In Vivo Gene Delivery by Nonviral Vectors: Overcoming Hurdles?

Abstract

The promise of cancer gene therapeutics is hampered by difficulties in the in vivo delivery to the targeted tumor cells, and systemic delivery remains to be the biggest challenge to be overcome. Here, we concentrate on systemic in vivo gene delivery for cancer therapy using nonviral vectors. In this review, we summarize the existing delivery barriers together with the requirements and strategies to overcome these problems. We will also introduce the current progress in the design of nonviral vectors, and briefly discuss their safety issues.

Introduction

Gene therapy holds the significance of correcting genetic defects, and there are many nucleic acid-based therapeutic strategies that can be used for gene therapy against cancer, including antisense and RNA interference (RNAi) mechanisms. Antisense oligonucleotides are typically 15–30 nucleotides long and block production of the disease-causing protein after complementarily hybridizing to their target messenger RNA (mRNA) and degrading the mRNA by activating RNaseH. RNAi is a separate process in which a specific mRNA is targeted for degradation in order to inhibit the synthesis of its encoded protein. Two types of small RNA molecules—microRNA (miRNA) and small interfering RNA (siRNA)—are central to the RNAi function. After delivery into the cytoplasm, the antisense strand of RNAi molecules recruits the corresponding mRNA in a sequence-specific manner to the RNA-induced silencing complex, which is followed by cleavage of the target mRNA, resulting in gene silencing. Both antisense oligonucleotides and RNAi-based therapy target mRNA to inhibit transcription of an overexpressing endogenous gene or a cancer-causing oncogene,¹ resulting in selectively inhibiting the expression of an unwanted protein (downregulation or loss of function). Plasmid DNA (pDNA) is also widely used to introduce a normal wild-type transgene into specific cells of the host where the endogenous gene is underexpressing,¹ resulting in expression of a deficient protein (upregulation or gain of function).

Currently, three different kinds of gene delivery systems have been explored: modified naked siRNA, viral vectors, and nonviral vectors. Modified naked siRNAs have increased nuclease stability and gene silencing efficiency, as well as reduced immune responses and off-target effects, when compared to unmodified RNAs.² Viral vectors have high gene transfection efficiency, but their residual viral elements can cause insertional mutagenesis and immunological problems. Nonviral vectors are constructed with biocompatible materials using innovative fabrication approaches, so that they can safely transport gene cargo in vivo, but their transfection efficiencies are not as high as viral vectors.

Therapeutic nucleic acids are potent in correcting genetic defects in cell and molecular levels, but in a real biological environment, their lack of serum stability and rapid clearance greatly compromise their in vivo delivery, which retards gene therapy application in clinical setting. In this review, we will summarize the existing delivery barriers, highlight the strategies to overcome these hurdles, and introduce the current progress of nonviral vectors constructed for cancer gene therapy.

Challenges of Gene Delivery

Off-target effect

siRNAs are capable of reducing expression of nontarget genes due to interaction of the siRNA guide strand with a partially complementary site on an “off-target” mRNA.³ Careful selection of siRNA sequences to avoid off-target effects is an important issue, and can be minimized by avoiding certain sequence motifs. The design and selection of a specific siRNA may involve the consideration of internal repeated sequence, GC content, siRNA length, specific base preference, secondary structure, etc.^4,5

Immune stimulation

Innate immune activation via RNA represents a significant undesirable side effect due to the toxicities associated with excessive cytokine release⁶ and inflammatory syndromes after systemic administration. The inflammatory response is mediated by toll-like receptors (TLRs) TLR3, TLR7, TLR8, and TLR9, which are located in endosome compartments, and recognize unmethylated CpG DNA as well as various moieties in double-stranded RNA or their degraded products.⁷ A number of parameters affect how delivery vehicles potentiate immune stimulation. These include the chemical composition of the delivery vehicles, physical properties such as the size and charge of formulated materials, pharmacokinetics and biodistribution of formulated nucleic acids, and routes of administration.⁶ TLR-mediated recognition and concomitant immune stimulation can be inhibited by chemical modifications, such as introduction of 2′-O-methyl (2′OMe)-modified nucleotides.⁸

Delivery

The most important—and most difficult—challenge in gene therapy is the issue of delivery. Not only must the therapy evade the reticuloendothelial system (RES) as it circulates after systemic administration, but it must also cross several barriers before it arrives in the cytoplasm or nucleus of its target cells.

Serum inactivation and enzyme degradation. Stabilization of RNA is regarded as a prerequisite for in vivo therapeutic applications, however, naked RNAs are rapidly degraded by nuclease in serum. Nuclease resistance can be enhanced by chemical modification of RNAs. Because the 2′-hydroxyl group of the ribose ring is not necessary for potent gene silencing by siRNAs, the most widely used RNA modifications are on sugar moieties. Common modifications are 2′-fluoro, 2′-O-methyl, and 2′-amine conjugations.⁹ The locked nucleic acid conformation has a 2′-O, 4′-C methylene bridge in the sugar ring. Such locked nucleic acid molecules have the desirable features of increased nuclease stability and silencing potency, as well as reduced off-target effects and immune responses.⁷ Backbone modification by changing the internucleotide phosphate linkage, such as phosphorothioate (P = S) and boranophosphonate (P = B) modifications, can be placed in the RNA duplex at any desired position to enhance nuclease stability.⁹

Though modified RNAs are favorable for in vivo delivery when compared to naked RNAs, they lack specific tumor targeting and would be quickly excreted by the kidney upon systemic administration, so a large amount of modified RNAs are needed to attain desired therapeutic effects.¹⁰ Generally, modified RNAs are delivered by local injection, but the local route is only accessible for certain cancers, such as skin cancer or head and neck cancer. Thus, most efforts have been concentrated on developing safe and effective nanoparticles for systemic gene delivery.

RES recognition. In order to condense negatively charged nucleic acids into delivery vehicles, most nanoparticles contain polycations such as cationic polymers or lipids. The positive charge aids cellular uptake but also promotes nonspecific interactions with nontarget cells and extracellular components such as serum proteins and extracellular matrix.¹¹ Binding of plasma proteins is the primary mechanism for the RES to recognize circulating nanoparticles,¹² causing the major injected dose to go to RES organs, such as the liver, spleen, and bone marrow, immediately after intravenous injection, with little accumulation in tumors.

The most common way to decrease nonspecific interactions is to shield the nanoparticle surface with hydrophilic, uncharged polymers such as polyethylene glycol (PEG). Surface PEG coating sterically hinders the interaction and binding of blood components with the nanoparticle surface and prevents opsonization and recognition by phagocytes of the RES,¹³ resulting in prolonged nanoparticle circulation in the blood. The nanoparticle circulation half-life may vary by changes in PEG chain length, PEG density, surface coating, particle size, and surface charge of the underlying nanostructure.¹⁴ PEGylated (aka stealth) nanoparticles with a reduced RES uptake and a prolonged circulation half-life are a prerequisite for enhanced tumor targeting.¹²

The EPR effect. Transport of macromolecules across the tumor endothelium is more efficient than that of normal endothelium because of its leaky and discontinuous vascular structures (permeation) with poor lymphatic drainage (retention), which is referred to as “enhanced permeability and retention (EPR) effect”.¹⁵ In other words, tumor endothelium allows the penetration of macromolecules and most nanoparticles. The EPR effect can be enhanced by PEGylation because the amount of blood that circulates through the tumor is usually far less than that of the RES organs, and only those nanoparticles that are not rapidly cleared from the circulation will have a chance to encounter the leaky tumor vasculature.¹²

Not all human tumors are equally leaky; for tumors with less leaky vasculatures, nanoparticles with small size (less than 30 nm) are desirable,¹³ but for most tumors, nanoparticles with a mean size around 100 nm are attractive for tumor targeting. Particles that are 100 nm in size also allow easier surface modification with PEG arranged in a brush mode, a conformation that efficiently prevents serum opsonization. Particles larger than 400 nm can not easily enter the capillary gaps in the tumor vasculature, whereas particles smaller than 70 nm are able to access the parenchymal cells in the liver (i.e., hepatocytes) after crossing the liver blood vessels (known as the sinusoidal space), since the sinusoidal vessels contain fenestrae that have an average diameter of 100 nm.¹¹ Small particles are also prone to renal excretion through the glomeruli in the kidney, and their large surface curvature presents difficulties for PEG shielding.

Entrance into cells: the PEG dilemma. PEGylation may protect nanoparticles from protein agglomeration and macrophage capture, controlling their biodistribution and tumor accumulation, but the PEG coating prevents the formation of essential nonbilayer intermediates¹⁶ and inhibits fusion with the cell and/or the endosomal membrane, thus reducing the potential of nanoparticle cellular uptake and cargo release from endosomes, decreasing silencing or transfection efficiency.

To solve this PEG dilemma, labile bonds can be introduced in the PEG chain, so that nanoparticles become unPEGylated upon reaching the target cells, leading to increased rates of membrane destabilization, transport of the loaded cargo inside the cells, and release from the endosome. pH-sensitive lipids composed of PEG, an acid-labile linker, and a hydrophobic tail can be used to construct fusogenic liposomes,¹⁷ or one can conjugate PEG and polycations with acid-sensitive linkers followed by compacting nucleic acids into pH-triggered deshielding lipoplexes/polyplexes.¹⁸ Enzyme-sensitive linkers can also be applied to conjugate PEG and lipids, followed by entrapping nucleic acids in liposome-like vehicles, so that the PEG moiety can be cleaved off in tumor sites where the specific enzyme is widespread.^19,20,21,22

Besides the external stimuli-triggered PEG shedding, conjugates of lipids and hydrophilic polymers are generally able to diffuse off membranes depending on the strength of the anchorage or the anchor chain composition.²² When particles (liposomes or lipoplexes) are coated with sheddable stealth PEG-lipid,²² the spontaneous shedding of the PEG-lipid from particles (de-PEGylation)^12,23 will continuously happen when particles circulate in blood, which eventually exposes the shielded cationic lipids and allows membrane fusion for nanoparticle uptake and endosome release.¹² In this case, the shedding rate is a crucial parameter that has to be addressed when designing sheddable PEG coating. If shedding occurs too quickly, the unprotected carrier will be rapidly cleared from circulation by the RES. In other words, RES competes with tumor for the uptake of nanoparticles.¹² However, if shedding is incomplete or occurs too slowly, the therapeutic efficacy of the loaded cargo might be compromised.

Targeting ligands are also frequently modified on the nanoparticle surface for enhanced cellular uptake by receptor-mediated endocytosis. Commonly used targeting ligands include aptamers,²⁴ cell penetrating peptides,²⁵ antibodies,²⁶ peptides or proteins,²⁷ and small molecule ligands.²⁸ Huang et al.²⁹ conjugated the LRP1 ligand angiopep-2 to a PAMAM dendrimer using bifunctional PEG, and complexed it with plasmid DNA (pORF-TRAIL) to treat brain glioma. Significant apoptosis induction and prolonged survival time after systemic administration indicated that the angiopep-2 peptide could be exploited as a specific ligand to cross the blood–brain barrier and target glial tumors. Sonsoles Diez et al.³⁰ constructed an asialoglycoprotein receptor-targeted lipopolymeric vector using the asialofetuin ligand for IL-12 gene transfer in hepatocellular carcinoma in vivo.

Endosome escape. The delivery of nonviral gene vehicles almost invariably involves endocytosis, and escaping from endosomes before they traffic into lysosomes is an essential step for nanovectors to avoid enzymatic degradation.

Cationic lipid complexes can bind to anionic lipids on the endosome membrane and form neutral ion pairs. These ion-pairs destabilize the endosome membrane and promote de-assembly of the lipoplex through the formation of the inverted hexagonal (H_II) phase,³¹ and finally release nucleic acids to cytoplasm.

Some protonable groups that are charged at acidic pH but less charged at neutral pH could be an alternative choice to designing ionizable cationic lipids that condense nucleic acids and promote endosome release.^11,32,33 The rational design lies in the pKa of the ionizable cationic lipid and the abilities of these lipids to induce the hexagonal H_II phase structure with anionic lipids of the endosomal membrane when protonated in the acidic endosome.

Acid-responsive mechanisms have been widely used to design delivery carriers in order to promote endosome release. Polymers and peptides with high buffer capacity between pH 7.2 and 5.0, such as polyethylenimine (PEI),³⁴ or peptides containing the cationic amino group lysine, arginine, and imidazole group histidine,³⁵ could buffer the endosome. This would cause more protons to enter into the endosomes, followed by chloride ions, leading to increased osmotic pressure and endosome rupture, releasing payloads into cytoplasm. This process is called the “proton sponge effect”.

Stimuli other than pH have been used to destabilize endosome membranes as well. Lipid or polymer derivatives which are sensitive to sulfhydryl reduction^36,37,38 and enzymatic cleavage^21,39 have been used to construct nonviral gene vectors. Upon exposure to the intracellular reducing agents or selective enzymes at the target sites, the vectors become destabilized and fuse with the endosome membrane, and finally release the entrapped cargo. Some fusogenic peptides can also be combined with nanoparticles to induce membrane fusion in endosomes through their structural changes in acidic conditions as compared to physiological pH.^40,41

Li and Huang⁴² developed a lipid-calcium-phosphate nanoparticle (LCP) for efficient siRNA delivery. The LCP entraps siRNA in a biodegradable nanosized calcium phosphate precipitate (CaP) core. The LCP de-assembles in endosomes due to the CaP core dissolving in acidic conditions, which increases the osmotic pressure, causing the endosome to swell and burst to release the entrapped siRNA.

Nuclear entry. The final target destination of antisense oligonucleotides, siRNA/miRNA, and mRNA, is the cytoplasm, whereas pDNA must be transported into the nucleus for gene expression. Nuclear transport generally occurs through nuclear pore complexes, however, nucleic acid condensates are impermeable through nuclear pore complexes due to their large size.⁴³ In dividing cells, the nuclear envelope disassembles during mitosis; pDNA transfection can only occur at this stage of the cell cycle due to elimination of the permeability barrier. The amount of DNA that reaches the nucleus is made lower due to the cytoplasmic nuclease that can degrade DNA, such that the majority of DNA that enters the cytoplasm never arrives in the nucleus. For nondividing cells, the mechanisms of DNA nuclear transport are of critical importance.

To facilitate nuclear targeting, many nuclear localization signal (NLS) peptides have been developed to allow DNA nuclear entry through nuclear pore complexes by active transport. NLSs are short clusters of amino acids that can bind to cytoplasmic receptors known as importins. NLS peptides can bind to DNA either through noncovalent electrostatic interaction or by covalent attachment. The most well-known and popularly used NLS is from the large tumor antigen of simian virus 40 (SV40).⁴⁴ Some DNA sequences themselves have nuclear import activity based on their ability to bind to cell-specific transcription factors, such as the SMGA promoter and flk-1 promoter.⁴⁵

In summary, in order to elicit superior in vivo therapeutic response to correct genetic defects, the nonviral vector must be able to tightly condense and protect nucleic acids to avoid enzymatic degradation, accumulate at tumor sites with little RES uptake, target specific cells, disrupt the endosomal membrane and release the therapeutic cargo to cytoplasm, and translocate the DNA cargo to the nucleus.⁴³ (Figure 1)

Figure 1.

Representative scheme of in vivo gene delivery barriers. EPR, enhanced permeability and retention; mRNA, messenger RNA; PEG, polyethylene glycol; RES, reticuloendothelial system; RISC, RNA-induced silencing complex; siRNA, small interfering RNA.

Nonviral Vector Design For in Vivo Gene Delivery

Lipoplex/polyplex

Nucleic acids are easily complexed with cationic lipids (e.g., DOTAP), cationic polymers (e.g., PEI), biodegradable cationic polysaccharides (e.g., chitosan), and cationic polypeptides (e.g., polylysine, protamine), via electrostatic interactions. By changing the charge ratio associated with nucleic acids and condensing agents, the transfection efficiency of condensates can be optimized.¹¹ Cationic polymers and lipids have shown superior gene transfection efficiency, but they have dose-dependent toxicity upon systemic administration. Among cationic polymers, the polymer chain length and the presence or absence of hydroxyl groups play a role in polyplex size and charge in addition to transfection efficiency and toxicity. The high gene transfection usually correlates with the high cytotoxicity.⁴⁶ Grandinetti et al.⁴⁷ reported that direct interaction between (PEI-pDNA) polyplexes and mitochondria during PEI transfection causes impaired mitochondrial function through membrane depolarization and could be the reason for high cytotoxicity in the PEI-based vehicles. This toxicity can be ameliorated by conjugation with biocompatible, hydrophilic polymers such as PEG,^48,49 and some pH and enzyme-sensitive linkers are widely used. For example, Walker et al. designed pH-triggered deshielding polyplexes to enhance endosome release, and Yin et al. entrapped RNA in poly (β-amine esters) complex nanoparticles that are degradable in the reductive environment due to the cleavage of disulfide bonds.^18,50 Furthermore, some synthetic polymers have been optimized using combinatorial chemistry and library screening in order to mimic viral functional delivery domains, such as surface ligands for cell entry, evasion from endolysosomal compartment, and entry into the cytoplasm of target cells. Many defined, precise sequences have been found to perform potent pDNA and siRNA delivery by using solid-phase synthesis.⁵¹ Since different topologies of the defined synthetic polymer structures can influence the complexation and biological properties of transfection agents, Schaffert et al.⁵² explored the gene delivery efficiency and cytotoxicity of linear polycations with modifications in different areas and found that these changes could cause robust differences in biological function.

Lipopolyplex

In order to increase serum stability, avoid RES uptake, and increase tumor accumulation, Li et al.⁵³ designed a liposome-polycation-DNA nanoparticle (LPD) for siRNA delivery. siRNAs were mixed with calf thymus DNA before being condensed with protamine. Protamine is a natural arginine-rich cationic polypeptide which is biocompatible, biodegradable, and nontoxic, so it is more desirable than synthetic polymers. The condensed particles were wrapped by cationic liposomes, followed by grafting PEG-lipid conjugate (DSPE-PEG) with or without a targeting ligand using a postinsertion method. The LPD can deliver a significant fraction of injected siRNA into the tumor after intravenous injection with little RES uptake. siRNA formulated in the targeted LPD completely silenced oncogenes in tumors, induced tumor cell apoptosis, and achieved superior tumor growth inhibition.^54,55 By adjusting the siRNA to protamine ratio, a positively charged LPD core can also be attained. LPD-II was formed by wrapping an anionic liposome composed of DOPA (dioleoylphosphatydic acid) around this LPD core, followed by grafting DSPE-PEG.⁵⁶ Based on the LPD formulation, Chono et al.⁵⁷ developed a liposome-polycation-hyaluronic acid (or heparin) nanoparticle (LPH). A remarkable advantage for LPH is that it showed very little immunotoxicity in a wide dose range compared to LPD. Chen et al.⁵⁸ used the scFv ligand targeted LPH to effectively deliver siRNAs (c-Myc/MDM2/VEGF) and miRNA-34a to a B16F10 lung metastasis model. Figure 2 shows the structure and preparation scheme of LPD (LPD-II)/LPH nanoparticles.

Figure 2.

The structure and preparation scheme of LPD (LPD-II)/LPH nanoparticles. LPD, liposome-polycation-DNA; LPH, liposome-polycation-hyaluronic acid (or heparin); pDNA, plasmid DNA; PEG, polyethylene glycol; siRNA, small interfering RNA.

mRNA can be loaded in lipopolyplexes (termed LPRs, lipid-polymer-RNA) using the similar method of LPDs. Perche et al.⁵⁹ reported that MART-1 mRNA lipopolyplexes with mannosylated liposomes (Man11-LPR100) targeting dendritic cells can be used as an efficient system for anti-B16F10 melanoma mRNA-based vaccines. A great inhibition of B16F10 melanoma growth was obtained after mice were intravenously immunized with MART-1 Man11-LPR100, indicating that tumor antigen mRNA-loaded Man11-LPR100 is an efficient system to induce an anticancer immune response. Although mRNA serves as a potential therapy in various medical indications, like other nucleic acid molecules, its strong immunogenicity and limited stability hamper clinical applications. To solve these limitations, Kormann et al.⁶⁰ designed a combination of nucleotide modifications. They found that replacement of only 25% of uridine and cytidine with 2-thiouridine and 5-methyl-cytidine synergistically abrogated mRNA interaction with TLR3, TLR7, TLR8, and retinoid-inducible gene I, thus substantially decreasing activation of the innate immune system. High expression of therapeutic proteins were detected in mice after intramuscular administration of double-modified mRNA, as well as in a congenital surfactant protein B (SP-B) deficiency disease model after two intratracheal doses of modified SP-B mRNA. Furthermore, entrapping modified mRNA in LPD nanoparticle has shown superior therapeutic effects in xenograft tumor models (unpublished data in Huang lab).

Although LPD nanoparticles are successful in the systemic delivery of siRNA, the delivered siRNA does not completely dissociate from the nanoparticles, such that most of the encapsulated siRNA is not bioavailable. In order to solve the problem regarding the inefficient release of siRNA from LPD, Li et al.⁴² developed a LCP nanoparticle. The LPD's DNA-protamine complex core was replaced by a biodegradable nanosized CaP precipitate prepared by water-in-oil microemulsions, and siRNA was entrapped in the CaP core. The rationale for the LCP design is that the CaP precipitate in the LCP core would dissolve and de-assemble at low pH in the endosome, increase the osmotic pressure, and cause endosome swelling and bursting to release the entrapped siRNA. The LCP can be further optimized by changing the precipitate core and the coating lipids.

Harashima et al. developed a liposomal gene carrier known as multi-functional nano device (MEND) for systemic gene delivery. Similar to LPD, MEND consists of a nucleotide core condensed with polycations and covered with lipid membranes. The surface of the lipid envelope can be modified with various functional devices, such as PEG for prolonged circulation, specific ligands for targeting, or fusogenic peptides for endosomal escape.⁶¹ In order to avoid the PEG dilemma, an enzyme-cleavable PEG system, PEG-peptide-DOPE conjugate, was used to modify MEND.²¹ In this strategy, the PEG is removed from MEND via the cleavage by a matrix metalloproteinase, which is specifically expressed in tumor tissues.

Harashima et al.^40,41 further modified a fusogenic peptide GALA (WEAALAEALAEALAEHLAEALAEALEALAA) on MEND for the sake of promoting endosome release. The 30 amino acid GALA contains a glutamic acid-alanine-leucine-alanine sequence that is repeated four times. Since the carboxyl groups of glutamic acid are negatively charged at physiological pH, electric repulsion between these groups forces GALA to be a random coil structure. In contrast, at acidic endosomal pH, protonation of the carboxyl group side chains of the glutamic acids dissipates electric repulsion, so the GALA structure changes into an α-helix, a structure that tends to induce membrane fusion. To avoid the recognition by biomacromolecules in vivo, a shorter version of GALA (shGALA) was developed, which was masked by the PEG layer of the MEND. The shGALA-modified MEND showed significant gene silencing in the tumor and inhibition of tumor growth.

Aptamer-siRNA chimeras

Aptamers are three-dimensional single-chain nucleic acid molecules that bind to a specific target molecule with high affinity and specificity.⁶² They are selected from a combinatorial library of randomized sequences through repeated rounds of selection, known as “systematic evolution of ligands by exponential enrichment” (SELEX),⁶³ and the target molecules can be small molecules, nucleic acids, proteins, carbohydrates, whole cells, and even organisms. Aptamers have high diversity and those possessing very high affinities to the target molecules can be isolated. In addition, they are relatively stable in storage and elicit little immunogenicity in therapeutic applications.

Most aptamer-siRNA targeted chimeras for cancer therapy are against prostate-specific membrane antigen (PSMA), a cell-surface receptor overexpressed in prostate cancer cells and tumor vascular endothelium. McNamara et al.⁶⁴ developed aptamer-siRNA chimeric RNAs. The aptamer portion of the chimeras mediated binding to PSMA, whereas the siRNA portion targeted the expression of tumor survival gene Plk1 and Bcl-2. The chimera effectively delivered siRNAs to LNCaP prostate cancer cells expressing PSMA, and triggered apoptosis and cell death both in cultured cells and in a prostrate tumor xenograft model.⁶⁴ In order to enhance silencing activity and specificity of siRNA, Dassie et al.⁶⁵ incorporated modifications, including adding 2-nucleotide 3′-overhangs and optimizing the thermodynamic profile and structure of the duplex, which enabled more efficient processing of the siRNA guide strand and reduced the effective concentration of siRNA portion. The optimized chimeras resulted in pronounced regression of PSMA-expressing tumors after systemic administration. Antitumor activity was further enhanced by grafting a 20 kDa PEG moiety to the 5′-terminus of the RNA strand, which increased the chimeras' circulating half-life and bioavailability without affecting chimera targeting and silencing. The aptamer portion of the chimeras can also be truncated from 71 to 39 nucleotides without loss of function, so that large-scale chemical synthesis can be facilitated. In principle, the aptamer-siRNA chimera approach can be applied to target reagents to many different cell types, provided that a cell-specific receptor exists and that an aptamer against the receptor can be selected.⁶⁵

Local Gene Delivery

Local delivery by external stimulations can avoid or delay RES uptake, reduce systemic toxicity, provide organ specificity, and help the delivery system reach the target cells. These environment-sensitive nanoparticles have been designed to release their contents based on the environmental changes leading to controlled drug release, such as temperature, light, ultrasound response, or magnetic stimulus.

Bubble liposomes by ultrasound exposure

Ultrasound technology provides an easy, safe, minimally invasive, and tissue-specific method for gene delivery into tumors.⁶⁶ The bubble liposome is one of the most favorable ultrasound-aided delivery methods. Bubble liposomes entrap the ultrasound imaging gas perfluoropropane, exploiting the cavitation bubbles produced by the pressure oscillations of ultrasound. This not only transiently enhances the permeability of a tissue or a cell membrane, reducing the thickness of the unstirred layer of the cell surface, and aiding DNA entry into cells,⁶⁶ but also affects the intracellular vesicles and trafficking after ultrasound exposure, thus enhancing the escape of gene cargo from the endosome to the cytoplasm and further transfer to the nucleus.⁶⁷ Negishi et al.⁶⁷ have used bubble liposomes and ultrasound exposure to enhance targeted liposome-mediated pDNA gene transfection, whereas Clumakova et al.⁶⁸ have demonstrated that their PLGA (poly(lactic-co-glycolic acid)) nanoparticle gene vector produced a significantly higher expression of the reporter gene in the tumor after a 5-minute ultrasonic treatment than that without ultrasound.

Heat and irradiation

Au nanoparticles (AuNPs) possess vivid optical properties with strong optical resonances associated with their surface plasmons. The optical absorption of AuNPs can be tuned from 690 to 900 nm in the near-infrared spectral range by varying the relative geometry and size of the core and shell. The plasmon-based optical properties of AuNPs assist in the photothermal ablation of solid tumors, providing a light-triggered release of short DNA strands conjugated to the surface of AuNPs.^18,69 Ni et al.⁷⁰ delivered DNAPK short hairpin RNAs by PSMA-targeting A10-3 RNA aptamers that selectively sensitized PSMA-positive cells to ionizing radiation, so that the toxicities to normal tissues were reduced.

Magnetofection

Liposomal magnetofection potentiates gene transfection by applying a magnetic field. Wang et al.⁷¹ constructed magnetic lipoplexes, which are self-assembling complexes of cationic lipids with plasmid DNA associated with superparamagnetic iron oxide nanoparticles. Liposomal magnetofection provided a threefold improvement in transgene expression over lipofection and knocked down the target protein in vitro. In vivo, the plasmid delivery efficiency into the tumor was significantly higher via liposomal magnetofection than lipofection.

Conclusion and Perspective

Various strategies have been developed to deliver gene cargos efficiently into target cells by nonviral vectors, which have attracted much attention in recent years. Rationally designed nonviral vectors have exhibited improved in vivo stability and pharmacokinetics, little RES uptake, high tumor accumulation, target specificity, efficient endosome release, and nuclear transcription of the encapsulated therapeutic nucleic acids. Even so, we have no doubt that the development of safe, stable, effective, and tumor-specific nanoparticles for systemic administration remains an unmet goal for successful clinical applications of cancer gene therapeutics.