Abstract
Nucleic acids are a promising type of therapeutic for the treatment of a wide range of conditions, including cancer, but they also pose many delivery challenges. For efficient and safe delivery to cancer cells, nucleic acids must generally be packaged into a vehicle, such as a nanoparticle, that will allow them to be taken up by the target cells and then released in the appropriate cellular compartment to function. As with other types of therapeutics, delivery vehicles for nucleic acids must also be designed to avoid unwanted side effects; thus, the ability of such carriers to target their cargo to cancer cells is crucial. This review will discuss classes of nucleic acids, hurdles that must be overcome for effective intracellular delivery, types of non-viral nanomaterials used as delivery vehicles, and the different strategies that can be employed to target nucleic acid delivery specifically to tumor cells. Moreover, nanoparticle designs that facilitate multiplexed delivery of combinations of nucleic acids will be reviewed.
Keywords: nanoparticles, targeted delivery, gene delivery, cancer therapy, nucleic acid
Graphical Abstract
Synthetic nanoparticles can be used for delivery of combinations of nucleic acid, with different materials conferring different advantages for safe and effective delivery. Targeting strategies can also be employed for cancer-specific delivery, including passive targeting based on nanoparticle physical properties; active targeting using cancer-specific ligands; transcriptional targeting using cancer-specific promoters; and targeting of cancer-specific proteins or signaling pathways.
1. Introduction
The delivery of DNA and other types of nucleic acid allows cells to be altered at the genetic level. By changing the gene expression profile of target cells, researchers can directly address the cause of diseases that have a genetic component, including not only inherited diseases but also acquired genetic diseases like cancers. For instance, malignant cells whose gene expression patterns are incorrectly regulated could potentially be treated by inducing overexpression of genes that induce apoptosis or by decreasing the expression of genes that promote cancer cell survival. However, delivery of nucleic acids to cancer cells remains a challenge. The ability to target cancer cells is crucial to maximize the efficiency of anti-cancer treatment as well as to avoid causing unwanted side effects on healthy tissues. Various types of nanoparticles have been studied for their ability to package nucleic acids by simultaneously protecting the payload from degradation and improving delivery precision.
In this review, we will discuss nucleic acids—largely DNA, mRNA, siRNA, miRNA, and immunostimulatory nucleic acids but also including a few examples of other classes—that have been studied for cancer therapy, as well as hurdles that must be overcome to deliver them to cells. We also cover types of nanomaterials used in the field and the properties that are important for gene delivery. Then, different methods of achieving cancer-specific delivery of nucleic acids will be reviewed, and finally, we will cover the promise and challenges of delivering multiple nucleic acids or combinations of nucleic acids with other therapeutic agents for cancer treatment.
2. Nucleic acids as cargo
This review will focus mainly on nucleic acids that are used to achieve one of three overarching goals: direct overexpression of a gene (e.g., using DNA or mRNA); knockdown of a gene using RNA interference (RNAi) by delivering small interfering RNA (siRNA) or microRNA (miRNA); and stimulation of the immune system by delivering nucleic acids that trigger pattern-recognition receptors (PRRs)[1]. While these are the most commonly studied in the field, examples of other types of nucleic acids will also be mentioned briefly, and their properties and delivery challenges are summarized in Table 1.
Table 1.
Examples of types of nucleic acid cargo, their properties, and delivery challenges.
| Nucleic Acid Cargo | Intended Action | Site of Action | Chemical Properties | Delivery Challenges |
|---|---|---|---|---|
| Plasmid DNA (pDNA) | Gene overexpression | Transcription to mRNA in nucleus |
|
|
| Messenger RNA (mRNA) | Gene overexpression | Translation to protein in cytoplasm |
|
|
| Small interfering RNA (siRNA) | Gene knockdown (affects specific gene) | RNAi pathway in cytoplasm |
|
|
| Micro RNA (miRNA) | Gene regulation (affects pool of genes) | RNAi pathway in cytoplasm |
|
|
| CpG oligodeoxynucleotides (ODNs) | Immune stimulation | Endosome |
|
|
| Cyclic dinucleotides (CDNs) | Immune stimulation | Cytoplasm |
|
|
DNA for overexpression of a gene is normally delivered as a double-stranded (ds) circular plasmid consisting of, at minimum, a promoter and a gene of interest.[2] DNA elements will be described in more detail below, as the choice of particular promoters and other design parameters are an important method by which to achieve cell-specific gene expression. In order to be expressed, the DNA must pass through not only the plasma membrane of the cell but also the nuclear envelope, and once in the nucleus, it can be transcribed into mRNA for export into the cytoplasm and translation into the protein of interest. Alternatively, single-stranded (ss) mRNA itself can be delivered. mRNA is chemically less stable than plasmid DNA and is more costly to produce at large scale; however, because it acts in the cytoplasm, it does not have to enter the nucleus as does DNA, simplifying its delivery path.[3, 4] Additionally, unlike plasmid DNA, mRNA carries no risk of undesired insertion into the genome, which could result in mutagenesis.
Translation of mRNA can be interrupted by RNAi, leading to decreased expression of a target gene.[5] The RNAi pathway has been studied for its uses as a tool for basic research[6] and for disease treatment,[7, 8] and its molecular mechanisms have been reviewed in detail elsewhere.[9] The pathway can be targeted by short lengths (~20 bp) of dsRNA. including siRNA, which is derived from cleavage of longer dsRNA strands by the enzyme Dicer, and miRNA, derived from cleavage of dsRNA with a hairpin loop structure. For both of these, one ssRNA strand can be incorporated into the RNA-induced silencing complex (RISC) in the cytoplasm. In the case of miRNA, the RNA-RISC structure then binds to mRNAs that have partial complementarity to the miRNA sequence. This incomplete sequence match allows miRNA to act on more than one mRNA sequence, leading to a range of potential effects, including repression or degradation of the mRNA molecule. The ability of miRNA to regulate pools of genes confers potentially powerful and far-reaching effects. By contrast, siRNA-RISC binds to an mRNA with complete complementary to the siRNA sequence, causing mRNA degradation and therefore decreased expression of a specific protein.[10] Because the effect of siRNA is generally limited to a single gene, unlike that of miRNA, its use could avoid off-targeted and unexpected side effects. Although siRNA is derived from long dsRNA strands under natural circumstances, researchers more commonly use synthetic ~20-bp siRNA directly as their cargo, bypassing cleavage by Dicer, as long dsRNA (>30 bp) has been shown to cause unwanted immune responses.[11]
However, in some cases, an immune response is the goal: there are several types of nucleic acids that have immunostimulatory properties; in fact, the RNAi pathway itself may have evolved at least in part as a way to combat dsRNA viruses.[12, 13] Some immune cells have receptors that recognize and are activated in response to certain types of nucleic acid as an innate defense mechanism against pathogens that carry foreign nucleic acids. Several examples of these exist, including DNA with unmethylated CpG sequences,[14] ssRNA in endosomes,[15, 16] and dsRNA,[17] all which are recognized by one or more of the well-studied Toll-like receptors (TLRs), whose signaling can lead to innate immune responses.[18] Cyclic dinucleotides (CDNs)[19] can induce a type I interferon (IFN) innate immune response by activating STING (stimulator of IFN genes). Even short siRNAs have been found to induce an innate immune response,[20] and thus, precautions must be taken to minimize off-target immune activation through engineering of the sequence or delivery vehicle; alternatively, if care is taken, this immune response can be leveraged as an additional deliberate function of siRNAs in addition to sequence-specific RNAi.
Finally, mRNA can also be cleaved by some types of DNAzymes, which are composed of a 15-nucleotide catalytic DNA sequence flanked by two binding regions.[21] The binding regions recognize a specific RNA sequence via Watson-Crick base complementarity, while the central catalytic region of this class of DNAzymes cleaves the bound mRNA. This molecule requires the presence of Mg2+ as a co-factor for activity, and it acts in a catalytic fashion: similarly to classical enzymes, it can dissociate from its mRNA substrate after cleavage and act on other mRNA molecules.
Common to all nucleic acids are their strong negative charge, their hydrophilic nature, and their relatively large size, with the molecular weight of siRNAs and miRNAs being on the order of 10 kDa and that of large DNA plasmids being on the order of 103-105 kDa. They are therefore difficult to deliver, as they cannot easily pass through cellular membranes and tissue spaces, and some types of nucleic acids, particularly RNAs and certain patterns of DNA, are easily recognized by the immune system and cleared. In addition, they are sensitive to degradation in the presence of nucleases. Even after successful entry into a cell, they must still overcome other intracellular barriers, including escape from the degrading conditions of the endosomal pathway, trafficking within the cytoplasm to specific intracellular locations, and/or entry into the nucleus.
Although the sequence of each nucleic acid can affect its function and its intended biological target, many of the physical and chemical considerations important to encapsulation in nanoparticles for delivery do not depend heavily on nucleic acid sequence. Thus, while differences in chemical properties and geometry among different nucleic acid cargos must still be considered, the co-delivery of multiple nucleic acids of the same type with differing sequences in the same delivery vehicle follow the same design principles as delivery of a single sequence of that same type. Co-delivery of different types of nucleic acids, on the other hand, can necessitate changes to nanocarrier design to efficiently deliver and direct different cargos to distinct cellular and subcellular locations where they can carry out their function. Combination therapy has been a major focus of study for cancer treatment, as, due to the heterogeneity of cancers among patients and even within a single patient, treatment with a single drug is often insufficient to destroy all of a patient’s malignant cells.[22] The intrinsic resistance of cancer cells to a drug, as well as well as acquired resistance through mutation, can lead to the selective survival of a population of malignant cells that does not respond to traditional treatment.[23] The delivery of multiple therapeutic agents in combination can target different cellular pathways at once, preventing the survival of a drug-resistant cell population, and the use of nanoparticles as delivery vectors for multiple drugs in combination has been extensively reviewed elsewhere.[24, 25]
Here, we will discuss the promise of combinatorial nucleic acid delivery for cancer therapy. Non-viral, synthetic nanoparticles, including polymeric or lipid-based nanoparticles and inorganic nanoparticles, can be designed for encapsulation or complexation of more than one nucleic acid, although, in every case, the advantages of co-delivering more than one gene or genetic sequence to achieve additive or synergistic effects must be balanced against the potential disadvantages of diluting the payload of each sequence. While research is being conducted on making viral gene delivery vectors safer and effective for delivery of multiple genes, they are still hampered by several challenges, including intrinsic toxicity or immunogenicity of viruses, which may be dose-limiting or prevent repeated administration, often needed for cancer treatment,[26–30] and cargo capacity that is limited by the viral capsid size, which makes the co-delivery of multiple genes difficult.[31] In this review, therefore, we will focus on non-viral nanoparticles for their versatility as combinatorial nucleic acid delivery vehicles.
3. Nanoparticles as nucleic acid delivery vehicles
Because of broad similarities in physical and chemical properties of nucleic acids and overlap in their extra- and intracellular trafficking routes, some major delivery challenges are applicable for many different types of nucleic acids. However, the specific delivery barriers that must be overcome may differ from one type of nucleic acid to another and should therefore influence the design parameters of nanocarriers. Some of these generally applicable considerations for nanoparticle-based nucleic acid delivery are discussed below and are also illustrated in Figure 1.
Figure 1. Challenges of nucleic acid delivery to tumors.
Effective and specific delivery of nucleic acids to tumors requires encapsulation or condensation of the cargo into nanoparticles. Nanoparticles must then remain stable in circulation, evading clearance and avoiding aggregation with other particles, and then leave the circulation to accumulate at the tumor site. Once there, particles must enter cells, and various intracellular barriers must be overcome depending on the type of nucleic acid cargo being delivered.
3.1. Nucleic acid binding or encapsulation
Nucleic acids on their own are vulnerable to degradation and various mechanisms of clearance. In order to reach cells intact, they generally must be loaded or condensed into nano-sized structures that can protect them from the environment and facilitate their trafficking to target sites, and a range of materials and strategies can be used to accomplish this. Many canonical delivery carriers are optimized to deliver one particular therapeutic agent, so encapsulating different types or cargo is a significant hurdle for combinatorial delivery. Many groups have taken advantage of the high negative charge of nucleic acids by complexing them electrostatically with cationic materials. Positively charged polymers like poly(L-lysine) (PLL), polyethylenimine (PEI), polyamidoamine (PAMAM), and poly(beta-amino ester)s (PBAEs) can bind and condense nucleic acids into nanoparticles via electrostatic interactions with amines.[32, 33] Similarly, cationic lipids have been used recently to co-encapsulate multiple types of nucleic acids into lipidoid nanoparticles (LNPs).[34] Liposomes contain both hydrophilic and hydrophobic moieties which can bind small molecule therapeutic agents for co-delivery with nucleic acids. Artificial vesicles like liposomes have an aqueous core that can be used to encapsulate the hydrophilic nucleic acids,[35–37] and other materials, like polyesters, can physically entrap nucleic acids via emulsification. In particular, PLGA-PEG carriers show desirable small molecule delivery, but their chemical properties and synthesis techniques are not conducive to co-encapsulation with nucleic acids. To improve nucleic acid loading in PLGA-PEG polymer nanoparticles, carriers can be loaded with synthetic nucleic acid analogs click nucleic acids (CNAs), which bind to nucleic acids in a sequence-specific manner.[38] On the other hand, to improve loading of small molecule drug into carriers optimized for nucleic acid delivery, the drug may be intercalated into therapeutic DNA prior to loading the agents into a delivery vehicle. Other nanoparticles, including inorganic nanoparticles, bind nucleic acids by covalent conjugation.[39, 40] One significant challenge that must be considered in combinatorial delivery is that the loading of one molecule into the carrier does not affect the loading of a second agent. The capacity of the carrier is a limiting factor in loading, but further unfavorable charge or hydrophobic interactions should be anticipated and considered when designing a vehicle for co-delivery. Loading the interfering agents into different zones or layers is a potential approach to overcome this challenge.
Differences in size and physical properties among nucleic acids can also affect loading and binding. For instance, CDNs, siRNA, and miRNA are much smaller than plasmid DNA and thus have fewer binding sites per molecule for electrostatic complexation, and this reduced multivalency can reduce binding affinity.[41–43] On the other hand, smaller cargos may be more amenable to chemical conjugation to nanoparticles. mRNA is intermediate in size between plasmid DNA and smaller RNA oligos, but, as mentioned above, ssRNA is a less stable molecule than dsDNA,[4] though both are vulnerable to enzymatic degradation by nucleases under certain conditions. It is thus particularly important to easily degradable nucleic acids like mRNA that binding and encapsulation by a biomaterial can have an important effect on the stability of the cargo in addition to condensing into a nanoscale particle. Electrostatic complexation or chemical conjugation of nucleic acids to cationic materials can protect them from enzymatic degradation;[44] encapsulation within a vesicle or solid nanoparticle can also provide a physical barrier to enzymes.[45, 46] Nucleic acid properties should be carefully considered when combining multiple types into a single carrier. Larger nucleic acids, such as plasmid DNA and mRNA, provide a high density of negative charge which can stabilize lipid or polymer carriers. Therefore, combination therapy with larger nucleic acid molecules can enhance the delivery and efficacy of smaller molecules, even if they act on distinct targets.[34] However, the relative amount of these stabilizing nucleic acids should be optimized, as very stable carriers with strong electrostatic interactions may not efficiently release their cargos.
3.2. Cellular uptake
Most of the nucleic acid cargos mentioned, including RNA oligos, mRNA, plasmid DNA, and immunostimulatory nucleic acids that act in the endosome or cytosol, must be internalized by target cells in order to function. Because the glycocalyx and lipid bilayer of cells have a net negative charge on the extracellular surface, cationic nanoparticles, in addition to their ability to bind nucleic acids, can facilitate cellular binding and uptake via electrostatic interactions.[47–50] Surface modification with certain moieties, such as cell-penetrating peptides (CPPs), can be used to further improve overall cellular uptake of nanoparticles.[51–55] In some cases, the oligonucleotides themselves can lead to cellular uptake by interaction with scavenger receptors,[56] as in the case of spherical nucleic acids (SNAs). In SNAs, nucleic acids are bound covalently to the surface of gold nanoparticles, the three-dimensional architecture of the nucleic acids drives uptake, even in later versions of SNAs in which the metallic core is dissolved and removed, leaving the nucleic acid shell.[57]
3.3. Endosomal escape and intracellular trafficking
Upon uptake, nanoparticles are generally localized in the endosomal-lysosomal compartment. Some immunostimulatory nucleic acids, such as agonists of TLR3 (dsRNA), TLR7/8 (ssRNA), or TLR9 (DNA with unmethylated CpG sequences) act in the endosome, where their receptors are located, and they can thus act in that compartment.
By contrast, plasmid DNA, mRNA, and RNA oligos avoid degradation in the endosome and then escape into the cytoplasm in order to be effective. Reversibly protonated cationic materials, including PEI and PBAEs,[58] can act as a buffer for protons that are pumped into endosomes, preventing the compartment from becoming overly acidified. This in turn leads to an influx of chloride into the endosome due to the building electrical gradient, followed by an influx of water due to osmotic pressure, which finally causes the endosome to lyse and release the nanoparticles into the cytoplasm. This “proton sponge” mechanism is regarded as one reason for effective endosomal escape by certain cationic nanoparticles.[59] Materials that can fuse with the endosomal membrane can also facilitate escape,[60] including CPPs,[33, 52, 53] which also aid in the initial cellular uptake. Other mechanisms have been reported recently to achieve endosomal escape for combinatorial nucleic acid delivery, including the inclusion of hydrophobic phenyl groups into dendrimers[61] and the linking of lipid moieties to a cationic aminoglycoside.[62]
Once out of the endosome, the nucleic acid cargo must be released from the nanoparticle, which can occur by a decrease in binding between cargo and delivery material and/or degradation of the material itself.[63, 64] Some common materials like polyesters and PBAEs degrade by hydrolysis, the latter over the course of hours at pH 7 and at physiological temperature. Interestingly, PBAEs are protonated at endosomal pH and thus less likely to deprotonate surrounding water molecules to form the stronger −OH nucleophile, and they are in fact less prone to degradation while inside the slightly acidic endosome than they are at neutral pH.[65] Thus, they can keep nucleic acid cargo intact while in the endosome and then release their cargo in the cytoplasm upon escape. Materials like PLGA and PLA have a wide range of different degradation kinetics, varying from days to weeks or even months, and their cargo can be released over time as they erode. The intracellular destination of nucleic acids may affect the choice of release mechanism when designing nanoparticles. For instance, siRNA, miRNA, mRNA, or other cargo that acts in the cytosol would benefit from a mechanism of fast, triggered cytosolic release. This can be facilitated by bioreducible moieties in the nanoparticle chemistry, as the cytosol of cells is a relatively reducing environment compared with the extracellular space.[66] Plasmid DNA, on the other hand, must enter the nucleus, and the specific barriers to nuclear entry have been reviewed previously.[67] In brief, diffusion of large biomacromolecules like plasmid DNA through the cytoplasm to the nucleus is very slow, although this can be improved upon by the attachment of a nuclear localization signal (NLS) to the plasmid or the nanocarrier,[68] promoting the binding of the DNA cargo to import proteins that can actively facilitate nuclear transport and entry.[67] It is therefore unsurprising that plasmid DNA delivery by non-viral nanocarriers has often been found to be much more efficient in dividing cells,[69–71] and while non-mitotic mechanisms of nuclear entry have been demonstrated,[70–72] nuclear transport remains a key challenge for non-viral delivery of plasmid DNA. Given the long history of using uncontrolled cell division as a method of targeting cancer cells while they are in mitosis,[73] however, this nuclear barrier to plasmid DNA delivery could in fact serve as a potential avenue to improving the selectivity of an anti-cancer gene therapy.
When combining multiple types of nucleic acids, or co-delivering a drug and nucleic acid, the desired release kinetics of the cargos must be considered. Sequential delivery may be desired, for example if a drug is intended to sensitize cells to a nucleic acid therapy or if the nucleic acid cargo knocks down drug efflux transporters to enhance the potency of a chemotherapy. In other cases, simultaneous action is preferred, such as for multiple nucleic acid cargos which act on an oncogenic pathway. To achieve precise release kinetics for each agent, the components may be incorporated into different compartments within the carrier, where the release of the agent is dependent on the degradation behavior of its encapsulating material. Layer-by-layer synthesis techniques can be used to deposit different drugs or nucleic acids to the surface of a carrier in a layered manner, which facilitates sequential release of the components.[74] Alternatively, nucleic acids may be conjugated to the particle surface using different chemistries with distinct release kinetics or stimuli-responsive behavior. To achieve complex release behavior of multiple agents, Badeau et al. have developed a modular system for logic gated release of various cells from a hydrogel.[75] By conjugating the cells with different linking moieties in series or in parallel, they achieved YES/OR/AND outputs from combinations of environmental stimuli, including enzymes and light. This strategy enables highly complex responsive release kinetics and has clear applications in combinatorial delivery systems.
3.4. In vivo stability of nanoparticles
The physicochemical properties of nanoparticles can affect their fate in vivo after administration. Nanoparticles are often cleared quickly from circulation by the mononuclear phagocyte system (MPS),[76, 77] a process that can be prevented or slowed by reducing interactions between nanoparticles and cells. One common method is to shield the surface of the nanoparticles by coating them with poly(ethylene glycol) (PEG), or PEGylation.[78] Grafting of PEG chains to nanoparticles has been reported to increase circulation time of various types of carriers and therefore allow accumulation at tumor sites,[79–82] as described in detail in the next section.
Because of the many necessary characteristics of nucleic acid delivery nanoparticles, all of the properties described here should be considered together. For example, while cationic materials are convenient for encapsulation or complexation of anionic cargo, positively charged materials can also cause toxicity by disrupting the cell membrane;[83] on the other hand, decreasing the molecular weight[84] or introducing biodegradable functional groups[85] into the polymer materials has been shown to increase the safety of the nanomaterial.[86–88] Thus, a combination of optimized parameters should be used to design an ideal nanoparticle.
4. Tumor Targeting
An overarching challenge in cancer therapeutics, including in nanomedicine, is achieving sufficient concentrations of an anti-cancer agent in the tumor while minimizing off-target toxicities in healthy tissues. Targeting methods take advantage of features that differentiate cancer from healthy tissue, including properties of the tumor microenvironment, overexpressed molecules on the cancer cell surface, and dysregulated gene expression. Nano-scale delivery vehicles are uniquely capable of passive targeting, active targeting via ligand functionalization, and controlled or triggered release. Additionally, nucleic acid cargos can themselves be engineered to take advantage of aberrant gene expression in cancer cells and achieve cancer specificity through their mechanism of action. Effective targeting is particularly important for combinatorial nucleic acid delivery. As discussed, certain combinations of nucleic acids can induce potent apoptotic or immunogenic effects, which are associated with toxicity in healthy tissues. Further, combining multiple nucleic acids will proportionally dilute the dose of each. Therefore, tumor targeting must be carefully considered when developing these delivery vehicles to maximize cargo delivery to cancer cells while avoiding dangerous off-target effects.
4.1. Passive Targeting
Hypervasculature, enhanced vascular permeability, and decreased lymphatic drainage are all hallmarks of rapidly growing tumors. The abnormal architecture of angiogenic tumor blood vessels underlies the enhanced permeability and retention (EPR) effect, which describes the tendency for systemically administered macromolecules within a 10–200 nm size range to accumulate in solid tumors.[89] Nanoparticle formulations of chemotherapies, such as Doxil® (for doxorubicin) and Abraxane (for paclitaxel), were developed in response to this phenomenon and have improved pharmacokinetics, an increased maximum tolerated dose, less systemic toxicity, and improved therapeutic efficacy compared to their free drug counterparts.[90] Nanoparticle size is a key predictor of passive targeting, because size affects clearance rate and route, extravasation into tumor tissue, cellular uptake, and interactions with the immune system. The range for passive targeting by the EPR effect is around 10–200 nm, although tumor accumulation has been described for slightly smaller or larger particles. Particle accumulation in the tumor compartment is in competition with clearance, either by the renal system or the MPS, which is comprised of macrophages predominately in the liver and spleen. Particles smaller than 10 nm are rapidly cleared by the renal system and tend to accumulate in the kidney, while larger particles accumulate in the liver and spleen.[91] Both clearance routes compete with accumulation in tumor tissue, so developing particles with low clearance rate and high circulation time is the primary goal for passive targeting by EPR. To further elucidate the effect of particle size on EPR and tumor accumulation, Perrault et al. systematically studied the biodistribution of 10–100 nm PEG-coated gold nanoparticles (mPEG-GNP) in subcutaneous MDA-MB-435 xenograft tumors.[92] Their results indicate that, for 40–100 nm particles, longer circulation times directly correlate with improved tumor accumulation. Smaller particles in the 20 nm range had lower tumor accumulation than larger particles, but these smaller particles had significantly enhanced permeation into tumor tissue.[92]
Inspired by the cylindrical or filamentous shapes of many viruses, non-spherical geometries have been explored for therapeutic drug and nucleic acid delivery. Long cylindrical, rod-like, disk-shaped, or filamentous particles have been shown to evade phagocytosis by resident macrophages. There is evidence that the contact angle between macrophages and particles affects phagocytic uptake; contact with a flatter surface, such as a rod or disk, causes macrophage spreading rather than phagocytosis.[93] Additionally, in dynamic fluid flow, filamentous particles elongate and align with flow, and hydrodynamic shear forces pull particles off of macrophages.[94] These effects slow the uptake of filamentous particles by the MPS and can dramatically increase circulation times. For example, Geng et al. developed filamentous micelles (filomicelles) comprised of a biodegradable co-polymer that remain in circulation 1 week after injection, while spherical versions are cleared within 2 days (Figure 2).[95] Increased circulation time is correlated with enhanced therapeutic efficacy of paclitaxel-loaded filomicelles in a subcutaneous xenograft model.[95] Micelle length appears to have been be a critical factor: an eight-fold increase in filomicelle length had the same therapeutic benefit as an eight-fold increase in paclitaxel dose.[95]
Figure 2. Shape effects of spherical vs filamentous micelles.
Filomicelles are self-assembled from diblock co-polymers (a) with nano-scale diameter and micro-scale length. The filomicelles extend in flow (b) and evade phagocytosis while spherical micelles in flow are internalized. When the micelles are injected systemically in mice, they persist in circulation for days, and longer micelles have a longer circulation half-life than shorter micelles. Filomicelles are efficiently internalized (d) by lung epithelial cells in static culture. Reproduced from Geng et al., “Shape effects of filaments versus spherical particles in flow and drug delivery,” Nature Nanotechnology 2:249–255, 2007,[95] with permission from Springer Nature.
Along with size and shape, surface properties are a key parameter affecting biodistribution and cellular uptake. For example, nanoparticles with a strong charge or a hydrophobic surface attract serum proteins, which adsorb to the surface and form a protein corona. Adsorption of opsonins, such as IgG or complement factor, tag the protein for clearance by the immune system. A protein corona can also block targeting ligands that have been conjugated to the particle surface. Therefore, a neutral hydrophilic charge is typically desirable for systemically administered nanocarriers. As mentioned above, PEGylation of particles is often used to reduce non-specific protein adsorption.[96, 97] PEG can be incorporated into block co-polymers, conjugated to the surface of inorganic particles, or incorporated into liposomal formulations, making it an attractive option for a variety of delivery applications. PEGylation reduces uptake by the MPS and dramatically extends circulation time, particularly when particles are coated with high-molecular weight PEG at a high density.[98, 99] Thus, many nanocarriers for cancer targeting are coated with PEG to increase blood circulation half-life and enhance passive targeting. However, bioinert hydrophilic carriers, due to their reduced binding to proteins and cellular components, also tend to have poor intracellular uptake and endosomal escape, and this related but unintended consequence is known as the PEG dilemma.[100] Strategies to overcome this dilemma, including active targeting and stimuli-responsive carriers, will be discussed in subsequent sections.
While EPR has been a significant discovery in preclinical models, passive targeting is a complex and highly variable process that depends on the size, degree of vascularization, and location or the tumor.[101, 102] There is also often a high degree of heterogeneity within tumors, with changes in cell density, interstitial pressure, and extracellular matrix composition affecting how nanoparticles move through different regions of tumor tissue.[103] These parameters are not easily recapitulated in vitro, and screening nanoparticles for in vivo delivery is traditionally low throughput and expensive. Recently, barcoded nanoparticles have been used to assess systemic nucleic delivery in a high throughput manner. Dahlman et al. incorporated barcoded oligos into nanoparticles formulated with different lipid compositions and pooled dozens of distinct formulations into a single systemic injection.[104] Using Illumina sequencing, this group was able to assess the effects of nanocarrier formulation on biodistribution and uptake of different nucleic acid nanoparticles. With this tool, the researchers have identified a successful liposome formulation for specific mRNA delivery to lung endothelial cells.[105]
4.2. Active Targeting
The large surface-to-volume ratio inherent to nanocarriers facilitates their interactions with biomolecules and cells. While methods described above have been employed to minimize this, thus extending the circulation time of nanocarriers, such interactions can also be leveraged as an advantage. Particles may be functionalized with active-targeting molecules that will specifically interact with the target and enrich particle accumulation in that site.[106–108] Nanoparticles, including those for combinatorial delivery, have been functionalized to target various surface macromolecules on cancer cells, including overexpressed or mutated proteins, altered glycoproteins or glycolipids, and cancer-associated fetal proteins. An alternative strategy is specifically targeting cell types which are susceptible to the particular nucleic acid combination. For example, targeting drug efflux pumps to deliver a nucleic acid combination to combat drug resistance, or targeting markers of difficult-to-treat stem-like tumor cells for combination miRNA delivery. Particles can also target molecules overexpressed in the tumor microenvironment, such as integrins, which are overexpressed on tumor vasculature.[109, 110] Optimizing active molecular targeting of nanocarriers requires careful selection of a conjugation chemistry by which to attach targeting molecules to the particle surface. Different conjugation strategies can be selected depending on the bulk particle material, targeting ligand, and desired application. It is possible to harness hydrophobic or electrostatic interactions to functionalize nanoparticles by nonspecific adsorption, which enables functionalization without chemical modifications or complex reactions.[111] However, this requires the use of large amounts of targeting molecule, and targeting ligands may be displaced by other biological molecules in a physiological environment. Thus, covalent conjugation is commonly used for irreversible attachment of targeting ligands. Biological molecules such as proteins or peptides contain primary amines, which react with activated carboxylic acid groups on a nanocarrier to form an amide bond.[112] Common carboxylic acid-activating compounds include carbodiimidazole (CDI), as well as carbodiimide compounds, such as 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), often in combination with N-hydroxysuccinimide (NHS), or dicyclohexylcarbodiimide (DCC).[113] Alternatively, cysteine residues in proteins or peptides contain thiol groups that can be reacted with maleimide-containing particles.[114] Proteins can also be functionalized with biotin for reaction with streptavidin-coated particles. To control ligand orientation on the particle surface, these reactive groups can be introduced at specific locations on the protein, either during synthesis or after protein purification.[108, 115] Particles can also be coated with Protein A or Protein G, which bind the Fc region of antibodies and facilitates properly oriented conjugation. Conjugating targeting molecules via a flexible linker, such as PEG, allows the conjugated ligand or antibody to rotate and move freely in space for optimal binding with a target.[116]
Conjugation methods can also dictate ligand density on the particle surface, which has proven an important parameter for optimization. For example, using various ratios of ligand-functionalized and unfunctionalized PEG for particle coating can significantly affect uptake by target cells.[117] While enriching the targeting ligand can lead to multivalent complexation and enhanced affinity, a saturation effect or even reduced binding has been reported.[118, 119] This effect has been attributed to multiple factors, including steric hindrance or suboptimal receptor clustering. Nanoparticle size and shape also dictates surface curvature and contact surface area, which can affect interactions between particles and target cells.[120] Therefore, although ligand density has been optimized in detail for particular nanocarriers,[121] optimal parameters vary greatly depending on the size, shape, and material composition of the particle as well as size, chemistry, and avidity of the particular targeting ligand.[122, 123] These parameters are summarized in (Figure 3).
Figure 3. Optimization parameters for cancer-specific nanocarriers.
Physical and chemical properties of delivery vehicles affect tumor accumulation, particle internalization and cargo delivery, and ultimately the therapeutic outcome. Classes of targeting moieties and their sizes are also summarized.
4.2.1. Antibodies and Fragments
Antibodies have been extensively explored as therapeutics because they can be engineered to target almost any antigen with a high degree of specificity.[124] Monoclonal antibodies have been used in cancer therapy for over 20 years, and patient responses are well-understood, so they are a natural choice for nanoparticle targeting.[125] While antibody therapies work by blocking or binding a receptor on a cancer cell, nanoparticles harness the specificity of antibodies while incorporating additional therapeutic modes of action by encapsulating a drug or imaging agent or acting as a therapeutic itself. For example, trastuzumab is a monoclonal antibody for HER2 that is used clinically for breast cancer, and conjugating this antibody to gold nanoparticles showed promise as a photodynamic therapy, a strategy to ablate tumor tissue with a high degree of precision.[126] Antibodies can also bind to endocytic receptors on target cells and facilitate cellular uptake, which is necessary for functional nucleic acid delivery.[127, 128] However, antibodies are bulky and significantly increase the size of conjugated nanoparticles. Safety concerns have also been raised, even with clinically approved monoclonal antibodies, regarding their immunogenicity.[129]
With the clinical success of antibody therapies came interest in developing molecules with the same specificity but smaller size. That initiative sparked the invention of next-generation antibodies, including single chain antibodies, domain antibodies, and nanobodies.[130–132] Single-chain antibody fragments have been extensively explored for active targeting, since they maintain the antigen-binding capability, are one fifth the size of full antibodies, and are considered safer than antibodies.[133] Antibodies and fragments are developed through display libraries, which preferentially and rapidly select for candidates which bind target cells and are internalized.[134, 135] Because fragments can be screened for intracellular uptake, functionalized particles have been explored for nucleic acid delivery. For instance, liposomes surface-conjugated to a melanoma-targeted antibody fragment showed a significant therapeutic benefit in melanoma lung metastases, while non-targeted control particles had no significant effect.[136] Because of their small size, antibody fragments targeting EGFR have been conjugated to very small particles (quantum dots and 10-nm iron oxide nanoparticles) for high-sensitivity diagnostic imaging.[137]
In particular, antibody targeting can enhance the particular action of nanoparticles for combinatorial nucleic acid delivery. Certain nucleic acids, such as drug efflux-targeting RNAi, are designed to act on drug-resistant cells, which can be targeted by surface receptors overexpressed on the membrane. For example, the drug efflux pump P-glycoprotein has been targeted using antibody-conjugated nanoparticles to deliver curcumin specifically to multi drug resistant cervical cancer cells.[138] In this case, active targeting significantly increased the specific uptake of PLGA nanoparticles and improved the efficacy of the therapy in vitro.
4.2.2. Ligands
Another approach to cancer cell targeting is functionalizing nanocarriers with ligands for overexpressed receptors, such as epidermal growth factor receptor (EGFR) and transferrin receptor. Many of these surface receptors are well-characterized cancer biomarkers, and ligands for these receptors have been identified and studied, which has led to their extensive development for active targeting. The natural ligands for overexpressed receptors range from proteins to carbohydrates to small molecules, such as vitamins.[139] Alternative small molecule ligands may also be developed using computational modeling and binding experiments, which expands targeting possibilities beyond the native ligand for a receptor.[140]
Cancer-specific receptors can have a variety of functions, but nucleic acid delivery benefits from active targeting to endocytic receptors, which facilitate cancer-specific uptake and intracellular delivery. The transferrin receptor is a commonly overexpressed endocytic receptor on cancer cells and can be targeted by transferrin (Tf) protein-coated particles.[141] Early preclinical work shows 20-fold higher tumor transfection with Tf-PEI-DNA particles than with non-targeted PEI-DNA particles, which corresponded with 30-fold lower off-target transfection in the lungs.[142] Transferrin-functionalized PEG-coated cyclodextrin particle CALAA-01 was the first RNAi therapeutic to be tested on patients in clinical trials. Systemically administered CALAA-01 was well-tolerated in phase 1 clinical trials, and there was dose-dependent nanoparticle accumulation and gene silencing in melanoma tumor biopsies.[143, 144] In another example, hyaluronic acid (HA) functionalization facilitates uptake via overexpressed CD44 receptors. HA can serve as a bulk scaffold material for self-assembled nanoparticles or as a coating to functionalize particles composed of alternate materials, such as mesoporous silica.[145, 146] In mouse studies, pre-treatment blockade with free hyaluronic acid reduces the concentration of HA-NPs localized in the tumor, indicating that active HA targeting plays a significant role in tumor accumulation.[147, 148]
Small molecule ligands also offer specificity to overexpressed receptors with minimal increase in particle size. Folic acid has been extensively studied as a targeting ligand due to its high affinity of the folate receptor, which is overexpressed in approximately 40% of cancer types, including breast, lung, ovarian, and colorectal cancers.[149, 150] Folate-conjugated therapeutics readily bind folate receptors and are internalized into the endosome through receptor-mediated endocytosis, which has enabled many folate-conjugated drugs and imaging agents to reach clinical trials.[151] Folic acid was one of the earliest molecules used for nanoparticle targeting, but recent advances have further enhanced its efficacy as a cancer-targeting ligand.[152] Advances in linker chemistry have enabled efficient recycling of folate receptors back to the cell membrane after particle internalization, which facilitates continuous gene delivery to cancer cells.[153] Incorporating PEG into nanoparticle formulations reduces non-specific endocytosis by healthy cells that do not express folate receptor, and PEG linkers also enhance folic acid binding with membrane-bound receptors.[154] Prostate-specific membrane antigen (PSMA) has also been widely investigated as a receptor target due to its expression in many prostate tumors as well as the neovasculature of other solid tumors. Small molecule 2-[3-[5-amino-1-carboxypentyl]-ureido] pentanedioic acid (ACUPA) is a high-affinity small molecule ligand of PSMA which has shown targeting properties in vivo and in vitro.[155, 156] BIND-14, a PEG-PLGA nanoparticle with ACUPA active targeting for docetaxel developed by BIND Therapeutics, was evaluated in clinical trials for PSMA-targeted docetaxel delivery to prostate cancer.[123] There was no significant toxicity, and a 12% overall response rate was observed in phase I clinical trials.[157] ACUPA has recently been employed for PSMA-targeted delivery of siRNA to prostate cancer cells. For example, Xu et al. developed a nanocarrier comprised of a pH-responsive polymer blend coated with PEG-ACUPA for PSMA targeting (Figure 4).[158] This delivery vehicle was used to deliver siRNA targeting prohibitin 1 (PHB1), which is overexpressed in prostate cancer.[158] Briefly, nanoparticles extravasate through leaky tumor vasculature and associate specifically with prostate cancer cells via interactions between ACUPA and PSMA. The pH-responsive polymer triggers rapid disassembly upon internalization, and the oligoarginine sharp domains facilitate endosomal escape and efficient siRNA delivery.[158] ACUPA-targeted NPs significantly suppressed tumor growth over non-targeted NPs after 30 days, showing the benefit of active targeting in this case.[158]
Figure 4.
A prostate cancer targeted multifunctional envelope-like nano device (MEND) (A) nanocarrier is synesized by siRNA self-assembly with two block co-polymers: sharp oligoarginine functionalized pH responsive Meo-PEG-b-P(DPA-co-GMA-Rn) and PSMA targeted ACUPA-PEG-b-PDPA. Schematic shows targteted intracellular siRNA delivery after IV administration of MENDs. This strategy enables efficient gene silencing and significantly slows LNCaP tumor growth (B) compared with control and non-targeted NPs. Representative images of tumor bearing mice on day 18 (C) and photographs of harvested LNCaP tumors afetr 30 days (D). Reprinted with permission from Xu, Xiaoding, et al., “Multifunctional envelope-type siRNA delivery nanoparticle platform for prostate cancer therapy,” ACS Nano 11(3): 2618–2627.[158] Copyright (2017) American Chemical Society.
Diseased or cancerous cells often display a different array of surface glycans compared to healthy cells, offering another targetable feature.[159, 160] Lectins can be used to target drugs or nanoparticles specifically to cells displaying these abnormal features.[161] One useful example is targeting asialoglycoprotein receptors on liver cancer cells with galactosamine-conjugated nanoparticles.[162–164] Lectin conjugation can also facilitate transport across the blood-brain barrier, which typically serves a major hurdle for delivery to brain tumors. For example, nanoparticles modified with wheat germ agglutinin enhanced delivery to the brain two-fold over unmodified nanocarriers.[165]
An alternative to direct ligand conjugation is coating nanocarriers with cancer cell membranes. Recent research describes homotypic cancer cell binding—a phenomenon where cancer cells preferentially bind to membranes which carry the same surface antigens.[166] To harness this feature, researchers have coated nanoparticles with modified and unmodified cancer cell membranes, which provides a stealth coating as well as tumor homing properties.[167] This straightforward approach enables particle functionalization with the complete range of tumor ligands in a biomimetic manner and facilitates cancer-specific accumulation and uptake.
4.2.3. Aptamers
As discussed, many ligand-based active targeting systems are restricted to a biological ligand binding its natural target receptor. Aptamers are single-stranded DNA or RNA oligos that serve as attractive targeting alternatives. These nucleic acid molecules form sequence-specific three-dimensional structures and can be engineered to bind virtually any target, from small molecules and single amino acids to proteins, carbohydrates, and whole tumor cells. Targeted aptamers are selected from a large library of random sequences using systematic evolution of ligands by exponential enrichment (SELEX).[168] In each round of this process, the aptamer library is exposed to the desired target, unbound sequences are washed away, and bound sequences are selectively eluted and amplified using PCR. The process is repeated in subsequent selection rounds with more stringent binding conditions to generate tightly binding aptamers with antibody-like affinity and specificity.
Aptamers have been extensively explored in the past decade and have yielded novel selection methods for cancer targeting. Cell SELEX selects for aptamers that bind to a monolayer of cultured cancer cells, rather than purified antigen.[169] This selection method generates a diverse pool of targeting aptamers that bind to or are internalized by the target cell population (Figure 5).[169–171] Methods have also been developed for in vivo SELEX, where aptamers are selected by binding tumors in situ in animal models.[172] This has resulted in aptamers with minimal off-target binding to healthy tissues that can bind multiple cell types in heterogeneous tumors.[173] Because aptamers can be generated for a range of molecules expressed in tumors, this targeting approach circumvents the problem of resistance, which is common when targeting a single receptor.[174]
Figure 5.
Schematic (A) illustrating the selection process for prostate cancer-specific internalizing RNA aptamers. Nanoparticles coated with prostate cancer-specific internalizing aptamers are specifically taken up in target PC3 cells (B) to a higher degree than in non-target HeLa cells. Bare particles without aptamer are taken up at low levels in both target and non-target cells, so aptamer conjugation is necessary for target-specific uptake. Uptake is distributed throughout the cytosol of targeted calls (C). When particles are loaded with Docetaxel (D), the aptamer conjugated particles (Dtxl-NP-Apt) are significantly more potent non-targeted particles (Dtxl-NP) at killing target cells. Reprinted with permission from Xiao, Zeyu, et al., “Engineering of targeted nanoparticles for cancer therapy using internalizing aptamers isolated by cell-uptake selection.” ACS Nano 6(1): 696–704.[171] Copyright (2012) American Chemical Society.
Aptamers are attractive for nanoparticle targeting because they have low molecular weight and can include chemical modifications for particle conjugation.[175] In one example, PSMA aptamer-targeted PLGA-PEG nanoparticles loaded with docetaxel were used to treat animals with subcutaneous LNCaP prostate tumors, and the survival rate was 100% in treated animals after 109 days.[176–178] Aptamer-nanoparticle conjugates have also been investigated for cancer detection and imaging with quantum dots and iron oxide nanoparticles.[179–181] More recently, peptide nanoparticles functionalized with doxorubicin and MUC1 aptamer were used for targeted delivery of chemotherapy to cancer cells while simultaneously monitoring drug release in real time.[182] The anti-VEGF aptamer therapy pegaptanib was clinically approved in 2004 to treat macular degeneration,[183] and other aptamers are in clinical development to treat conditions ranging from thrombosis[184, 185] to leukemia.[186] Aptamer functionalized nanoparticles could be a promising approach for high-specificity binding to newly emerging cellular targets.
4.2.4. Integrin Targeting
The integrin profile of tumors is distinct from that of healthy tissues, and several integrins are upregulated on tumor endothelial cells and cancer cells.[187, 188] The αVβ3 integrin is significantly upregulated in many cancers and can be targeted with a simple RGD peptide motif. RGD-targeted nanoparticles bind selectively to αVβ3 integrins, and this interaction facilitates selective endocytosis into angiogenic endothelial cells and cancer cells.[189] This approach has been used to efficiently deliver nucleic acids to tumor vasculature, for example to deliver siRNA against VEGF-R2 and inhibit both angiogenesis and tumor growth.[190] Because αVβ3 is expressed on both endothelial cells and cancer cells, RGD facilitates cancer cell targeting in addition to endothelial targeting. This dual targeting is a major advantage of RGD-functionalized nanocarriers. RGD-targeted chitosan nanoparticles containing siRNA have been used to successfully downregulate drug efflux transporter P-glycoprotein expression and reverse multidrug resistance in a breast cancer model.[191] Other integrins can serve as therapeutic targets, including arresten (α1β1), canstatin (αvβ3 and αvβ5), angiostatin (αvβ3), tumstatin (αvβ3), endostatin (αvβ3, αvβ5, and α5β1), and endorepellin (α2β1).[192] Nanoparticles have been used to deliver these anti-angiogenic agents[193] or the genes that encode them.[194]
Despite promising preclinical data for active targeting in nanomedicine, no FDA-approved nanocarriers have employed active targeting strategies. Active targeting increases the complexity and potential immunogenicity of a drug delivery system, which makes it more difficult, time consuming, and expensive to develop. Further complicating their optimization is the binding site barrier effect, where antibodies or nanoparticles bind target cells with high affinity and cannot penetrate throughout the tumor.[195, 196] BIND-014 was an early targeted nanocarrier to enter clinical trials, and it benefitted from rigorous and systematic optimization of particle properties (size, surface properties, drug loading etc.) in preclinical studies and biodistribution validation in multiple animal models (mouse, rat, and monkey).[123] This rigor is essential when increasing the complexity of a platform and should be a model for active targeting in the future.
Target identification remains a major hurdle, and even well-characterized targets are almost always heterogeneously expressed within tumors. Additionally, many of the receptors overexpressed on cancer cells, including transferrin and folate receptors, are also expressed on proliferating healthy cells, so targeting these receptors can lead to off-target side effects and systemic cytotoxicity.[197] Thus, efforts should be focused on developing companion diagnostic methods to characterize target expression and predict patient response. Radiolabeled tracers based on the RGD peptide sequence are already in clinical development for monitoring αVβ3 integrin expression in patients.[198] Such advanced diagnostic tools are needed to study biomarker distribution within patient populations and assess the feasibility of actively targeted nanocarriers for personalized medicine.
4.3. Stimulus-Responsive Targeting
As tumors grow and develop, cancer cells exist in a constantly changing environment, influenced by high cell density and low blood supply. Solid tumors have an abnormally acidic pH, are subjected to low oxygen, and have a high concentration of certain enzymes.[199] These properties of the tumor microenvironment are targetable features that can be exploited for nanoparticle targeting and controlled release. To respond to an environmental stimulus, nanocarriers must include responsive chemistries that change the properties of the particle when it encounters a trigger. This triggered response can expose binding domains, dismantle a protective coating, or change particle surface properties to facilitate cancer-specific uptake.
Zwitterionic polymer nanoparticles are neutrally charged in physiological conditions, which confers stability, resists serum protein adhesion, and prevents clearance.[199] In the slightly acidic tumor microenvironment, the zwitterionic polymer becomes protonated and switches to cationic, facilitating uptake into cancer cells.[200] Cleavable PEG linkers can similarly be used to facilitate tumor cell uptake. Inert PEG coatings are commonly used in drug and gene delivery to enhance particle stability in circulation and increases the circulation half-life. However, as mentioned above, PEG coating also tends to decrease cellular uptake of particles.[100] To address this, researchers have coated particles with cleavable PEG, which is released in response to a trigger. For example, matrix metalloproteinases (MMPs), enzymes associated with angiogenesis and tumor growth, have been explored for triggered PEG de-shielding in a tumor-specific manner.[201] A multifunctional envelope-type nano-device (MEND) functionalized with MMP-cleavable PEG maintained the prolonged serum stability characteristic of PEG functionalized particles. Further, carriers conjugated with cleavable PEG exhibited superior in vitro and in vivo tumor transfection over carriers with non-cleavable PEG.[202, 203]
Responsive vehicles can also facilitate cytosolic release of nucleic acid cargo, which is particularly important for an efficient therapeutic response from RNA, since these molecules act in the cytoplasm. Polymer nanoparticles often incorporate disulfide bonds, which degrade in the reducing environment of the cytoplasm and release their cargo. Using a bioreducible polymer reduces the toxicity of the nanocarrier while ensuring full protection of the easily degraded RNA cargo. pH can also trigger release in the acidic endosomal environment for site-specific cellular delivery. Acidic pH can trigger a shape change or expansion which disrupts the endosome and releases nucleic acid directly in the cytosol. Lipid-based liquid crystalline nanoparticles, termed nano-transformers, expand in a pH 5 environment from needle-like structures to nanospheres.[204] This shape transformation was proposed to have improved endosomal escape by promoting membrane fusion with the endosome. In another approach, Luo et al used miRNA-catalyzed release to specifically trigger payload release in the presence of miRNA-21, which is overexpressed in many cancers.[205] Similarly, the DNA “nanosuitcase” developed by Bujold et al. opens conditionally in the presence of a miRNA or mRNA and releases its therapeutic oligo cargo.[206] This approach allows triggered intracellular release in response to a genetic biomarker, limiting its effects to cancer cells. Responsive vehicles for intracellular delivery ensure efficient and specific cargo release to allow all components in a combinatorial system to act simultaneously, which is likely important to achieve synergistic effects.
Systemically administered nanocarriers can also be triggered by an external stimulus to enhance gene delivery at the tumor site. The properties of thermoresponsive polymer particles change in response to externally applied heat or cold, which can be harnessed for tumor-specific gene delivery.[207] Poly(N-isopropylacrylamide) (PNIPAM)-based polymer nanocarriers undergo a hydrophilic to hydrophobic transition when temperature is raised from 37° C to 42°C, which causes the particles to aggregate and enhances endosomal escape, resulting in 2 orders of magnitude enhanced transfection at hyperthermic sites.[208] Light-responsive particles have also been used to enhance endosomal escape in a spatially controlled manner.[209] To harness this for siRNA delivery, a phtotosensitizer was combined with siRNA and encapsulated using the Lipofectamine™ commercial transfection reagent, and cells stimulated with light showed 10-fold higher silencing than non-stimulated cells.[210] Though this strategy has shown efficacy in vitro, photodynamic therapy in vivo is limited by the penetration depth of visible light through tissue. Magnetic coating or encapsulation of magnetic material can be used for magnetically guided nucleic acid delivery to a targeted site.[211, 212] An applied magnetic field is used to concentrate or retain gene delivery particles at the tumor site, which has been used to enhance the delivery of DNA and siRNA in a process termed magnetofection.[211, 212]
Ultrasound has been extensively used in combination with microbubbles for regio-specific delivery of anticancer agents. Briefly, a nucleic acid nanocarrier is co-delivered with gas microbubbles, which are clinically approved for use as a contrast agent.[213] Co-localization of microbubbles and nanocarriers is essential for successful transfection, so the nanocarriers must be coupled to the surface of the bubble using covalent conjugation. Then, ultrasound is applied to the target area, and passage of ultrasound through tissue creates pressure waves, which cause the microbubbles to undergo cavitation and release energy that opens transient pores in surrounding cells. The co-delivered nanocarrier enters the cell through these sub-micron pores, which leads to significantly enhanced transfection at the site where ultrasound was applied. DNA-containing PEI polyplexes conjugated to microbubbles combined with ultrasound stimulation enabled gene delivery to implanted tumors in a mouse kidney with 40-fold higher expression in tumor tissue than control non-sonicated tissue.[214] In a similar approach, liposome-bearing microbubbles enhanced the delivery of an anti-fibrotic miRNA to diseased liver in rats.[215] Microbubble cavitation can also be used to open the blood-brain barrier and allow systemically administered nanocarriers to reach tumors in the brain.[216] Because nanocarriers enter through pores in the cell membrane, they bypass endosomal uptake and are delivered directly to the cytosol.[217] The method is also non-invasive and provides precise spatial and temporal control over nucleic acid delivery. However, the precise location of the tumor must be known to effectively apply the ultrasound to the target site. Image-guided focused ultrasound has been used to localize ultrasound signal more precisely, particularly for opening the blood brain barrier.[218, 219] Still, this approach is limited to primary tumor sites rather than dispersed metastases.
4.4. Local Administration
When possible, local administration can increase particle concentrations at the target site while decreasing healthy tissue exposure. Direct intratumoral administration is not an option in most cases, as accessing the tumor would involve an invasive procedure. Surgical tumor resection can be used as an opportunity to deliver a therapeutic directly to the site, where any remaining cancer cells can be treated to prevent recurrence. Implanted drug delivery depots can deliver a therapeutic at a controlled dose over the course of weeks or months with a single implantation surgery.[220] Nucleic acid therapeutics have been encapsulated in hydrogel depots for local delivery and reduced systemic toxicity. Naked DNA can be incorporated in hydrogels for regenerative medicine and anti-cancer applications, but only low levels of gene transfer have been observed due to the absence of a carrier.[221, 222] DNA/PEI polyplexes have been successfully encapsulated in hydrogels and show effective gene transfer in vitro and in choriallantoic membrane assays.[223, 224] siRNA-loaded micelles have similarly been encapsulated in injectable polyurethane scaffolds for sustained local gene silencing.[225] The bulk hydrogel can be tuned to achieve the desired release kinetics, including varying material, molecular weight, crosslinking density, size, and geometry. Another advantage of hydrogels is the ability to encapsulate multiple separate agents in a single gel and simultaneously target cancer by multiple modes of action. Locally implanted hydrogel patches have been used to simultaneously deliver chemotherapy, siRNA, and gold nanoparticles for photothermal therapy.[226] This strategy can also be applied to deliver nucleic acid cargos with different properties, for example RNA and DNA, at the same site with particles optimized for each particular cargo. Nanocarrier properties can be independently tuned for efficient and targeted nucleic acid delivery, but maintaining stability and bioactivity of encapsulated particles is a significant hurdle to successful transfection.
Certain tumors can be accessed with non-surgical and non-invasive delivery routes, which reduces the potential for complications and enables frequent repeated particle dosing. Aerosol delivery provides direct access to lung tissue and can be used as a delivery route for lung cancer.[227, 228] An inhalable cationic liposome formulation for plasmid DNA is in clinical trials to treat cystic fibrosis, and a hyperbranched PBAE nanoparticle recently was used for mRNA delivery to lung epithelium (Figure 6).[229, 230] Patel et al. reported that hDD90–118 polypexes remained stable after aerosolizing with a vibrating mesh nebulizer, which produced micro-sized droplets ideal for distribution throughout lung tissue.[229] They were able to transfect 24.6% of lung epithelial cells after a single dose, with transfection seen in all five loves of the lung.[229] Delivery to other tissues, including the liver, spleen, and heart, was negligible, and there was no observed local or systemic toxicity.[229] While this platform has not been employed in lung cancer models, cancer treatment is an obvious potential application of these new inhalable technologies. The skin is uniquely accessible for local delivery, so topical applications are being explored for nanoparticle gene delivery to skin cancer.[231] Nanocarriers can be used to control permeation through the skin, and transport properties can be controlled independently of the cargo.[232] Chitosan nanoparticles have been used for antisense oligonucleotide and plasmid DNA delivery to skin, with reporter gene effects persisting on rat skin for 6–7 days.[233, 234] Free nucleic acid had no measurable transfection, indicating that a drug silvery system is required to effectively protect and deliver topically applied nucleic acids. Finally, oral delivery of nanoparticles allows direct access to tumors of the gastrointestinal tract.[235, 236] Chitosan nanoparticles have also been explored for this purpose, as they are stable upon oral administration, can encapsulate nucleic acid cargos, facilitate transport across the intestinal wall, and can be functionalized for active tumor targeting.[237] Delivery to different tissues requires overcoming different barriers, which can be physical (surfactant, mucous, stratum corneum), biological (enzymes, resident immune cells, blood brain barrier), and chemical (harsh pH of the GI tract). Future work in this area must focus on enhancing particle stability, controlling release kinetics, and improving transport and permeability across these tissue-specific barriers in orthotopic tumor models.
Figure 6.
A vibrating mesh nebulizer (A) was used to prepare luciferase mRNA delivery vectors for aerosol administration. Nano-scale polyplexes were encapsulated in micron-sized droplets and administered to a whole-body chamber. Hyperbranched PBAE hDD90–118 polyplexes enabled high levels of luciferase delivery in the lungs (B) after 24 hours, and local delivery by inhallation resulted in highly specific delivery to lung tissue and negligible off-target luciferase (c) measured by bioluminescence. Particles maintained a similar size and morphology before and after nebulization, characterized by electron microscopy (D). Particles also have a narrow size distribution before and after nebulization (E). Reprinted with permission from Patel, Asha Kumari, et al., “Inhaled Nanoformulated mRNA Polyplexes for Protein Production in Lung Epithelium,” Advanced Materials (2019):e1805116,[229] with permission from John Wiley and Sons.
4.5. Cancer-specific nucleic acid therapies
Cancer is fundamentally a condition caused by dysregulated gene expression, and nucleic acid therapies can treat the genetic basis of the disease by counteracting observed genetic changes. Advances in high throughput sequencing, microarray technologies, and novel computational models have resulted in a rapidly growing understanding of the genetic basis of cancer. We are now beginning to understand how particular genetic mutations and expression profiles correlate with disease stage, drug resistance, and potential for metastasis. Therefore, selecting certain nucleic acid cargos can target disease that is more aggressive, drug resistant, or likely to metastasize. Combinatorial strategies greatly increase the number of possible nucleic acid combinations that may be used to target a heterogeneous tumor population. Additionally, as illustrated thus far, combinatorial approaches have been successful in slowing tumor growth in these aggressive and difficult to treat cases.
One common approach is restoring the function of a mutated tumor suppressor with exogenous nucleic acids. Mutations in tumor suppressors have been associated with resistance to chemotherapy and radiation.[238, 239] Tumor suppressor function can be restored by introducing nucleic acids, including DNA and mRNA, that encode for the wild-type protein.[240] For example, systemic delivery of PTEN mRNA using polymer-lipid hybrid nanoparticles was shown to slow tumor growth in multiple models of prostate cancer.[241] Modified mRNA shows enhanced stability compared to plasmid DNA therapies in systemic circulation with more predictable and desirable protein kinetics. P53 is another tumor suppressor that has been targeted in many clinical and preclinical trials.[242, 243] Studies have shown that introducing p53 induces apoptosis in many cancer cells, but healthy cells only experience cell cycle arrest, which suggests inherent cancer-specificity to p53 gene therapy.[244, 245] However, certain healthy cell types, including epithelial and hemopoietic cells, are sensitive to p53-induced apoptosis.[246, 247] Also, p53 is part of a complex interconnected web of factors, and the efficacy of these therapies is limited by mutations in downstream factors and epigenetic changes, as well as heterogeneous p53 expression throughout tumors.[247] While p53 mutations have been explored most extensively due to their prevalence, other tumor suppressors (e.g., Rb7, PTEN, or mda-7) have also been targeted.[248]
An alternative approach is silencing of overexpressed oncogenes using RNAi. Many strategies involving siRNA or miRNA have been developed to target oncogenes for cancer therapy. Ideal targets are upregulated in cancer cells, are vital for cancer progression, and do not have a rapid turnover rate. Examples of targeted pathways include angiogenesis (VEGF), proliferation (FAK), survival (Bcl-2, survivin), cell cycle (PLK1, cyclin B1), and resistance to chemotherapy or radiation (c-myc).[249] Knocking down these pathways can have a potent anti-cancer effect but can also affect the viability of healthy cells and tissues, leading to systemic toxicity. Brummelkamp et al. approached this problem by developing siRNA that specifically targets the oncogenic K-RASV12 allele without any effect on wild-type K-RAS expression, which is required for normal cell survival.[250] They showed that knocking down K-RASV12 with a viral vector completely prevented CAPAN-1 pancreatic cancer cell growth in vitro as well as in a subcutaneous tumor model. This approach is promising for specifically targeting cancers with mutations or chromosomal translocations that produce mRNA transcripts distinct from those expressed in healthy cells.
Cancer cells are also more vulnerable to cell death through certain proapoptotic pathways, including TNF,[251] Fas,[252] and Bcl.[253] Tumor necrosis factor (TNF)-related apoptosis-inducing ligand (TRAIL) induces apoptosis with a strong cancer selectivity, due to overexpression of TRAIL-binding death receptors on cancer cells.[254] Because recombinant TRAIL proteins have poor pharmacokinetics and a short circulation half-life, nucleic acid therapies are a promising approach to achieve sustained TRAIL expression. Co-delivery with small molecule sensitizers, including clinically approved chemotherapies, has been shown to reverse TRAIL-resistance and improve antitumor efficacy in a synergistic manner. A dendrimer nanocarrier co-encapsulating doxorubicin and a plasmid DNA expressing human TRAIL induced synergistic growth inhibition in U87 glioma cells.[255] This treatment administered intravenously induced observable apoptosis in an orthotopic murine glioma model, and the co-delivery vehicle extended median survival to 57 days, compared to 34 days with doxorubicin alone. Another strategy for TRAIL therapy is to transfect or transduce tumor-homing stem cells ex vivo to secrete TRAIL in the vicinity of the tumor.[256] This strategy combines regio-selective delivery of the stem cells with the cancer-specific TRAIL therapy and shows significant survival benefit in an aggressive brainstem glioma model.[257] A secretable TRAIL construct has also been developed to enhance the bystander effect to neighboring cancer cells and further potentiate the antitumor effect.[258]
Gene expression in cancer cells is modulated by differential expression of transcription factors, which can be exploited to restrict the expression of therapeutic genes to tumor cells. By placing transgenes under the control of certain promoters, it is possible to achieve tissue-specific, cell-specific, or exogenously stimulated expression.[259] For example, hTR and hTERT promoter drive telomerase activity, which is a common feature of most cancers and can activate genes in a cancer-targeted manner.[260] Tumor-specific promoters ideally have strong expression in cancer cells and little to no expression in healthy cells. For example, alpha fetoprotein (AFP) is transcriptionally silent in adult liver but is expressed in 70–80% of hepatocellular carcinoma cases.[261] When potent pro-apoptotic genes are placed under the control of the AFP-promoter, cell death is restricted to AFP-producing HCC cells, and no acute systemic toxicity is observed.[262] Transcriptionally targeted BikDD DNA delivered by DOPC-cholesterol liposomes prolonged survival in multiple xenograft and syngeneic orthotopic murine HCC models (Figure 7).[262] Histological staining indicated that treatment-induced apoptosis was restricted to liver tumor cells, and cell death was not observed in healthy liver.[262] Many other tumor-specific promoters have been identified for particular cancer types, each with a different prevalence, expression profile, promoter strength, and tumor target. Another approach is to use promoters that respond to the tumor microenvironment. Hypoxia response elements (HREs) can be used to drive gene expression in hypoxic tumor environments, where cells are typically resistant to chemotherapy and radiation. Glucose-responsive promoters, such as hexokinase 2 and GRP78, respond to low glucose and high catabolism in tumors.[263–265] Finally, inducible promoters have been developed to respond to exogenous stimulation. Examples of stimuli include radiation, hyperthermia, and small molecule drugs.[266–268] These promoters do not depend on the expression profile of the tumor, so the strength and duration of activation can be controlled. Ultimately, transcriptional targeting offers an opportunity to control expression at the cellular level, which can dramatically reduce off-target toxicity.
Figure 7.
DOTAP-cholesterol liposomes loaded with transcriptionally targeted eAFP-VISA-BikDD or non-targeted CMV-BikDD were I.V. injected in orthotopic ML-1 tumor-bearing mice. Both particles significantly reduced tumor burden (a) and representative photos are shown from 1 week after the last treatment. Mouse survival (b) was significant on treatment groups, and the transcriptionally targeted DNA therapy extended survival significantly compared with the non-targeted DNA. Tissue samples from mice in (a) were fixed and stained for apoptosis using a TUNEL assay (c). The percentage of apoptotic cells were quantified in random fields from both tumor and healthy liver. While the targeted and non-trageted therapies induced similar numbers of apoptotic cells in the tumor, the transcriptionally targeted DNA induced less apoptosis in the healthy liver tissue. Reprinted with permission from Li, L. Y., et al., “Targeted hepatocellular carcinoma proapoptotic BikDD gene therapy,” Oncogene 30(15):1773, 2011,[262] with permission from Springer Nature.
Combinatorial delivery can also be used to counteract phenotypic changes in a particular subset of cancer cells, such as drug resistant or stem-like cells. A combination of miRNAs can be used to target multiple pathways involved in the stem-like phenotype of brain tumor initiating cells. This minor population of glioblastoma cells has been associated with tumor reoccurrence and drug resistance. Using pooled miRNAs to revert the stem-like traits in these cells has proven successful in inhibiting growth, neurosphere formation, and shrinking orthotopic xenografts in mouse glioblastoma models.[269, 270] Therefore, particular combinations of nucleic acids can target certain phenotypes associated with aggressive or reoccurring disease. Pooled siRNAs have also been used as a treatment strategy to overcome multiple drug resistance. In one example, two siRNAs (anti-Pgp and anti-Bcl-2) were combined with epirubicin in a calcium phosphate inorganic nanoparticle as a treatment for drug-resistant liver cancer cells and tumors.[271] The study concluded that the combination therapy was effective due to simultaneous targeting of two drug resistance mechanisms: pump (anti-Pgp) and non-pump (anti-Bcl-2).[271] Therefore, pooled nucleic acid combination therapies can act synergistically and address tumor heterogeneity by acting on multiple pathways with spatial and temporal synchronization.
Nucleic acid therapeutics are distinct in that Watson-Crick base pairing allows for sequence-dependent manipulation of three-dimensional structure. Changing the architecture of the nucleic acid can confer favorable properties or additional specificity. For example, Conde et al. developed an RNA triple helix comprised of two different tumor suppressor miRNAs and one antagomir which inhibits a pro-cancer oncomiR.[271] Combining these molecules into a triple helical structure imparts high structural stability and facilitates co-delivery.[271] Similarly, Li et al. developed tetrahedral RNA structures with precise control over their geometry.[272] The 2’ -F modified RNA tetrahedrons were both thermodynamically and enzymatically more stable compared with DNA and non-modified RNA.[272] The tetrahedrons were also functionalized to include siRNAs and RNA aptamers, which facilitated siRNA delivery in vitro and in vivo with aptamer-specific tumor targeting.[272] The enhanced stability and targeting properties of RNA structures facilitate carrier-free nucleic acid delivery, which reduces the design complexity and the number of potentially toxic or immunogenic components
While many combinations of nucleic acid and drug cargos have been shown to have additive or synergistic effects, new data regarding the pathways involved in tumor growth, progression, and metastasis should be considered. As our understanding of cancer progression evolves, combination treatments must be continuously optimized to identify potent combinations of anticancer agents. These combinations should act broadly on many types of cancers, particularly in drug-resistant, aggressive, and metastatic disease.
4.6. Multifunctional Targeted Nanocarriers
The use of a cancer-specific cargo reduces the burden of developing a perfectly targeted nanocarrier, since off-target delivery will have minimal effect in normal cells. Further, combinatorial nucleic acid therapies can be selected to address tumor heterogeneity or to target a particular cancer cell subtype. The targeting methods described here work at different levels: local delivery to the tissue of interest, responsive uptake in the tumor microenvironment, active targeting to certain cell types, and genetic targeting to particular expression profiles. Thus, these orthogonal targeting mechanisms can be combined into highly targeted multifunctional nanocarriers. The multifunctional envelope-type nano-device (MEND) developed by Hatakeyama et al. integrates multiple strategies for successful cancer-specific nucleic acid delivery.[203] MEND is a nanocarrier comprised of nucleic acid condensed by a cationic polymer coated with a lipid envelope, which has been functionalized with various combinations of responsive and targeting features. For example, one iteration of the platform combines MMP-cleavable PEG and pH-sensitive fusogenic peptides to overcome the PEG dilemma and specifically enhance cellular uptake of siRNA in solid tumors.[273] Combining transcriptional targeting with particle targeting is another promising dual-targeting strategy, demonstrated by Cocco et al.[274] They developed a PLGA-PBAE blend nanoparticle functionalized with tumor targeting c-CPE peptides and used these particles to deliver a diphtheria toxin subunit A (DT-A) gene under the control of the cancer-specific p16 promoter. The dual-targeted particles efficiently transfected primary patient cells and significantly slowed tumor growth in chemotherapy-resistant ovarian cancer models.
Multifunctional particles also have the potential to achieve stepwise release of therapeutic agents. For example, core-shell particles can release siRNA molecules from an outer polymer layer in a glutathione-responsive manner followed by slow release of a chemotherapy to overcome multi-drug resistance.[275] While multifunctional targeted particles show promise in preclinical studies, the lack of targeted particles in clinical trials for nucleic acid delivery is a testament to the complexity of implementing active targeting in a translational and scalable way. Developing targeted nanocarriers requires additional testing, time, and expense, and the resulting product is often expensive to manufacture. Despite these hurdles, several next-generation actively targeted nanomedicines for nucleic acid delivery have entered clinical trials, as summarized in Table 2. As more potent nucleic acid therapies are developed, particularly immunotherapies, a high degree of specificity will be essential to avoid dangerous off-target effects. Thus, these targeting strategies may serve as enabling technologies to bring nucleic therapies to patients.
Table 2.
Nucleic acid delivery vehicles that have been investigated in clinical trials. SD = Stable Disease, PR = Partial Response by RECIST criteria
| Name | Nanocarrier | Cargo | Targeting | Response | Phase | Identifier |
|---|---|---|---|---|---|---|
| ALN-VSP02 | Liposome | 1:1 anti-KF11 and anti-VEGF | Passive | SD in 6/15 and PR in 1/15 patients at high dose | 1 Completed | NCT01262235 NCT01158079 |
| Atu27 | Liposome | Anti-PKN3 siRNA | Passive | SD in 14/34 patients | 2, Completed | NCT00938574 |
| CALAA-01 | Cyclodextrin NP | Anti-RRM2 siRNA | Transferrin | SD in 1/24 patients | 1, Terminated | NCT00689065 |
| EPHARNA | Liposome | Anti-EPHA2 siRNA | Passive | 1, Recruiting | NCT01591356 | |
| MK-4621–002 | JetPEI™ | MK-4621 RNA targeting RIG-1 | Intratumoral Injection | SD in 4/15 patients | 2, Recruiting | NCT03065023 NCT03739138 |
| NU-0129 | SNA Gold NP | Anti-Bcl2L12 siRNA | Passive | 1, Recruiting | NCT03020017 | |
| PNT2258 | Liposome | DNA oligo blocking BCL2 | Passive | SD in 4/15, PR in 2/15, and CR in 3/15 patients | 2, Completed | NCT01733238 |
| SGT-53 | Liposome | p53 DNA plasmid | Anti-transferrin receptor scFv | SD in 7/11 patients | 2, Recruiting | NCT02354547 NCT02340156 NCT02340117 NCT03554707 |
| TargomiRs | EDV™nanocells | miRNA mimic | Anti-EGFR antibody | SD in 4/6 and PR in 1/6 patients | 1, Completed | NCT02369198 |
| TherGAP | JetPEI™ | Plasmid DNA (DCK::UMK fusion gene) | Local delivery by endoscopic ultrasound | SD in 6/22 patients | 2, Recruiting | NCT01274455 NCT02806687 |
| TKM-080301 | Liposome | Anti-PLK1 siRNA | Passive | SD in 3/6 and PR in 1/6 patients | 1/2, Completed | NCT01262235 NCT02191878 NCT01437007 |
5. Nanomaterials for combinatorial nucleic acid delivery
Nucleic acids confer many advantages as therapeutic agents, one of which is their ability to be co-delivered for combinatorial therapy. Viruses are an obvious vehicle for gene delivery, as they have evolved to be highly efficient at nucleic acid transfer.[276] Presently, the majority of gene therapy clinical trials have relied on viral vectors, and the only FDA-approved gene therapy uses a viral vector.[277] However, while new generations of viral vectors are continually being developed,[278] viruses can have several limitations as delivery vehicles, including safety concerns related to their immunogenicity,[279] risks of mutagenesis,[280] and limited cargo capacity.[281] The last of these, in particular, is an important concern when considering combinatorial delivery of multiple nucleic acid sequences or types. In this section, we focus on non-viral delivery vectors engineered to prevent the problems generally seen with viruses, and their uses for combinatorial nucleic acid delivery will be described. Examples of some prominent types of materials used for this purpose are illustrated in Figure 8.
Figure 8. Types of nanomaterials used for nucleic acid delivery.
Broad classes of materials and nanostructures used as nucleic acid delivery vehicles are summarized, including lipid-based nanoparticles (A), cationic polymer-based nanoparticles (B), nanoparticles based on other polymer types (C), inorganic nanoparticles (D), and nanostructures that use DNA itself as a structural component (E). Part E adapted from Doye et al., “Coarse-graining DNA for simulations of DNA nanotechnology,” Physical Chemistry Chemical Physics 15(47):20381–20772, 2013,[369] with permission from the Royal Society of Chemistry.
5.1. Lipid-based nanoparticles
Lipid-based materials have been studied extensively for delivery of nucleic acids. Easily manufactured, lipid-based nanoparticles can carry many types of nucleic acid within a hydrophilic liposome core, including in combination,[282] and they have been of high interest in particular for the delivery of oligonucleotides like siRNA.[283, 284] Cationic lipids, which can associate with anionic nucleic acids and negatively charged target cell surfaces, can be synthesized using combinatorial chemistry to form a wide range of lipids or lipidoids with varying properties,[285–287] allowing researchers to formulate different nanoparticles in a facile way. Lipid nanoparticles can facilitate cellular uptake by interaction with lipids in the cell membrane, which also aids in endosomal escape,[288–291] or by interacting with endogenous lipoproteins that are taken up via binding to specific cell receptors.[292–294]
However, lipid nanoparticles carry certain drawbacks. The lipid-lipid interactions that help these nanoparticles to enter cells and escape the endosome may also result in high toxicity,[295] and some lipids have been found to be inherently immunogenic on their own, even without the addition of nucleic acid cargo.[296] Thus, careful study of the chemical properties of the lipids used is critical for the use of this type of nanocarrier. Lipid nanoparticles are also prone to interactions with components of the blood that may cause aggregation or destabilization. PEGylation can been used to prevent clearance of lipid nanoparticles,[297] and SNALPs, stable nucleic acid lipid particles, were developed to have both internally stabilizing components as well as a PEG shield.[298] Though challenges remain, lipid-based nanoparticles are among the most highly studied nanocarriers for nucleic acids, both pre-clinically and in clinical trials,[299] especially for delivery of oligonucleotides. Alnylam Pharmaceuticals reported a phase I trial in 2013 that used a lipid nanoparticle carrying two siRNA sequences to target vascular endothelial growth factor (VEGF) and kinesin spindle protein (KSP),[300] with multiple siRNAs being loaded into the same particles. VEGF is upregulated in many solid cancers due to the tumor cells’ need to rapidly increase angiogenesis to provide nutrients to the growing tumor; KSP is involved in mitosis and is overexpressed in quickly dividing cells. In cancer patients with tumors in the liver, the authors reported toxicities comparable to or lower than those seen from traditional chemotherapy and demonstrated anti-tumor activity after administration of their formulation. The biodistribution of these nanoparticles was largely in the spleen and liver, allowing them to passively target hepatic tumors.[300] The authors noted that their therapeutic was detectable in extrahepatic tumors as well and that these tumors were also controlled by the treatment; though an explanation is not provided, it may be that the size of the nanoparticles allowed them to accumulate in these other tumors via the EPR effect. It should be noted that an increase in cytokine levels was noted after treatment,[300] suggesting that, while siRNAs can be modified to elicit minimal immune responses, some immunostimulatory effects may still occur using this and similar systems.
The NOV340 liposomal delivery system, or SMARTICLES®, developed by Mirna Therapeutics Inc. and Marina Biotech, addresses some of the drawbacks of lipid-based nanoparticles.[301] The system uses ionizable lipids that are negatively charged or neutral at neutral pH or higher, allowing the nanoparticles to avoid interaction with cell membranes or tissues that might prevent their free circulation. At lower pH, such as that found at tumor sites, NOV340 becomes cationic, and this triggered change facilitates its interaction with and uptake by cancer cells. A phase I clinical trial using the NOV340 system to deliver a miR-34a mimic to patients with advanced malignancies suggested some anti-tumor activity, though severe adverse immune responses were also observed in some patients.[302] A pre-clinical study using the NOV340 liposomal system to deliver a combination of miRNA mimics (miR-34a and let-7b), showed that the combination of both sequences was more effective than either on its own in a mouse model of non-small cell lung cancer (Figure 9).[303] Thus, while, once again, careful work must be undertaken to better understand the safety profile of lipid-based systems in a clinical setting, combinatorial delivery of nucleic acids still holds promise for therapeutic effect.
Figure 9.
A liposomal formulation (NOV340) facilitated co-delivery of miR-34a and let-7b. After multiple injections (a), tumor lesions in the left lobe of lungs in animals treated with one or both miRNAs were fewer and smaller in number as seen by hematoxylin and eosin staining (b) and quantified in (c). Tumor sizes were lower in animals treated with both miRNAs compared to each individual miRNA (d), and tumor proliferation was lowe after treatment (e-f). Survival was also extended by treatment with miR-34a. Figure reprinted by permission from Springer Nature: Kasinski et al., “A combinatorial microRNA therapeutics approach to suppressing non-small cell lung cancer,” Oncogene 34(27):3547–3555, 2015.[303]
In another illustrative example, Chen et al. used a lipid-based nanoparticle formulation for the co-delivery of miR-34a, an miRNA commonly downregulated in human cancers that can cause apoptosis via multiple mechanisms upon administration to cancer cells, and three siRNA sequences against c-Myc, MDM2, and VEGF, all genes commonly upregulated in solid tumors.[136] Their formulation consisted of hyaluronic acid, miRNA, and siRNA pre-complexed with cationic molecule protamine and then encapsulated into cationic, PEGylated lipids for enhanced stability. The final nanoparticle was surface-modified with the tumor-targeting small chain antibody fragment (scFv) GC4, specific for melanoma. Exploiting the combined effects of multiple RNA cargos, the encapsulation capability of a hydrophilic polymer and cationic agent, the enhanced circulation time conferred by PEG, and the targeting capacity of the scFv, they were able to show not only accumulation of the nanoparticles at the tumor site in a B16-F10 murine melanoma model but also additive inhibition of metastatic tumor growth after delivery of miRNA and siRNAs in combination.[136]
As well as being multiplexed with each other, RNA oligos can also be delivered in combination with traditional chemotherapy drugs. Given a known mechanism of drug resistance by cancer cells, researchers can use siRNA to knock down a gene that promotes resistance and simultaneously deliver a chemotherapeutic. Saad et al. used a cationic liposomal formulation to co-deliver siRNAs against MRP1, an efflux pump that removes drugs from cells, and BCL2, which confers resistance by preventing apoptosis, along with the common anticancer agent doxorubicin.[304] While treatment of lung cancer cells with doxorubicin led to upregulation of both of these resistance-causing genes, co-delivery with either the two siRNA sequences decreased resistance and led to significantly greater susceptibility of the cells to doxorubicin-mediated killing.[304]
As the above examples demonstrate, lipid-based nanoparticles have been, and continue to be, studied extensively for the delivery of nucleic acids, especially siRNA and miRNA. This may be due in part to the greater challenge of binding short oligonucleotides via complexation compared to larger nucleic acids with greater multivalency, as discussed in Section 3; liposomal formulations do not rely on high-avidity electrostatic interactions with nucleic acids in order to encapsulate them in their hydrophilic core and may therefore by favorable for delivery of smaller cargo. This type of formulation is also easily adaptable to the co-encapsulation of multiple cargoes for combinatorial delivery. However, despite the long history of their use in research and in clinical development, lipid-based carriers must still contend with problems of excessive toxicity and potentially lethal side effects described above. Intrinsic immunogenicity as well as enhancement of the immunogenicity of RNA cargo has been of particular concern in the cases described above, including examples that reached the stage of clinical trials.[300, 302]
5.2. Cationic polymeric nanoparticles
As previously mentioned, PLL-based NPs are poorly able to escape the endosome, unlike newer cationic polymers,[305, 306] which contain reversibly protonatable amines that can aid in escape to the cytosol. PEI and its branched or functionalized derivatives, on the other hand, has been a basis for many types of polymer-based intracellular delivery systems in large part because of its reversibly protonated state: PEI, with a mixture of primary, secondary, and sometimes tertiary amines, can exist in a positively charged state that allows tight binding of nucleic acids, while other amines can still buffer additional protons once within the endosome, leading to escape.
An important hurdle to the use of PEI alone as a delivery agent is its toxicity, as the cationic properties that aid in cargo complexation and cellular internalization can also disrupt the cell membrane and lead to adverse effects. High-molecular weight (MW) PEI is effective in gene transfer but tends to be toxic, while low-MW PEI is less toxic but can also be less effective.[307, 308] To address this challenge, researchers have developed PEI-based polymers that have high MW but can degrade into smaller segments.[307] This can be done by linking multiple low-molecular-weight PEI oligomers with ester linkages, which degrade by hydrolysis; reducible disulfide bridges; or other degradable functional groups.[308]
Condensation of nucleic acids into nanoparticles by PEI and similar polymers is driven largely by electrostatic interactions, and some of the properties of such particles can be tuned by changing the ratio of cationic polymer to anionic genetic material. This type of flexibility allows researchers to easily mix multiple different nucleic acids together to co-complex with PEI and form nanoparticles for combination therapy. Chen et al. used branched PEI, which allows the use of relatively low-MW PEI while retaining higher efficacy due to the increased amine density of the branching structure, to co-deliver siRNAs against VEGF receptor 2 (VEGFR2) and epidermal growth factor receptor (EGFR).[309] Both VEGFR2 and EGFR are upregulated in many solid tumors and participate in angiogenesis and tumor growth, respectively. They were able to achieve knockdown of both genes in non-small cell lung cancer cells in vitro and in vivo. Interestingly, knockdown of VEGFR2 alone was more effective in tumor control than knockdown of both genes together; however, the high dosage of VEGFR2 siRNA needed caused adverse side effects, which were mitigated by the combination of both genes at a lower dose. Similarly, the combination of this dual-gene knockdown with a low dose of traditional chemotherapeutic agent cisplatin led to a strong therapeutic effect while minimizing the toxic effects normally seen with higher doses of the drug, illustrating the ability of combination therapy to achieve similar results to a single-agent therapy while reducing toxicity.[309] Although the commercially available branched PEI polymer was not chemically modified in any way for this study, gene knockdown in vitro was demonstrated in the absence of serum, and a therapeutic effect was seen in vivo after local intratumoral injection.[309] As mentioned, unmodified PEI-based polyplexes can be destabilized by interaction with serum proteins and cells after systemic administration, but the methods used by the authors allowed them to sidestep these hurdles, a potentially viable therapeutic strategy in the case of solid tumors that are accessible for local administration of nanoparticles.
Another class of cationic polymers, PBAEs, shares some of the advantages of PEI while also being less toxic due to the presence of degradable ester linkages and has been widely studied for gene delivery.[310] As with lipid-based nanoparticles, combinatorial chemistry has been used to study PBAEs for DNA or RNA delivery,[311–313] generating PBAEs with a variety of different physical and chemical properties. The ease of manufacture of these polymeric formulations makes them a versatile group of materials for gene delivery compatible with large plasmid DNA, RNA oligomers, and smaller CDNs. Aside from their hydrolytically degradable ester linkages, PBAEs can be synthesized with disulfide bridges that target nucleic acid cargo to the cytoplasm.[270, 314, 315]
While PBAEs, like other cationic polymers that form electrostatic complexes, bind to oligonucleotides less efficiently than to larger plasmids due to reduced avidity, the relatively low toxicity of PBAEs permits the use of higher amounts of the polymer or branched polymer structures with greater amine density to compensate for lower binding affinity. Using these strategies, these biomaterials have been shown to be effective for delivery of short sequences like siRNA,[312, 314–317] miRNA,[270] and CDNs.[43] A recent report by Lopez-Bertoni et al. used disulfide-containing PBAEs to co-deliver miR-148a and miR-296–5p, both miRNAs that inhibit the stem cell-like phenotype of human glioblastoma (GBM) cells in an effort to prevent tumor growth in a mouse model.[270] Not only did the PBAE-delivered miRNAs affect the expression of stem cell markers and formation of tumorigenic neurospheres in vitro, but they also inhibited tumor growth in an orthotopic human GBM mouse model. Importantly, the two miRNA mimics in combination were more effective than either of the sequences alone, when measured by tumor size over time, and 67% of the animals treated with the miRNA combination were long-term survivors (>133 days post-tumor implantation, with 67% of these survivors having no detectable tumor), compared to a median survival of <70 days in the control group,[270] demonstrating the power of nanoparticle-based nucleic acid therapies that target multiple biochemical pathways (Figure 10).[270]
Figure 10.
A cationic and bioreducible PBAE was used to form nanoparticles with each of two miRNAs (miR-148a and miR-296–5p) or a combination of both in order to prevent the growth and tumorigenicity of stem-like GBM cells (A). After intratumoral injection of nanoparticles, the combination of both miRNAs was more effective than either individual sequence in reducing tumor size (B) and causing necrosis of tumor tissue (C). Combination miRNA delivery also significantly extended survival (D). Figure adapted with permission from Lopez-Bertoni et al., “Bioreducible Polymeric Nanoparticles Containing Multiplexed Cancer Stem Cell Regulating miRNAs Inhibit Glioblastoma Growth and Prolong Survival,” Nano Letters 18(7):4086–4094.[270] Copyright 2018 American Chemical Society.
Another method of indirectly delivering RNA oligos is to deliver plasmid DNA that encodes short hairpin RNA (shRNA) that will be processed into siRNA once in the cytoplasm. Yin et al. used a disulfide-containing PBAE to co-deliver shRNAs encoding siRNA against survivin and Mdr-1, both important for tumor cell survival and drug resistance.[318] They showed in vitro that PBAE-mediated transfection of drug-resistant human breast cancer cells with either of these shRNAs facilitated increased susceptibility of the cells to doxorubicin treatment. In a mouse model, the combination of both shRNAs with doxorubicin was synergistic and led to slower tumor growth than any of the three components on its own.[318]
Related to PBAEs are poly(amido amine)s (PAMAMs), which contain the more stable amide linkages in place of degradable ester linkages but can be synthesized to contain disulfide bridges that release cargo in the cytoplasm.[319, 320] Branching PAMAM structures, or dendrimers, can be synthesized in a highly controlled manner to have high amine density for complexation with nucleic acids.[321] Like other cationic polymers, PAMAM dendrimers condense anionic genetic material into nanoscale particles, and the properties thereof are dependent in part on the balance of positive and negative charges in the polyplex. Maksimenko et al. used anionic oligomers, including DNA olionucleotides and anionic dextran sulfate, as a binding component in their PAMAM dendrimer nanoparticles, and, by this method, they were able to improve the efficacy of transfection and decrease the required dose of the therapeutic gene.[322] In this case, the co-delivered oligonucleotide was not intended to have a direct biological target but rather served to alter the physicochemical properties of the nanoparticle formulation. This method of stabilizing PAMAM dendrimer formulations with an oligonucleotide was also used by Vincent et al. to form nanoparticles for the co-delivery of two genes, angiostatin and TIMP-2, in the form of DNA plasmids.[323] Both of these genes were intended to inhibit tumor angiogenesis, with the former specifically inhibiting endothelial cell proliferation and the latter regulating the activity of matrix metalloproteinases (MMPs) involved in tissue remodeling and new vessel formation. Using their PAMAM dendrimers complexed with a DNA oligonucleotide and carrying two other DNA plasmids encoding functional genes, they showed that the oligonucleotide enhanced transfection efficacy and that co-delivery of both anti-angiogenic genes was more effective at slowing tumor growth in vivo than delivery of either gene separately.[323]
An advantage of synthetic polymers such as the ones described here is the ease with which they can be tuned or chemically modified to exhibit desirable properties and to bear functional groups of interest, including chemical functionalities as well as biological groups like targeting moieties. For instance, PBAE or PAMAM nanoparticles can be PEGylated to improve biodistribution after administration.[324, 325] The chemical versatility of polymers is a major advantage for increasing functionality or modifying nanoparticle properties, like toxicity. However, unlike liposomal nanocarriers, cationic polymer-based polyplexes do not strictly encapsulate nucleic acids but rather bind them electrostatically, and they therefore bind less strongly to cargos with lower MW (i.e., lower avidity). It is interesting to note that, while both lipid- and polymer-based nanoparticles have been used to deliver many different types of nucleic acids, many of the polymer-based systems tend to be used for larger plasmid DNA molecules rather than siRNA or miRNA, in contrast to lipid-based systems, and/or use various strategies to improve binding, including the addition of non-coding oligomers as binding help[322] and the use of excess cationic polymer to enhance binding stability,[270, 312, 314, 315] the latter of which in turn requires a polymer with low toxicity to allow the increased amount of material to be safely added to the nanoparticle system.
Polymers can also be combined with other classes of materials to take advantage of a wider range of favorable characteristics. Dahlman et al. used lipid-polymer hybrids to achieve in vivo siRNA delivery with specific biodistribution.[326] The authors conjugated a lipid to a low-MW polyamine, and the resulting novel material was formulated with PEGylated lipids and RNA oligos to form lipid-polymer hybrid nanoparticles. Importantly, these hybrids were found to target endothelial cells, rather than the liver,[326] which is the most common target of in vivo siRNA delivery due to the high accumulation of lipid nanoparticles in the liver. In a follow-up study, taking advantage of the ability of a single nanoparticle to carry multiple oligos, the authors delivered an miR-34a mimic, whose expression has been shown to limit tumor growth due to regulation of genes involved in cell cycle progression and apoptosis, and an siRNA against the oncogene Kras in a murine model of lung cancer.[327] This formulation’s material properties allowed it to accumulate in the lungs, and delivery of both miR-34a and siKras simultaneously was more effective at controlling tumor progression than either sequence alone. Additionally, the combination nucleic acid delivery also enhanced the effect of paclitaxel, resulting in significantly longer survival times (Figure 11),[327] further demonstrating the benefits of combining multiple nucleic acids as well as other traditional therapeutic cargos. Designs such as this combine the advantages of multiple systems in an effort to optimize nucleic acid delivery. In this case, the lipid component could interact with cells as well as improving the stability of the nanoparticles; the polymer could be easily modified and used as the basis for reaction with the other materials; and an additional lipid component brought with it PEG functionalization for better in vivo stability, serving as an example of combinatorial use of delivery material as well as cargo.
Figure 11.
Lipid-polymer hybrid nanoparticles were used to co-deliver an miRNA (miR-34a) and siRNA (siKras) to lung tumor. The combination of both decreased the number of cancer cells in vitro (A) and significantly slowed tumor growth in vivo (B). The RNA combination also increased the number of CC3+ apoptotic cells in the tumor (C), and combining miRNA and siRNA with cisplatin further improved animal survival over any of the component treatments alone (D), indicating that combining these modalities can have an additive effect on tumors. Figure reproduced from Xue et al., “Small RNA combination therapy for lung cancer,” Proceedings of the National Academy of Sciences of the United States of America, 111(34):E3553-E3561.[327] Copyright 2014 National Academy of Sciences.
5.3. Other types of polymeric nanoparticles
One of the most commonly studied types of polymers for gene and drug delivery is polyesters,[328] including poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and the co-polymer poly(lactic-co-glycolic acid) (PLGA). As PLGA and related polymers are components of FDA-approved devices, they are generally considered to be safe, a major advantage over other biomaterials that have significant toxicity.[329] However, encapsulation efficiency of nucleic acids in PLGA-based nanoparticles is generally poor compared to the loading efficiency in liposomes or cationic polymers and normally requires additional emulsifying excipients, and nucleic acids have been found to be unstable after encapsulation.[330] To improve association with nucleic acids, cellular uptake, and endosomal escape, a common strategy is to use blends or co-polymers of PLGA with cationic materials, such as PEI,[331, 332] PLL,[333, 334] chitosan,[335] or spermidine,[336] thereby combining the assets of cargo-binding polycations with the biocompatible matrix within which nucleic acids can be encapsulated. For instance, Wang et al. synthesized a co-polymer consisting of a hydrophobic portion based on cholesterol [N-(2-bromoethyl)carbamoyl cholesterol] and a hydrophilic, cationic main chain [poly(N-methyldietheneamine sebacate)] and used the resulting amphiphilic polymer to form core-shell nanoparticles co-encapsulating the chemotherapeutic paclitaxel and either DNA plasmids or siRNA oligos.[337] The authors showed in a 4T1 murine breast cancer model that co-delivery of a DNA plasmid encoding IL-12, an anti-tumor cytokine, could enhance the effects of paclitaxel and slow the growth of the 4T1 tumor. The same nanoparticle system could also be used to co-deliver paclitaxel and siRNA against BCL2 in vitro, knocking down an anti-apoptotic of tumor cells while delivering a cytotoxic agent and demonstrating the ability of the same nanoparticle system to delivery multiple types of nucleic acid due to similarities in their electrostatic properties.[337]
In addition to synthetic polymers, natural polymers like cyclodextrins (CDs) have been used for gene delivery, including in clinical trial.[144] As cyclic oligosaccharides, CDs naturally form an amphiphilic structure that allows them to encapsulate cargo while shielding the cargo from potentially destabilizing interactions with other biomolecules.[338] The encapsulation capacity, the availability of multiple functional groups for chemical modification, and the common use of CDs as excipients and absorption enhancers make CDs a versatile platform for gene delivery. Many efforts to use CDs to deliver nucleic acids use a combination of CDs and other components like PEI[339] or lipids[340] to enhance gene delivery or targeting ligands[341] to prevent off-target effects on healthy tissue.
5.4. Inorganic nanoparticles
Calcium phosphate (CaP) was used in early efforts to achieve DNA delivery by co-precipitating DNA with CaP to form nanoparticles.[342, 343] Once internalized into cells, CaP has been shown to dissolve at the low endosomal pH, which increases the osmotic pressure of the endosomal compartment and leads to cytoplasmic release of nucleic acid cargo.[344] While CaP crystal growth can be challenging to control, CaP nanoparticles have been used in combination with organic biomaterials to improve gene delivery[344–347] to aid in overcoming delivery hurdles like uptake, stability, or endosomal escape. An interesting example is that of lipid-coated CaP nanoparticles (LCPs). In LCPs, nucleic acid is loaded into the CaP core, which is then stabilized with a surrounding lipid bilayer. Yang et al. used LCPs to encapsulate a pool of siRNAs, then coated the core with PEGylated lipids and the tumor-targeting ligand anisamide, which has been used in a number of cancer-targeted nanomedicines.[348] The authors co-delivered siRNAs against HDM2, a suppressor of p53 activity; the oncogene c-myc; and the pro-angiogenic factor VEGF. The PEGylated lipid coating and the targeting ligand together promoted accumulation of the LCPs in non-small-cell lung cancer after injection in vivo. While the authors do not show the anti-tumor effects of each siRNA alone, if any, the pool of three siRNAs in the lipid-modified CaP nanoparticles was able to significantly slow tumor growth and promote cancer cell death.[348]
Mesoporous silica nanoparticles (MSNs) have also been used to facilitate gene and drug delivery due to the biocompatibility of silica and the tunability of the porous architecture of these solid particles.[349] Highly porous MSNs can be loaded with therapeutics like nucleic acids, as well as functionalized for various chemical properties. Taratula et al. developed a multimodal MSN-based delivery system, in which thiol surface-functionalization was used to bind thiol-terminated siRNA to the nanoparticles; surface thiols were also used for PEGylation of the particles; and the porous structure was used to load traditional chemotherapy drugs like cisplatin and doxorubicin.[350] Using inhalation to deliver their loaded and functionalized MSNs directly to lung cancer cells, the authors report that simultaneous knockdown of BCL2 and MRP1, both genes that are important for different mechanisms of drug resistance, led to increased cell death and greater susceptibility to doxorubicin and/or cisplatin.[350] This strategy takes advantage of the various surfaces and compartments available for loading and functionalization on particles like MSNs, which allows different cargoes and materials to be utilized for simultaneous delivery (Figure 12).
Figure 12.
Mesoporous silica nanoparticles allow the co-delivery of multiple cargos and functionalization with a range of materials. In this example, MSNs were functionalized with thiol-reactive groups. The pores were loaded with chemotherapeutics doxorubicin or cisplatin, and the thiol-reactive groups were used to load two siRNA sequences and a tumor-targeting peptide sequence to the surface. Figure reprinted with permission of Taylor & Francis, Ltd, from Taratula et al., “Innovative strategy for treatment of lung cancer: targeted nanotechnology-based inhalation co-delivery of anticancer drugs and siRNA,” Journal of Drug Targeting 19(10):900–914, 2011.[350]
Among inorganic materials, gold nanoparticles (AuNPs) are among the most well studied today. Because their surface can be easily chemically functionalized,[351–353] there are multiple ways in which they can be used to bind nucleic acids. They can be surface-coated layer-by-layer (LbL) with alternating layers of anionic nucleic acids and polycations like PEI[354, 355] or PAMAMs and PBAEs.[74] In one study by Bishop et al., AuNPs were functionalized with anionic 11-mercaptoundecanoic acid (11-MUA), allowing them to be coated with cationic PAMAM or PBAE.[74] These positively charged NPs could then be coated with layers of siRNA or DNA, alternating with layers of cationic polymer. The cationic polymers facilitated cellular uptake and endosomal escape, and reducible disulfide bond-containing PAMAMs were used to cause release of siRNA in the cytoplasm following hydrolytic degradation of PBAEs and release of DNA.[74] AuNPs bring additional advantages, such as their ability to act as sensors due to their optical properties[356] and to deal cause damage to cancer cells via their photothermal properties.[357, 358] Spherical nucleic acids (SNAs) have become prominent in nucleic acid delivery research and are composed of AuNP cores coated with a shell of nucleic acids, using DNA in the earlier studies and later extending to other types of cargo like siRNA.[39] As mentioned briefly above, the properties of SNAs depend heavily on their three-dimensional structure, which affects their interaction with other biomolecules as well as their ability to be taken up by target cells. Later versions of SNAs used a coating of silica between the core and the nucleic acid shell, allowing the AuNP core to be dissolved entirely, leaving a hollow, spherical shell of nucleic acid while minimizing the inorganic delivery vehicle itself.[57]
Like gold, other inorganic nanoparticles can also facilitate simultaneous imaging and therapeutic delivery. Magnetic nanoparticles are easily and scalably manufactured, and they can be detected by MRI[359] for tracking or diagnostic purposes as well as moved toward target cells using a magnetic field.[360] Quantum dots (QDs) are an interesting option as a delivery vehicle with optical properties, as they are designed to be bright, photostable imaging agents and are normally of small size amenable to in vivo trafficking and cellular uptake.[361, 362] As with other inorganic nanoparticles mentioned here, many of these are generally used in combination with other gene delivery materials, such as coatings of PEI,[363] PEG,[364] and other polymers[365, 366] to allow binding and other capabilities required for gene delivery. For example, Dong et al. reported graphene quantum dots (GQDs) conjugated with PEG-PLA for stability, and miRNAs and antisense oligodeoxynucleotides (ASODNs), which have a similar function but greater chemical stability than antisense RNA oligos, were loaded onto the functionalized GQDs via surface absorption.[367] The extremely high surface area of GQDs allowed the co-loading and -delivery of an miRNA inhibitor to suppress miR-21, often upregulated in solid tumors, and an ASODN targeting survivin (Figure 13). The GQDs demonstrated stable optical properties after functionalization, allowing them to be imaged via fluorescence, and they were able to transfect HeLa cells in vitro, with the combination of both inhibitors having a stronger effect on cells than either of them separately.[367] While inorganic nanoparticles have promise and can provide interesting properties to nanocarriers, they have not proceeded to the clinic to the same extent as organic nanoparticles, though strong research on them is ongoing.
Figure 13.
The high surface area of graphene quantum dots allowed them to be functionalized with two probes to inhibit miR-21 and survivin as well as polymers PLA and PEG. The resulting GQDs were biocompatible in addition to retaining favorable optical properties. Figure reprinted with permission from Dong et al., “Multifunctional Poly(l-lactide)–Polyethylene Glycol-Grafted Graphene Quantum Dots for Intracellular MicroRNA Imaging and Combined Specific-Gene-Targeting Agents Delivery for Improved Therapeutics,” ACS Applied Materials and Interfaces, 7(20):11015–11023.[367] Copyright 2015, American Chemical Society.
5.5. DNA-based nanostructures
The base complementarity of DNA allows researchers to form self-assembled structures at the nanoscale[323] with the potential for a high degree of precision and complexity, as reviewed in more detail elsewhere.[368] Oligonucleotides can be self-assembled into simple geometric shapes, branching junction structures, and higher-order periodic structures by selecting the nucleotide sequence to control conformation and binding among strands.[369] While these nanostructures can be well controlled, high purity and precise stoichiometry are required to prevent errors. More complex structures, termed “DNA origami,” can be formed from long single-stranded DNA stabilized with short oligonucleotide sequences. This type of design allows more flexibility in the required stoichiometry as well as greater complexity of the final structure. Importantly, unmodified as well as functionalized DNA nanostructures have been shown to be able to be internalized by cells, normally a challenge for naked DNA, and do not cause the type of cytotoxicity often seen with cationic materials.[368] Nucleic acids can also be easily loaded into DNA-based nanostructures through traditional chemical functionalization or Watson-Crick base complementarity, either as the therapeutic material or as a targeting moiety, as in the case of aptamers. This class of nanomaterials, therefore, can use DNA itself as a structural component as well as facilitating the delivery of biologically active nucleic acids. Lee et al. used self-assembled DNA tetrahedrons to deliver siRNA to HeLa cells.[370] Because tight control of oligonucleotide stoichiometry affords this method a high degree of precision in the geometric shapes formed, they were able to fabricate monodisperse, “molecularly identical” nanoparticles using base hybridization. Pre-conjugation of folic acid to the siRNA strands also allowed the researchers to achieve gene knockdown specifically in cancer cells overexpressing the folate receptor.[370]
The precision and detailed bottom-up planning needed for base hybridization-mediated DNA self-assembly can itself be a hurdle to efficient nanocarrier design, in addition to requiring a large amount of high-purity DNA sequences to use as building blocks. Rolling circle replication (RCR) allows researchers to generate large quantities of long DNA strands via an enzymatic process, and these DNA strands form nanoscale liquid crystals without requiring careful selection of complementary base sequences. This DNA self-assembly method results in “nanoflowers” (NFs) that can be used as a carrier for other therapeutics or even be used as the therapeutic molecule itself, with greater flexibility in the DNA sequences that can be used. Jin et al. fabricated NFs using DNA containing built-in functionalities, including DNAzymes targeting early growth response 1 (EGR-1) and survivin, both involved in the development and progression of breast cancer, and an aptamer binding cancer cells that overexpress nucleolin.[371] By incorporating magnesium pyrophosphate during the self-assembly, they formed NFs that were stable at neutral pH but collapsed at acidic conditions, such as that found in the endosomal component, due to the dissolution of magnesium pyrophosphate. This provided a source of Mg2+ ions to be used as co-factors for the DNAzymes, which subsequently resulted in knockdown of survivin and EGR-1 expression by the cells. Importantly, decreasing the expression of both genes led to better tumor control in vivo than decreased expression of either gene alone. Although in vivo NF administrations were intratumoral rather than systemic, their in vitro studies suggested that their aptamer targeting ligand could potentially lead to specific uptake of NFs by only the cancer cell type of interest.[371]
As with other types of nanoparticles, NFs can be combined with other types of materials to take advantage of their different characteristics. For instance, Zhu et al. designed intertwining DNA-RNA nanocapsules (iDR-NCs) to deliver DNA and RNA together for anti-tumor immunotherapy.[372] They designed DNA with a CpG sequence to stimulate TLR9 and amplified this using RCR; an shRNA sequence targeting Stat3, which has immunosuppressive functions in the presence of tumors, was encoded in DNA and also amplified by rolling circle transcription (RCT) in the same pot to. The resulting DNA and RNA sequences were self-assembled into “microflowers” (MFs), structures on the order of 1–2 μm in diameter, kept fairly large in part because of mutually repelling negative charges. The MFs were then condensed into nanocapsules using a PPT-g-PEG, a cationic polypeptide grafted with PEG chains. Finally, the authors complexed the iDR-NCs with a tumor neoantigen, Adpkg. The resulting particle, consisting of (1) CpG-rich DNA, (2) Stat3-targeting shRNA, (3) a cationic polymer, and (4) a tumor antigen was able affect multiple aspects of the tumor immune environment simultaneously and lead to a durable, specific T cell-mediated anti-tumor response (Figure 14).[372]
Figure 14.
Intertwining DNA-RNA nanocapsules (iDR-NCs) were fabricated into a multimodal delivery vehicle. A CpG-rich DNA sequence and Stat3 shRNA sequence were amplified by rolling circle replication or transcription, respectively, forming microflowers (MFs) (A). The MFs were shrunk, or condensed, into iDR-NCs using PEG grafted to a cationic, hydrophobic polypeptide (PPT-g-PEG), which was also used to load tumor antigens into the NCs (B). These iDR-NCs could then be delivered to APCs in lymph nodes as a vaccine to promote gene knockdown, immunostimulation, and a tumor-specific response (C). Figure is reproduced with permission from Zhu et al., “Intertwining DNA-RNA nanocapsules loaded with tumor neoantigens as synergistic nanovaccines for cancer immunotherapy,” Nature Communications 8:1482 (2017).[372]
5.6. Genome-editing delivery systems
A growing niche in intersection of gene therapy and combinatorial delivery is the field of genome editing. Genome editing technologies [373] are special cases that illustrate the importance of nanoparticles that can reliably co-deliver multiple nucleic acid cargos. A number of genome-editing tools exist, including zinc-finger nucleases (ZFNs)[374] and transcription activator-like effector nucleases (TALENs),[375], both of which consist of the enzyme Fok1 that cleaves DNA at a specific site determined by the DNA-binding domains flanking the endonuclease, and this cleavage event and subsequent repair results in modification of the original sequence. Because the site specificity relies on the binding affinity of the protein domains for a particular DNA sequence, designing new ZFNs and TALENs to target different sequences requires potentially extensive protein engineering.
While these nucleases can be encoded in nucleic acids like DNA or mRNA and delivered as such, we focus here on the CRISPR-Cas [clustered regularly interspaced short palindromic repeat/CRISPR-associated protein (Cas)] platform, which is most commonly accomplished by co-delivery of two separate nucleic acids. In the CRISPR-Cas system, which has been reviewed in detail elsewhere,[373] the DNA sequence specificity is determined by a single guide RNA sequence (sgRNA) rather than by the binding affinity of a protein, simplifying the process of designing systems with new specificities, and a nuclease, commonly Cas9, serves to cut the genomic DNA at the specific site. Traditionally, viral vectors were used to deliver (1) DNA or mRNA encoding the Cas9 nuclease along with (2) DNA or mRNA encoding the sgRNA,[376, 377] but researchers have since developed non-viral methods of co-delivering Cas9- and sgRNA-expressing nucleic acids or sgRNA itself in order to improve the translational potential of this technology.[378]
The delivery of plasmid DNA and mRNA faces challenges similar to those described previously—the former must be trafficked into the nucleus while the latter acts in the cytosol; mRNA also avoids the risk of insertion into the nucleus but may be less stable—but must also contend with the need for both components, Cas9 and sgRNA, to complex with one another before entering the nucleus themselves to edit the genomic DNA. As a result, many groups have developed modified versions of the types of nanoparticles described above to co-deliver the larger Cas9-expressing nucleic acid and the smaller sgRNA. Due to differences in expression kinetics and persistence of the two components, some studies have also found that mRNA encoding Cas9 should ideally be delivered before sgRNA to optimize efficiency.[379] In this study, Miller et al. used zwitterionic amino lipids (ZALs) to co-deliver sgRNA and mRNA encoding Cas9, taking advantage of the ability of lipids with high positive charge to bind to smaller nucleic acids (sgRNA) and the ability of zwitterionic lipids to bind to longer mRNA, resulting in co-encapsulated cargo. Though they noted the preference for sequential delivery of mRNA and sgRNA, concerns about delivery kinetics and persistence must be balanced against the necessity for both nucleic acids to reach the same cell, which is most easily facilitated by delivering both components within the same nanoparticle, and the authors demonstrated successful in vivo editing by increasing the ratio of Cas9 mRNA to sgRNA within the same ZAL particles.[379]
An example of polymer-mediated genome editing was recently reported by Rui et al.[380] In order to co-encapsulate Cas9 plasmid DNA, which is large and must enter the nucleus, and sgRNA, which is much smaller and must complex with the Cas9 protein in the cytosol, the authors used modified PBAEs termed reducible branched ester-amine quadpolymers (rBEAQs). The hyperbranched architecture of rBEAQs increased the amine density of the polymer, permitting tighter electrostatic binding, and the presence of reducible disulfide bridges in the polymer backbone permits targeted cytosolic release of smaller nucleic acids like sgRNA.[380] These PBAE-derivatives were able to achieve approximately 40% gene knockout in vitro by co-delivery of the two cargos.
For successful genome editing using CRISPR/Cas technology, nanoparticles must be designed for successful co-encapsulation and co-delivery of both components, generally achieved using nucleic acid cargos. The differences in the sites of action and the physical and chemical properties of the two components requires that the particles be designed for optimal delivery of both types of cargo. Because this technology is intended to cause permanent genetic change, unlike many of the previously discussed nucleic acid delivery methods that result in only transient retention of the nucleic acid cargo, any non-specific or off-target effects present an important risk. One strategy for addressing this is to design plasmids with promoters allow expression of either Cas or the sgRNA only in certain circumstances, such as in the presence of tetracycline.[381, 382] While this field is still growing, genome editing with nanocarriers has already proven to be a powerful approach and highlights the benefits of being able to co-deliver different nucleic acids.[383]
6. Conclusions and Perspectives
Nucleic acids have the ability to specifically target single genes, affording them high precision, and multiple different nucleic acid cargos can be delivered in combination, providing a strategy for overcoming drug resistance mechanisms in heterogeneous cancer cell populations. The physical and chemical properties of nucleic acids, however, requires that they be carefully complexed or encapsulated in a vehicle that will facilitate safe and effective intracellular delivery to target cells. Various aspects of the nucleic acid sequence, the activity of the nucleic acid or its product, the physicochemical properties of the nanoparticle, or administration method can also determine the tissue or cell type affected.
While there are many different types of nucleic acids, their broadly similar physical and chemical characteristics allows them to be co-loaded or co-complexed into a vehicle for combinatorial delivery, although specific delivery hurdles must still be addressed for different types of nucleic acids. A wide range of different biomaterials can be used to complex or encapsulate genetic material, with cationic lipids and polymers being the most prominent in the field. While a large proportion of the early work in the field of gene delivery focused on development of materials or nanoparticles that could facilitate effective transfection, either in vitro or in vivo, as our understanding of transfection agents improves, as well as our understanding of the biological mechanisms underlying cancer, an increasing number of reports have aimed to achieve enhanced therapeutic effects through combinatorial delivery of genes that target multiple distinct biochemical pathways.
Apart from the challenges of designing the optimal nucleic acids for the desired biological effects and the challenges of engineering the chemical vehicle to overcome delivery limitations in a controlled laboratory setting, for a new translational genetic medicine to be successful, there are also important manufacturing and scale-up considerations. For example, while many bottom-up self-assembly processes allow polymers, lipids, and other biomolecules to bind to and encapsulate nucleic acids into nanostructures, these processes can often be variable and heterogeneous, generating a polydisperse population of nanostructures.[384, 385] This heterogeneity can be exacerbated by scaling up from small scale bulk mixing conditions found in a laboratory to larger size volumes, especially as mixing environments vary with different types of processing equipment. To help alleviate these concerns, increasing attention in the nanomedicine field has turned to continuous manufacture processes such as microfluidic mixing and extrusion techniques that can be more easily scaled in a homogenous and robust manner.[386–388] Similarly, to minimize variability in patient outcomes, homogeneity in nanomedicine properties is critical and may necessitate improvements to both separation and characterization methods to ensure precision. Finally, as these nanostructured materials encapsulate fragile nucleic acids and are often wholly composed of degradable materials, the shelf-life and durability of the materials is also an important concern for clinical translation. Excipient selection, storage conditions, and often freeze-drying procedures (lyophilization) must be optimized and can make the difference between a suitable nanomaterial for genetic medicine in the clinic and one that unsuitable due to aggregation, degradation, or altered biophysical properties.[389, 390]
Thus, for successful nanoparticle-mediated combinatorial nucleic acid delivery to cancer, there are important challenges that span the biological, chemical, and engineering disciplines. Advanced nanomaterials are capable of overcoming these obstacles and progress in this field is accelerating as new approaches enter the clinic. The continuing development of nanomaterials for targeted nucleic acid delivery hold enormous promise for the treatment of cancer.
Acknowledgements
The authors thank the NIH for support (R01CA228133 and R01EB022148).
Biographies
Hannah J. Vaughan received her B.S. in Biomedical Engineering from Duke University and came to the Johns Hopkins University School of Medicine in 2016 as a PhD candidate. Her research under the mentorship of Dr. Jordan Green is focused on targeted non-viral nucleic acid delivery for cancer therapy and diagnostics.
Jordan J. Green is a Professor of Biomedical Engineering, Chemical & Biomolecular Engineering, Materials Science & Engineering, Oncology, Neurosurgery, and Ophthalmology at the Johns Hopkins University School of Medicine. He received his Ph.D. in Biological Engineering from the Massachusetts Institute of Technology in 2007. Prof. Green’s main research interests are in creating biomaterials and nanobiotechnology to engineer cells and developing advanced therapeutics.
Stephany Y. Tzeng earned her PhD in Biomedical Engineering at the Johns Hopkins University School of Medicine in 2014 and joined their research faculty in 2017. Her research focuses on nanomedicine for cancer treatment, with particular focus on gene therapy and immunotherapy.
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