
receptor-expressing HepG2 cells, while a decreased transfection efficiency was observed in HEK 293T
cells. Interestingly, the significant rise in transfection efficiency in the HepG2 cells could not be
related to an increased uptake. The lack of correlation between uptake and transfection efficiency
had already been reported by a previous study of Ruponen et al. [190], where it is suggested that
downstream intracellular processing mechanisms might contribute more to the final transfection
efficiency than uptake alone. Wang et al. attributed the increase in transfection efficiency to the
relaxation of DNA/polycation interactions which would allow the gene transcription machinery to
more easily reach the complexed DNA [192], a hypothesis previously postulated by Ito et al. [193]. In
45
their study, the authors describe the formation of ternary HA coated PEI/pDNA complexes and
specifically investigate the influence of HA on transcription of complexed DNA. They observe with in
vitro assays that HA increases both DNA transcription as well as DNA relaxation and conclude that HA
might display high mobility group (HMG)-like properties, which enhance transcription by granting the
transcription machinery easier access to the genes. Similarly, Xu et al. noticed an increase in
transfection efficiency independent of uptake when evaluating ternary HA/CLPEI/pDNA complexes in
NIH/3T3 cells [194]. These polyplexes consist of pDNA complexed by low Mw PEI-fragments
crosslinked with disulfide linkages (CLPEI), which are electrostatically coated with HA for increased
stability in the extracellular environment. The authors hypothesized that increase in transfection
efficiency was because of the HA which helps DNA unpacking triggered by the intracellular CLPEI
degradation, by further loosening the interactions between complexed DNA and low Mw
polycations. Likewise, it was observed by de la Fuente et al. that the combination of HA and low Mw
CSOs in their hydrogels resulted in very high transfection efficiencies, which was also explained by
the proposed role of HA as a transcription activator and facilitator of DNA unpacking [172].
Instead of polymers, other commonly used NA vectors are cationic liposomes [2]. The group of Dan
Peer has focused on the use of HA by covalently attaching it to the surface of preformed liposomes.
They have previously shown that liposomes with mitomycin D preferentially accumulate at CD44overexpressing tumor tissue when coated with high Mw HA [161]. They noticed that the effect of HA
is two-fold: (1) HA acts as a targeting ligand in vitro to cells overexpressing hyaluronan receptors; (2)
HA provides the particles with a longer circulation time in vivo by eluding unspecific interactions with
blood components. What is more, the authors discovered a new advantage of HA, being its use as a
cryoprotectant [162]. In a later study, the liposomes coated with HA are used for the in vivo delivery
to leukocytes of siRNA against cyclin D1 [195]. The targeting to the leukocytes was not done by using
the HA as a ligand, rather HA was used as an easily modifiable intermediate to graft targeting
antibodies. Furthermore, HA was used for liposomal stability during rehydration after lyophilization
and to prolong the circulation time in vivo.
46
Some groups argue that a post-coating as described before, where HA is covalently attached to
preformed liposomes, does not grant enough control over loading density which is deemed highly
necessary for proper targeting [166]. By conjugating high Mw HA molecules on DOPE lipids via an
EDC-facilitated amide-bond, liposomes with an outer HA shell could be made with sufficient control
of HA loading. These liposomes were evaluated in terms of delivery of both pDNA [166] and siRNA
[196]. Surace et al. prepared liposomes consisting of cationic lipid [2-(2,3-didodecyloxypropyl)hydroxyethyl] ammonium bromide (DE) and DOPE, complexed with pCMV-Luc plasmids [166]. These
lipoplexes were then evaluated in CD44-expressing MDA-MB-231 cells, as well as in MCF-7 control
cells with very low CD44 expression (Figure 7). Taetz et al. used DOTAP:DOPE liposomes to validate
HA receptor-targeted gene silencing of the HA coating in CD44-expressing A549 cells and in nonhyaladherin-expressing Calu-3 cells [196]. Both studies demonstrated a very low toxicity in all cell
lines, attributed to the negative surface charge. Also, the targeting properties of the NPs were
confirmed by an increased transfection efficiency in CD44-expressing cells, compared to a lower
transfection in the control cell lines.
In conclusion, the use and advantages of HA in the field of nucleic acid delivery are being actively
investigated in combination with many known nucleic acid delivery vectors and in many different
particle structures. Therefore, it should be noted that every method has its own merits and pitfalls,
yet a clear consensus on NP structure has not been found yet.
4.4. Influence of hyaluronic acid molecular weight
In the wealth of literature concerning the use of HA as an aid in drug or nucleic acid delivery, a lot of
contradictory findings have been documented. As previously stated, this could be explained in part
by a different NP structure or the different cell types used. However, the Mw of the HA molecules
should also be taken into account, because it has been documented that the functions of free HA in
the body is mostly determined by its Mw [197]. For example, high Mw HA chains are nonimmunogenic whereas low Mw HA oligomers are suspected to elicit immune-responses. It has also
47
been documented that HA oligomers have a tumor-suppressive ability [156], whereas HA of a
different Mw are thought to have a general stimulating effect of tumor growth [155]. These effects
are presumably modulated by HA-binding receptors overexpressed in the tumors, indicating the
interaction of HA with hyaladherins is also dependent on Mw of the HA molecule [18]. Yet, as a
targeting ligand towards the CD44 receptor, HA has been used in its low Mw form [171,191], its high
Mw form [166,196] and a middle Mw form [170]. Some groups acknowledge the alleged effect of
Mw on HA functions and have investigated accordingly. Hornof et al. have tried to optimize an
electrostatically shielded particle with different Mw’s of HA (<10 kDa, 10 – 30 kDa, 30 – 50 kDa)
[191]. The authors find that a low Mw coating (<10 kDa) is most adept at providing the carriers both
with colloidal stability in physiological media as with a targeting function. However, the HA fractions
used in this study only cover a very small part of the entire range of Mw, mostly in the low Mw
range. A more comprehensive study by Mizrahy et al. compared 5 different fractions, ranging from
low Mw 6.4 kDa HA to high Mw 1500 kDa HA, for the effects on innate immune response and CD44
targeting [198]. Contrary to the previous findings, they document that the binding affinity towards
the CD44 receptor is increased with increasing HA Mw. In conclusion, because it is generally accepted
that the different functions of HA in the body are solely regulated by its Mw, this factor should not be
overlooked when employing HA in a drug or nucleic acid delivery setting.
5. Cyclodextrin-based nucleic acid nanotherapeutics
5.1. Introduction to cyclodextrins
Cyclodextrins (CDs) can be defined as naturally occurring cyclic oligosaccharides, consisting of -1,4
linked D-glucopyranose units (Figure 2), that are produced by enzymatic conversion of starch. CDs
are characterized by a typical amphiphilic topology, with a hydrophilic exterior and an inner
hydrophobic cavity, which endows CDs with the capacity of forming molecular inclusion complexes
with hydrophobic ‘guests’ [199]. Depending on the number of glucose building blocks, one can
distinguish between a hexameric (-CDheptameric (-CD and octameric (-CD) form. In a
48
pharmaceutical context, CDs have been frequently applied for the formulation of poorly watersoluble drugs, to improve drug stability or to enhance drug permeability across biological membranes
[200]. However, a wealth of scientific reports available in the literature describes the implementation
of native or chemically modified CDs in nanocarrier design to improve the formulation of
biomacromolecules or enhance their delivery across the extra-and intracellular barriers. More than a
decade ago, a body of evidence already indicated the potential applications of CDs in therapeutic
formulations of peptides, proteins and oligonucleotides [201,202]. It was shown that the
complexation with CDs entailed improved cellular delivery, resistance to endonucleases and reduced
immunogenicity [199,200,202]. Currently, CD containing delivery systems, mostly composed of CD
containing polymers, are of interest for gene and siRNA delivery. Besides cationic CD polymers, also
CD-based polyrotaxanes, CD-based dendrimers and monodisperse CD derivatives have been used for
pDNA and/or siRNA delivery, which will not be discussed in further detail here. Recent
comprehensive reviews specifically on CD-mediated drug delivery are available [199,200,203].
5.2. Cyclodextrin polymer nanoparticles
The concept of using CD-containing polymers (CDPs) as drug carriers already exists for several
decades and they have been explored in the field of gene delivery since 1999 [200,204]. The general
added benefit of incorporating a CD moiety in a pharmaceutical nanocarrier is two-fold [199]: (1) CDs
have the ability to act as membrane absorption enhancers and (2) the unique feature to form CD
molecular inclusion complexes enables multifunctional surface modification of nanocarriers to
improve in vivo stability and cellular targeting, the latter being of particular interest in relation to NA
NPs. CDPs in gene therapy can be structurally divided in polymers with CDs in their backbone or preexisting polymers to which CD derivates are grafted. A general feature of these polymers is that they
are positively charged to allow electrostatic complexation with nucleic acid therapeutics.
49
5.2.1. Polymers with cyclodextrin backbone
The group of Mark E. Davis was the first to synthesize CDPs with polycationic signature for gene
delivery [200,205-209] and significant advances have been made since then. The cationic CDPs were
able to condense pDNA (~5 kbp) into nanosized polyplexes (CDPlexes) that achieved equal in vitro
transfection efficiencies compared to PEI and Lipofectamine™ [204]. Further optimization of the
polymer structure and its transfection activity was attained by capping the polymer ends with
imidazol groups [210]. It is hypothesized that this modification bestows the CDPlexes with endosomal
buffering capacity that would allow enhanced endosomal escape of the transgene complex on
account of the proton-sponge effect [211]. Moreover, the authors claim that the imidazol-modified
CDPlexes show enhanced intracellular unpackaging which could also contribute to the observed
improvement in transgene expression [210]. To optimize CDPs for systemic administration, it was
demonstrated by Pun and Davis that CDPlexes built from linear cationic CDPs and pDNA could be
furnished with a PEG hydrophilic shell. This was achieved through the addition of adamantane
modified PEG chains (Ad-PEG) to preformed CDPlexes, resulting in the formation of Ad-CD inclusion
complexes at their surface. PEGylation of polyplexes prevents their aggregation in the systemic
circulation and minimizes non-specific interactions [2]. In addition, targeting of specific cell types can
be made possible by coupling receptor ligands to the distal end of the Ad-PEG chains. CDPlexes
coated with galactosylated Ad-PEG were specifically internalized by hepatoma cells through
receptor-mediated endocytosis via the asialoglycoprotein receptor [212]. Likewise, the use of
transferrin (Tf) as a targeting ligand enables efficient internalization of CDPlexes in malignant cells
overexpressing the transferrin receptor [213]. Davis and coworkers also pioneered the translation of
this concept towards siRNA delivery (Figure 11) [214-216]. Ad-PEG-Tf modified CDP complexes with
siRNA (siCDPlexes) targeting the ribonucleotide reductase subunit M2 (RRM2) showed significant
tumor growth inhibition in murine models [217-220] and dose-escalating studies in non-human
primates demonstrated that multiple systemic NP administrations, with doses up to 9 mg siRNA/kg,
were well tolerated. The delivery concept was termed RONDEL™ (“RNAi/Oligonucleotide
50
Nanoparticle Delivery”, Calando Pharmaceuticals) and was applied in the first in-human phase I
clinical trial with a targeted siRNA containing NP, administered to patients with solid cancer
refractory to standard-of-care therapies [221]. Targeted NPs were detected in post-treatment tumor
biopsy sections and significant RRM2 downregulation was observed both on the mRNA and protein
level. Moreover, in one patient sample 5’-RLM-RACE (RNA Ligand Mediated – Rapid Amplification of
Complementary DNA Ends) PCR demonstrated the presence of an RNAi-specific RRM2 mRNA
cleavage fragment, indicating that the in vivo delivered anti-RRM2 siRNA can effectively activate the
RNAi pathway [221].
Figure 11. (A) Chemical structure of -cyclodextrin containing polycation, end-capped with imidazol
groups (CDP-im), for in vivo delivery of siRNA and pDNA. (B) Schematic representation of the ionic
self-assembly of siRNA and CDP-im into nanosized particles and their functionalization with
poly(ethylene glycol) (PEG) chains and targeting ligands (e.g. transferrin, Tf) via the formation of
inclusion complexes with adamantane (Ad) (RONDEL™ delivery technology). Adapted from ref
[214,221].
51
Likewise, this delivery platform was exploited by Brahmamdam et al. in a murine cecal ligation and
puncture (CLP) model of sepsis. Immune suppression through apoptotic loss of competent immune
effector cells significantly contributes to morbidity and mortality in this pathology. In vivo delivery of
RONDEL™ formulated siRNA to splenic CD4+ T and B lymphocytes, targeting the cell death proteins
Bim and PUMA, could significantly diminish the apoptotic depletion of these immune cells and
reversed sepsis-induced immune suppression [222]. Regardless of PEGylation however, it was
demonstrated in mice, monkeys and humans that the siRNA/CDP NPs are rapidly cleared from the
circulation [215,220,223]. The different components of the RONDEL™ delivery technology selfassemble into particles with a size ~70 nm, large enough to avoid renal filtration. Additionally, the
presence of a PEG corona should suffice to prevent rapid uptake by the MPS and dissociation in the
systemic circulation. In a recent report, Zuckerman et al. explained this seemingly paradoxical
clearance phenomenon by a more detailed look into the renal filtration barrier [220]. Apparently,
sub-100 nm polyplexes may extravasate across the glomerular fenestrated endothelium and
disassemble at the glomerular basement membrane (GBM), through interaction with the abundant
negatively charged proteoglycans (e.g. heparan sulfate). The released siRNA can subsequently be
filtered into the urinary space. Interestingly, the authors hypothesize that this clearance mechanism
maybe of general importance for cationic siRNA polyplexes that are small enough to cross the
glomerular endothelium [220].
Focusing on the intracellular delivery potential and with the aim to enable photo-induced
spatiotemporal control on the RNAi effect, Bøe et al. investigated the influence of PCI on the gene
silencing performance of siRNA CDPlexes, previously developed by the group of Mark Davis [224].
The authors chose to work with the unmodified CPDs, lacking the endosomolytic imidazol groups, to
maximize light-directed control over the intracellular siRNA delivery. With photochemical treatment
a gene knockdown of ~90% was obtained at the mRNA level in two different cell lines, while the RNAi
effect was nearly absent when PCI was not applied. Given that PCI treatment is under investigation
52
for in vivo anti-tumor applications [225], the authors foresee that combining unmodified siRNA
CDPlexes with PCI could be an attractive strategy for systemic light-directed tumor targeting [224].
CDPs have also been modified with OEI units to enhance pDNA delivery. Linear high Mw CDPs were
constructed by Srinivasachari and Reineke through Cu(I)–catalyzed azide-alkyne 1,3 dipolar
cycloaddition (“click reaction”), linking a diazo--cyclodextrin monomer with a series of dialkyneoligoethyleneimine monomers. The resulting CDPs achieved good reporter gene transfection
efficiency in HeLa cells [226]. In another report, OEI units (OEI600) were crosslinked with 2hydroxypropyl (,  or )-cyclodextrin following their activation with carbonyl diimidazol (CDI). The
resulting polymers demonstrated a lower toxicity and comparable or higher transfection efficiency in
SKOV-3 human ovarian carcinoma cells or SK-BR-3 human breast cancer cells compared to bPEI (25
kDa) [227,228]. An interesting feature for these polymers is their biodegradability by virtue of the
carbamate linkages, ensuring lower polymer toxicity [199]. The -hydroxypropyl-CD-OEI600 polymer
was further modified with an oligopeptide (MC-10) targeting the human epidermal growth factor
receptor 2 (HER2), that is frequently overexpressed on breast and ovarian cancer. Targeted polyplex
formulations with the interferon- (INF- gene could transfect SKOV-3 cells in vitro and in vivo,
resulting in an enhanced anti-tumor effect in SKOV-3 tumor-bearing mice, compared with the nontargeted formulation [229]. Likewise, a folate-targeted polyplex formulation, based on -CD-OEI
grafted with folic acid, could mediate gene expression in melanoma-bearing mice with a comparable
efficiency as adenoviral transduction [230].
5.2.2. Cyclodextrin modifications of pre-existing polymers
Instead of a polymeric CD backbone, several research groups have described the attachment of CD
moieties to pre-existing (cationic) polymers. This chemical engineering strategy allows to better
modulate their gene transfer potential and/or inherent toxicity. In addition, this again provides a
platform to include additional functionalities often required for in vivo use [199]. Among this class of
CDPs, PEI modifications obviously take a prominent position, with PEI being the gold-standard in non-
53
viral gene delivery. However, as mentioned before, the use of PEI as a transfection reagent is
severely hindered by its high cytotoxicity. It was the group of Mark Davis that first investigated the
effect of CD grafting on commercially available linear and branched PEI in depth [231]. Transgene
expression in PC3 cells was reduced as a function of increased CD attachment for both PEIs, while it
mitigated their in vitro toxicity. The inclusion of Ad-PEG in the formulation proved to be essential to
maintain NP colloidal stability at physiological salt concentration. Intriguingly, the PEGylated CDgrafted linear PEI achieved better in vitro gene transfection than linear PEI alone and could induce
liver gene expression following systemic administration in mice [231]. An elegant example of the
versatility of employing CDs in NP design was given by Zhang et al., who reported on the
supramolecular inclusion-driven self-assembly of NPs with a core-shell architecture [232]. The
nanoassemblies are composed of -CD grafted PEI and poly(-benzyl-L-aspartate) (PBLA), with PBLA
microdomains forming the hydrophobic core that is surrounded by a positively charged PEI
hydrophilic shell. The self-assembly into NPs is driven by the formation of inclusion complexes
between the hydrophobic benzyl groups and the pending CD moieties. The amphiphilic character of
the resulting NPs could be harnessed for the combined drug delivery of small hydrophobic drugs,
encapsulated in the hydrophobic center, and hydrophilic macromolecular drugs (e.g. proteins and
nucleic acids), complexed to the hydrophilic surface. This dual drug delivery concept was
demonstrated with dexamethasone and luciferase encoding pDNA in an osteoblast cell line [232].
Besides PEI, the grafting of CD units has also proved valuable to enhance the gene delivery potential
of other cationic polymers, such as poly-L-lysine (PLL) [233] and chitosan [234]. The group of Maria
Alonso proposed hybrid polysaccharide nanocarriers, containing both chitosan and cyclodextrin
building blocks, for gene delivery to the pulmonary epithelium. The particles were constructed by
TPP induced ionic gelation of native chitosan in the presence of anionic -cyclodextrin derivatives
and pDNA. The hybrid particles, prepared with low Mw chitosan (~10 kDa), succeeded in enhanced
gene expression in differentiated Calu-3 cells when compared to the conventional chitosan
formulation [234].
54
As a final example, the CD derivatization of neutral instead of cationic polymers has been put
forward by Kulkarni et al. in the context of siRNA and pDNA delivery [235,236]. They reported the
development of a degradable cationic polymer construct based on the self-assembly of
monosubstituted cationic -CD (amino--CD+) derivatives with a poly(vinyl alcohol) (PVA) polymer
backbone bearing PEG2000 and cholesteryl or adamantane modified grafts. The multivalent host-guest
interactions of the amino--CD+ with the pendant Chol/Ad-PVA-PEG enables electrostatic
complexation with nucleic acids into stabilized nanosized particles. The hydrophobic ‘guest’ anchors
are coupled to the PVA-PEG polymer via an acid-labile acetal linkage that can be hydrolyzed in
endolysosomal compartments following internalization by the target cells, mediating complex
disassembly. Polyplexes formed with siRNA or pDNA achieved comparable or higher biological
activity in vitro compared to 25 kDa bPEI or Lipofectamine 2000 while being far less toxic [235,236].
6. Other polysaccharides in nucleic acid nanotherapeutics
The polysaccharides discussed above are notable to many research groups involved in gene therapy.
Nonetheless, other less well-know polysaccharides may also be beneficial toward nucleic acid
delivery. This section aims to cite some remarkable contributions to gene transfer and/or gene
silencing made by -glucan, alginate, arabinogalactan, pullulan and pectin-based nanotherapeutics.
6.1. -Glucans
-Glucans are a heterogenous group of carbohydrates built from repeating D-glucose units, linked
together via -glycosidic bonds, and are primarily found in the cell wall of fungi [237]. As mentioned
before for other polysaccharides, many research groups also pursued cationic modification of the
original -glucan polysaccharide, e.g cellulose [238,239] and shizophyllan [240,241], to enable
electrostatic complexation of therapeutic nucleic acids. However, -1,3-glucans such as curdlan,
lentinan and shizophyllan, also have the intrinsic ability to form macromolecular complexes with
55
homopolynucleotides based on hydrogen bonding [242], a feature that was recently adopted by
Takedatsu et al. for the delivery of antisense oligonucleotides to macrophages [243]. These authors
employed schizophyllan (SP), that consists of a -1,3-D-glucan main chain with one -1,6-D-glucosyl
side chain every three glucose units (Figure 2). SP exists as a single chain at alkaline conditions and
adopts a triple helix structure at neutral pH. When neutralization of an alkaline solution of SP occurs
in the presence of homopolynucleotides, a hybrid triple helix can be formed in which two SP main
chain glucoses interact with one polynucleotide base [242,243]. Although SP can only hybridize with
homopolymeric sequences, the authors could form complexes of SP with functional AsON by
endowing the latter with a poly(dA) tail. -1,3-D-glucans are selectively bound by the pattern
recognition receptor (PRR) dectin-1 present on the surface of APCs and subsequently internalized by
phagocytosis. Complexation of AsON with SP would therefore enable their selective delivery to
macrophages, which are attractive target cells due to their role in promoting pathogenic
inflammation in several immune-related diseases such as atherosclerosis and inflammatory bowel
disease (IBD). Takedatsu and coworkers demonstrated the delivery of AsONs targeting the
macrophage-migration inhibitory factor (MIF) via SP complexation to CD11b+ macrophages, leading
to effective MIF suppression. Moreover, intraperitoneal administration of the antisense MIF/SP
complexes could attenuate intestinal inflammation in a murine dextran sodium sulphate (DSS)induced colitis model [243].
Besides -glucan polymer chains, also microparticulate -glucan shells have been exploited for
nucleic acid delivery. Aouadi et al. impressively demonstrated oral siRNA delivery to macrophages of
the gut-associated lymphatic tissue (GALT) to be able to downregulate pathogenic inflammatory
responses. Phagocytosis was mediated by formulating the siRNA in micrometer-sized -1,3-D-glucan
shells (GeRPs) obtained by solvent extraction of baker’s yeast. The GeRP core consisted of yeast
tRNA, Endo-Porter [244] (EP), PEI and siRNA in a layer-by-layer format to assist with cytoplasmic
siRNA release triggered by phagosomal acidic pH. Orally delivered GeRPs were able to systemically
silence TNF- in macrophages recovered from several MPS related tissues, since GALT associated
56
macrophages may disseminate from the gut. Moreover, silencing of the mitogen-activated protein
kinase kinase kinase kinase 4 (Map4k4) via oral gavage of GeRPs could protect mice from a lethal
dosis of lipopolysaccharide (LPS) owing to inhibited production of TNF- and IL-1.
To simplify this complex five-component GeRP formulation, in a follow-up study only the EP
amphipathic peptide was used to complex the siRNA in the GeRP core [245]. EP was identified as the
critical component in this formulation, enabling siRNA entrapment based on electrostatic interaction
and potentiating phagosomal destabilization. Moreover, omitting PEI from the original GeRP
formulation is believed to be beneficial to constrain in vitro and in vivo toxicity [245]. It has to be
noted that the specific interaction of (particulate) -1,3-glucans with membrane-bound and cytosolic
pattern recognition receptors (PRRs), such as dectin-1 and NLPR-3 inflammasome, may trigger a proinflammatory response [246,247]. Although this may promote their usefulness as adjuvants in
(genetic) vaccination strategies [237,248], in part it also contradicts their application as delivery
agents of anti-inflammatory therapeutics. Nonetheless, in the above-mentioned reports no effects
were observed in control experiments with neither uncomplexed shizophyllan [243] nor empty -1,3D-glucan shells [249].
6.2. Alginate
Alginate is an anionic naturally occurring polysaccharide that can be obtained from brown algae and
bacteria such as Azotobacter and Pseudomonas [250]. Alginate is defined as a linear block copolymer
composed of regions with consecutive -1,4-D-mannuronic acid residues (M-blocks), -L-guluronic
acid residues (G-blocks) and alternating M and G residues (MG-blocks) (Figure 2). The copolymer
composition and its Mw may vary significantly with source and species. Because of its
biocompatibility, low toxicity and muco-adhesiveness, also alginate is an attractive polymer for
biomedical applications [14]. In addition, in analogy with chitosan, alginate can form 3D hydrophilic
networks (hydrogels) via ionic gelation induced by divalent cations (e.g. Ca2+), an attractive feature
toward controlled drug release applications. Gel formation predominantly involves the association
57
between G-blocks on adjacent polymer chains in the presence of divalent cations as ionic crosslinker.
Alginate-based macro-and microscopic hydrogels have mainly been investigated for controlled small
molecule and protein delivery in the context of wound healing and tissue regeneration [250]. A
limited number of reports also describes alginate-based nanostructures as nucleic acid carrier.
In accordance with reports on e.g. HA and DS, alginate has also been exploited in pursuit of
alleviating PEI mediated cytotoxicity [251]. Patnaik et al. showed successful gene and siRNA
transfection in various cell lines in vitro with PEI-alginate nanocomposites composed of different Mw
PEI (25 kDa or 750 kDa) [252,253]. For transfection purposes, the obtained PEI-alginate
nanocomposites were complexed with pDNA or siRNA. To PEGylate the complexes, PEG4000bis(phosphate) was incorporated into the amalgam wherein 10% of the PEI amine groups ionically
interact with the anionic phosphate end-groups on the homobifunctional PEG. At an optimal alginic
acid content, the nanocomposites outperformed native PEI while being almost non-toxic [252,253]. It
should however be questioned in future work to what extent these electrostatic complexes will hold
together in biological media before proceeding toward in vivo studies. Jiang et al. endowed
preformed PEI/DNA polyplexes with an alginate coat and demonstrated that this enhanced reporter
gene expression in vivo in comparison to the uncoated complexes [254,255]. More recently, He et al.
also reported on PEI-alginate conjugates by grafting 2 kDa PEI to an alginate backbone with the
corresponding aim to optimize the balance between efficiency and toxicity [256].
Making use of alginate’s ability to form a crosslinked gel-like structure, Shardool and Mansoor
encapsulated pDNA encoding the anti-inflammatory murine IL-10 into alginate NPs with a 60%
loading efficiency. The negatively charged surface of the particles was electrostatically modified with
a tuftsin targeting peptide sequence (TKPR) via a hexameric L-arginine motif preceeding a tetrameric
L-glycine spacer. Tuftsin is known to stimulate receptor-mediated phagocytic uptake and serves to
direct the DNA loaded alginate NPs to macrophages. Transfecting IL-10 in murine J774A.1
macrophages with the tuftsin modified particles could significantly block subsequent LPS stimulated
TNF- expression.
58
6.3. Arabinogalactan, pullulan and pectin
In addition to the dextran-spermine conjugates [61], described under 2.3.2., the group of Y.
Barenholz also reported on the grafting of spermine to an arabinogalactan backbone [58,257].
Opposite to dextran-spermine conjugates, the arabinogalactan-spermine derivatives however failed
to induce substantial transgene expression. The authors attempt to explain this disparity by pointing
at the difference in polysaccharide structure. While dextran is a linear -glucan, arabinogalactan is a
highly branched natural polysaccharide, composed mainly of galactose and arabinose residues.[14]
The branching structure may be responsible for the low level of spermine conjugation observed in
arabinogalactan compared to dextran. Importantly, this will also entail a decreased percentage of
secondary amines in arabinogalactan, interfering with the polymers ability to induce a proton sponge
effect following endocytic uptake [257].
Pullulan is known as a neutral and non-toxic exopolysaccharide of fungal origin composed of -1,6linked maltotriose units. It is produced from starch by the fungus Aureobasidum pullulans. Pullulan
has a unique linkage pattern that contributes to its adhesiveness and ability to form fibers and
biodegradable films. These characteristic physicochemical properties explain its use in food and
cosmetic industry as well as various biomedical applications [258]. Because of its non-ionic nature,
several research groups turned to cationic modifications of pullulan to enable the formation of
electrostatic complexes with pDNA for in vitro and in vivo gene delivery [259-261]. Interestingly it has
been shown that pullulan, equal to arabinogalactan, has intrinsic liver targeting properties, resulting
from its interaction with the asialoglycoprotein receptor present on liver parenchymal hepatocytes
[262]. However, it has to be noted that extensive chemical modification of the native polysaccharide
may greatly influence its affinity for the liver [261].
Another polysaccharide rarely reported for gene delivery is pectin, an important constituent of the
cell wall of fruits and vegetables, best known for its use as gelling and thickening agent. It is a very
complex and heterogenous polysaccharide, both in polymer length as in chemical composition.
Pectin extracts predominantly consist of homogalacturonan regions (HG) composed of repeating α-
59
1,4-D-galacturonic acid units. HG regions, which are often designated as ‘smooth regions’ are
attached to so-called ‘hairy regions’, that are built from alternating 1,2-linked -D-galactose and -Lrhamnose residues. The hairy regions are highly branched with neutral arabinan, galactan and
arabinogalactan side chains. Although occurring less frequently, pectin extracts can also contain
other highly branched structures, adding to its complexity [263]. It is hypothesized that the
galactose-rich side chains of pectin stimulate binding to galactose-binding lectins (galectins),
expressed on the cell surface of the cancer cells, making pectin a potential carrier polymer in cancer
therapy [17,264]. However, although the galacturonic acid groups in pectin are partially methylated,
unmodified galacturonic acid moieties confer its anionic nature which is not beneficial for nucleic
acid complexation. Therefore, Katav and coworkers modified citrus pectin with primary, tertiary or
quaternary amine groups and formulated polyplexes with GFP encoding pDNA. Reporter gene
expression in HEK 293 cells was highest for the quaternized pectin although further improvements on
endosomal escape and intracellular disassembly may be required [264]. Alternatively, ionotropic
gelation of pectin in the presence of divalent cations (Ca2+, Mg2+, Mn2+) and pDNA may lead to the
formation of DNA loaded micro-and nanogel particles thereby avoiding potential cytotoxicity induced
by cationic modifications, albeit with limited success in transfection efficiency [265]. Very recently,
Zhou et al. reported on the derivatization of hyperbranched amylopectin with various oligoamines.
The newly constructed amylopectin derivatives exhibited lowered erythrocyte lysis and cytotoxicity
when compared with branched PEI, but did not outperform the latter polymer with regard to
transfection efficiency [266].
7. Conclusions and future perspectives
Polysaccharides are promising candidates in nanocarrier development for nucleic acid delivery, due
to several intrinsic advantages (table 1). Most natural polysaccharides discussed in this review share
the important benefit of biocompatibility and biodegradability. However, to be successful in nucleic
acid delivery, a first requirement for polysaccharide-based nanocarriers is to ensure efficient nucleic
60
acid encapsulation and protection. Because of its polycationic nature, chitosan is frequently applied
for the assembly of polyelectrolyte complexes (PECs) with negatively charged nucleic acids.
Conversely, other polysaccharides require complexation with polycations or chemical modification
with oligo-or polyamines to enable nucleic acid complexation. Fortunately, polysaccharides possess
many functional groups that allow straightforward derivatization. Although reworking the structure
of a native polysaccharide can both be beneficial to nucleic acid complexation and intracellular
nucleic acid delivery, the inherent biocompatibility can no longer be guaranteed. It is therefore
imperative to perform an in depth evaluation of the potential toxicity of modified polysaccharides
and resulting NPs. In view of that, even though the potential of nanocarriers to improve delivery of
nucleic acid therapeutics cannot be denied, their transition from bench to bedside will strongly rely
on the emerging field of nanotoxicology to assess the parameters governing the toxic effects inflicted
by nanomaterials. Polysaccharides and their derivatives have been used to optimize nucleic acid
nanocarrier design toward in vivo applications. Prolonging the blood circulation time of NPs, by
endowing them with stealth properties, could promote passive tumor targeting via the enhanced
permeability and retention (EPR) effect. Many polysaccharides also possess excellent bioadhesive
properties, justifying their use in mucosal drug delivery strategies (e.g mucosal DNA vaccination). In
addition, several cell types express carbohydrate-binding receptors recognizing specific
oligosaccharide motifs, enabling active targeting at the cellular level depending on the type of
polysaccharide used. However, to clearly state a therapeutic benefit, additional in vivo data in
validated and representative animal disease models are urgently required. As reviewed above, many
types of polysaccharides with variable chemical composition, molecular structure and polymer
length, have been employed in nucleic acid nanocarrier design. In concert with a range of different
methodologies available for NP production, this entails a plethora of existing polysaccharide-based
nanotherapeutics, albeit often with unpredictable biological and physicochemical properties. Besides
thorough physicochemical characterization, future research should therefore aim for a clearer view
on the (un-)specific interactions at the nano-bio interface, both on the extra-and intracellular level.
61
More detailed insight in the underlying mechanisms of the nucleic acid delivery process will strongly
contribute to the design of tomorrow’s safe and effective nucleic acid delivery systems.
62
Table 1. Predominant beneficial properties contributing to improved nucleic acid delivery typically
ascribed to the most commonly used polysaccharides in biomedical and pharmaceutical applications
Polysaccharide advantage
dextran
cyclodextrin
chitosan
hyaluronan
Low cytotoxicity
Non-immunogenic§
Nucleic acid complexation
Ionotropic gelation
Ease of chemical modification
Membrane permeation enhancer
muco-adhesive
‘stealth’ properties*
Receptor targeting
§
also dependent on other parameters, such as molecular weight (e.g. hyaluronic acid) and nano-
architecture (e.g. dextran and chitosan)
*stealth properties include steric shielding, decreased unspecific binding and reduced complement
activation
Acknowledgments
The Research Foundation Flanders (FWO) and the Ghent University Special Research Fund are
gratefully acknowledged for their financial support.
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Graphical abstract