Non-viral Gene Delivery

Sum, Chi Hong; Shortall, Samantha Marisha; Wong, Shirley; Wettig, Shawn David

doi:10.1007/978-3-319-78259-1_2

Chi Hong Sum⁴,
Samantha Marisha Shortall⁴,
Shirley Wong⁴ &
…
Shawn David Wettig⁴

Part of the book series: Experientia Supplementum ((EXS,volume 110))

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1 Citations

Abstract

Although viral vectors comprise the majority of gene delivery vectors, their various safety, production, and other practical concerns have left a research gap to be addressed. The non-viral vector space encompasses a growing variety of physical and chemical methods capable of gene delivery into the nuclei of target cells. Major physical methods described in this chapter are microinjection, electroporation, and ballistic injection, magnetofection, sonoporation, optical transfection, and localized hyperthermia. Major chemical methods described in this chapter are lipofection, polyfection, gold complexation, and carbon-based methods. Combination approaches to improve transfection efficiency or reduce immunological response have shown great promise in expanding the scope of non-viral gene delivery.

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Materials for Gene Delivery Systems

Introduction of Genes via Sonoporation and Electroporation

Principles of Electroporation for Gene Therapy

Keywords

2.1 Introduction

Viral vectors remain the most widely studied gene delivery vector for gene therapy applications, due primarily to the high delivery efficiencies inherent to these vectors; however, the various safety (Raper et al. 2003; Hacein-Bey Abina et al. 2003, 2008) and other, practical, concerns associated with viral vectors continue to drive research into non-viral gene delivery methods. Non-viral methods can be broadly classified into physical- and chemical-based methods. Chemical-based vectors are generally comprised of materials that are substantially less toxic and/or less immunogenic compared to viral vectors and, through the use of targeting ligands, can be targeted to specific cell populations, while physical methods allow for spatial precision (i.e., site-specific delivery of DNA) (Al-Dosari and Gao 2009). Major chemical methods of gene delivery will be described in detail in sections below, along with select physical/combined methods, specifically those that have and continue to generate a significant number of research publications. Other methods, such as electroporation, direct injection, microinjection, and other injection-based systems, remain under investigation, in some cases in ongoing clinical trials; the drawback with all of these methods is the invasive nature of the techniques. This can be balanced by, in some cases, the very high effectiveness of the localized injection of DNA into the target tissue(s) bypassing systemic and intracellular barriers.

2.2 Physical Methods

In the direct injection method, first reported in 1990 (Wolff et al. 1990), DNA uptake occurs primarily in the area of the needle track, pointing to the importance of physical damage to the tissue (Al-Dosari and Gao 2009). The majority of ongoing clinical trials involving direct administration of naked nucleic acid (427 in total as of 2016) are directed to cancer or cardiovascular disease (http://www.wiley.com/legacy/wileychi/genmed/clinical/). While the direct injection method is simple, with no concerns regarding vector toxicity, very low gene expression is observed as a result of the various barriers that impede cellular entry and expression (such as the negatively charged cell membrane, intra- and extracellular endonucleases, aggregation by serum proteins leading to clearance by the reticuloendothelial system, etc.). Microinjection, a technique originally developed for manipulation and transfer of material into living cells (Feramisco et al. 1999), allows for injection of genetic material directly into the nucleus via a micropipette. While microinjection gained a great deal of popularity in the early 1980s, it did not demonstrate efficiencies as high as those observed with viral vectors, is time-intensive (only a certain number of cells can be injected in a given amount of time), and requires a high degree of technical skill on the part of the operator (Cline 1985). Advances in manufacturing capabilities now allow for the production of microneedle patches, which enables dermal and transdermal delivery of drugs in general and DNA (Mellott et al. 2013) in particular [see a number of excellent reviews by Prausnitz (2004) and Kim et al. (2012)]; however, the method still suffers from a number of disadvantages, including primarily localized delivery, as well as imprecise and/or small dose limitations, and a high degree of complexity with respect to the manufacture of the patches themselves (van der Maaden et al. 2012).

Electroporation is perhaps one of the most effective and extensively studied and used methods for non-viral DNA transfection, and Mellott et al. provide a recent and extensive review of the technique (Mellott et al. 2013). The method utilizes electrical pulses to permeabilize cellular membranes (creating pores hence the term electroporation) allowing for the introduction of DNA into the cells and was first introduced by Neumann et al. (1982). While the method has been extremely successful both in vitro and in vivo, in a wide variety of cell lines and/or tissues, the most significant disadvantage is that the method is highly complex, requiring careful determination of physical and biological parameters for each application, as well as a skilled technician for its administration. This, coupled with the invasive nature needed for most applications (electrodes must be in contact with the tissue of interest), has limited its application, particularly in clinical settings. Improvements in the technology has led to ongoing clinical trials beginning in 2004 (Gothelf and Gehl 2010; Gehl 2014; Heller and Heller 2015), which show the high safety, tolerability, and efficacy of gene electrotransfer, particularly to the skin. Recent advances on gene electrotransfer to the skin or muscle are particularly promising, due to the accessibility of the tissues. Donate et al. observed that increased temperature upon electroporation improved DNA uptake, expression duration, and reduced tissue damage as a consequence of lower voltage requirements (Donate et al. 2016).

The final physical method of DNA transfection to be introduced is that of ballistic injection. Ballistic injection (also known as the “gene gun”) differs from direct or microinjection in that it is a needle-free injection method. Instead, plasmid DNA is combined with metal nano- or microparticles (typically gold) and “fired” as a projectile from a pressurized ballistic device (hence the term “gene gun”). The technique was first introduced in 1987 (Sanford et al. 1987) and has gained popularity for applications such as DNA vaccines due to the ability of the technique to readily deliver DNA across the stratum corneum deeper into the epidermis of the skin (Wells 2004), but further application has been limited due to the inability to reach deeper tissues, in vivo, although there are numerous examples of successful delivery to the skeletal muscle, heart, tumor, and embryonic tissues in murine or rat models (see Mellott et al. (2013) and Wells (2004) for reviews on the ballistic method). Additional limitations are inflammation and potential tissue damage and a lack of specificity with both targeted and nontargeted cells being transfected (Mellott et al. 2013).

From the above, it is clear that there is not only great potential for the physical methods identified above, but they have also had excellent success, particularly in in vivo applications. In all cases, the limited clinical application of these methods results from the technical complexity of the method, the invasive nature of the method, or even a combination of both factors. As will be seen in the next section, chemical-based non-viral transfection methods offer numerous solutions to both the technical complexity and the invasive nature of some of the physical methods while still maintaining reasonable or even high levels of transfection efficiency.

2.3 Lipofection and Polyfection

Cationic lipid- and polymer-based carriers make up the most common non-viral gene delivery vectors. The first lipid-mediated delivery of nucleic acid was reported in the 1980s (Fraley et al. 1980; Felgner et al. 1987; Wu and Wu 1987). These systems were and are still considered highly promising gene carriers based on their ability to interact with both the anionic hydrophilic backbone of nucleic acids and with the anionic hydrophobic plasma membrane. Significant progress has since been made in characterizing and improving the interactions between both lipid/DNA (lipoplex) and polymer/DNA (polyplex) complexes and cellular barriers.

While they comprise an incredibly diverse vector space, cationic lipids do share an overall structure, such as (1) a cationic head group, (2) hydrophobic tail domains, and (3) a linker between these groups. Electrostatic interactions between the positively charged lipid head groups or polymer molecules with the negatively charged DNA leads to the self-assembly of lipid-DNA or polymer-DNA complexes in aqueous solution. X-ray studies have identified two key packing geometries: lamellar and hexagonal (May and Ben-Shaul 2004). In a lamellar structure, DNA double-stranded (dsDNA) molecules are sandwiched between lipid bilayers. In a hexagonal structure, the dsDNA molecules are intercalated within a hexagonal or honeycomb-like matrix. These structures may give rise to higher-order structures in solution. Lipid phase behavior has been reviewed more thoroughly elsewhere (May and Ben-Shaul 2004; Ewert et al. 2004). Complex stability is highly dependent on many factors: the lipid structure (Ren et al. 2000; Mintzer and Simanek 2009), the ratio of vector to DNA (Thierry et al. 1997), the structure of the DNA itself (length, circular, or linear), and environmental factors such as temperature, pH, and salt concentration. The structure of the lipid composition heavily influences transfection efficiency as it determines the degree of interaction with the cellular membrane and DNA cargo release. Neutral helper lipids such as dioleoylphosphatidylethanolamine (DOPE) (Boussif et al. 1995) can be added to improve transfection. Enhanced transfection efficiency due to DOPE may be due to its destabilization of lipid membranes, first proposed in 1995 (Farhood et al. 1995). DOPE and similar helper lipids form inverted hexagonal structural domains when complexed with DNA, which is linked with its higher transfection efficiency (Koltover et al. 1998). Investigations (Koltover et al. 1998; Hui et al. 1996; Lin et al. 2003; Mochizuki et al. 2013) suggest that the presence of DOPE or similar helper lipids may increase fusion between the lipid complex and the endosomal compartment thereby facilitating endosomal escape and DNA release.

Two landmark cationic lipids, N-[1-(2,3-dioleyloxy)propyl]-N,N,N-trimethylammonium chloride (DOTMA) (Felgner et al. 1987) and N-[1-(2,3-dioleoyloxy)propyl]-N,N,N-trimethylammonium methyl sulfate (DOTAP) (Stamatatos et al. 1988), have greatly shaped the development of lipofection. DOTMA, DOTAP and their derivatives are now widely commercially available. They share an overall structure: a quaternary ammonium head group and two hydrocarbon tails linked to a glycerol backbone. Structural modifications have been shown to have direct impacts on transfection efficiency; notably, close positioning of the cationic head group to the hydrocarbon chains were observed to be necessary for higher transfection activity (Ren et al. 2000). Another cationic lipid, 2,3-dioleyloxy-N-[2-(sperminecarboxamido)ethyl]-N,N-dimethyl-1-propanaminium trifluoroacetate (DOSPA) (Gebeyehu et al. 1992) became the basis for another widely used commercial transfection reagent, Lipofectamine.

Cationic polymers encompass another large non-viral vector space: positively charged polypeptides capable of DNA complexation and cell membrane interaction. Poly-l-lysine (PLL) was one of the first polymers identified for gene transfer applications (Laemmli 1975). Comprised of long lysine peptide chains, PLL has been investigated in both linear and branched forms (Okuda et al. 2004; Yamagata et al. 2007). Hyperbranched PLL may further provide greater transfection efficiency (Kadlecova et al. 2013) as greater molecular weight has correlated with higher transfection efficiency, though also greater toxicity. As PLL is highly positively charged, PEGylation is useful in reducing its toxicity and increasing stability (Männistö et al. 2002). Another cationic polymer, polyethylenimine (PEI) has been extensively studied since its introduction as a transfection agent in 1995 (Boussif et al. 1995) and is now considered the gold standard for transfection. Available as both a linear and branched polymer, it has a high buffering capacity, which contributes to its ability to escape from the endosomal pathway after cell uptake (Behr 1997).

Many more examples of cationic lipids and polymers have been reviewed in depth elsewhere (Mintzer and Simanek 2009). More recently, hybrid vectors comprised of both polymer and lipid agents (lipopolyplexes) have been investigated (Rezaee et al. 2016). Modified and unmodified PEI has been complexed with DOTMA (Matsumoto), DOTAP (Lee), and DOPE. The potential synergistic interactions between the polymer and lipid agent are being elucidated. Hybrid viral/lipid and viral/polymer vectors (Keswani et al. 2015) may be able to combine the best of both worlds: the highly efficient targeting and cell-penetrating abilities of viral peptides with the safety and low immunogenicity of the non-viral components.

Transfection into cell culture and in vivo using cationic lipids and polymers has generally been impaired by their high toxicity. Due to their highly charged surfaces, cationic lipid and cationic polymeric particles are rapidly cleared by the mononuclear phagocyte system (MPS) or are subject to aggregation by serum components (Senior et al. 1991; Chonn et al. 1992; Zelphati et al. 1998). To combat this, polyethylene glycol (PEG), an amphiphilic molecule, can be conjugated to the nanoparticle in both linear and branched forms. PEG [reviewed in Suk et al. (2016)] has high biocompatibility, low toxicity, and low immunogenicity and is very easy to manipulate, making it an ideal molecule for nanoparticle conjugation (Winkler 2015). Modification with PEG chains, also known as PEGylation, has been shown to help protect against clearance and serum aggregation (Harasym et al. 1995; Meyer et al. 1998; Rejman et al. 2004). This has led to the rise of PEGylated liposomes or “stealth liposomes,” long-circulating liposomes capable of evading the immune system, which are better reviewed by Allen (1994) and more recently by Immordino et al. (2006). PEGylation has also seen much benefit in prolonging the circulation duration of highly cationic polyplexes as well, such as PLL and PEI (Ogris et al. 1999; Mishra et al. 2004). While highly beneficial for extending the serum lifetime of liposomal or polymer vectors, it has been noted that PEGylation may decrease transfection efficiency as it may impair uptake of the vector into targeted cells.

Uptake of most non-viral gene delivery vectors is known to be primarily by endocytosis (Zabner et al. 1995). After charge interaction between the cell surface and the nanoparticle, the particle is engulfed by one of many endocytic pathways. Liposomes have been visualized within the endosomal compartment by electron microscopy (Zhou and Huang 1994; Zabner et al. 1995) and confocal microscopy (Zabner et al. 1995). Pinocytosis and phagocytosis are thought to play a minor role in lipoplex uptake by specialized cells (Matsui et al. 1997). The major bottlenecks to chemical vector-mediated gene transfer once the vector breaches the cell membrane are endosomal escape and bypassing the nuclear membrane. Endosomal escape may be mediated through fusion or penetration of the endosomal membrane using peptides derived from viruses known to escape from the endosome. Such peptides include hemagglutinin (HAA) derived from influenza, TAT derived from HIV, and many others. However, the exposure of such antigenic peptides to serum and how to link these peptides to the vector must then be considered. Alternatively, some lipid and polymer agents with high buffering capacity, most notably PEI, have exhibited effective endosomal escape through the “proton sponge effect” (Behr 1997). Low pH within the endosome leads to increased protonation of the captured lipoplex or polyplex. This is thought to lead to an excessive inflow of ions and water into the endosome causing it to swell and rupture, releasing its contents into the cytoplasm.

The nuclear membrane presents another barrier to transfection in nondividing cells (van der Aa et al. 2006; Lam and Dean 2010). Non-viral gene delivery and expression is far more successful during mitosis when the nuclear membrane has broken down (Tseng et al. 1997). In quiescent cells, transport across the nuclear membrane is regulated by the nuclear pore complex (NPC) (van der Aa et al. 2006). These NPCs shuttle macromolecules (>40 kDa) between the cytoplasm and the nuclear compartment via specific nuclear import and export proteins. Import proteins recognize nuclear localization signals (NLS), typically short clusters of basic residues. A number of both peptide and DNA NLS sequences have been discovered (Escriou et al. 2003; Miller and Dean 2009). The combination of translocation sequences to overcome intracellular barriers with the appropriate cationic carrier to penetrate cellular barriers will be of great use to non-viral gene delivery.

Overall, cationic lipids and polymers comprise the largest category of non-viral gene delivery vectors, with many rapid innovations within the development field. Successful vectors such as PEI have shown great capability of both cell uptake and endosomal escape, while many other polymers and lipids have been far less efficient at navigating cellular barriers. Future innovations will therefore highly depend on elucidating the mechanisms behind these successes and also in improving the ability of non-viral vectors to bypass intercellular barriers.

2.4 Magnetofection

Magnetofection describes nucleic acid delivery mediated by the application of magnetic fields to susceptible vectors (Plank et al. 2000, 2011; Scherer and Plank 2008; Scherer et al. 2002; Mykhaylyk et al. 2007). Although development of magnetically assisted gene delivery stemmed independently by multiple groups after advancements made in magnetic drug targeting as early as the late 1970s (Mykhaylyk et al. 2007), the term “magnetofection” was first coined in by Plank in 2000 (Plank et al. 2000). Early successes paved the way for a plethora of developments both in vitro and in vivo, involving magnetofection; this contributes to the development of numerous commercially available products. As such, magnetofection poses to be a vital research tool in many applications of non-viral gene delivery.

Magnetofection involves the addition of magnetic vectors to cells/tissues followed by the application of permanent (static) or pulsating/oscillating (dynamic) magnets for magnetic field induction (Fig. 2.1). High transfection efficiency associated with magnetofection is mainly attributed to the rapid sedimentation of magnetic vectors onto cell surfaces upon implementation of a magnetic field (Mykhaylyk et al. 2007; Plank et al. 2011; Sauer et al. 2009; Kami et al. 2011; Steitz et al. 2007; Liu et al. 2011a; Song et al. 2010; McBain et al. 2008; Vainauska et al. 2012; Sun et al. 2012; Kamau et al. 2006). The application of magnetic fields improves transfection kinetics by reducing time for transfection from hours to minutes. Rapid sedimentation serves to overcome the diffusion barrier that limits vector accumulation at target cells and, consequently, lowers the required vector concentrations at target tissues (Scherer et al. 2002; Plank et al. 2011; Vainauska et al. 2012). Thus, magnetofection offers several significant advantages: (1) rapid vector internalization, (2) high transfection efficiencies at lower vector doses, (3) lower cytotoxicity, (4) reduced incubation time, and (5) the potential for synchronized transfections. After improved accumulation at the cell membrane, cell uptake of the non-viral vector would then proceed through endocytic routes as is typical for the vector in question (Plank et al. 2011). Cai et al. (2005) examined magnetic carbon nanotubes to mechanically penetrate past cell membranes using a physical means of entry. This will be further examined in the carbon nanomaterials section. Rational vector design remains crucial in attaining high transfection efficiencies. Since non-viral magnetic vectors are assembled in a combinatorial manner, the different components (magnetic core, nanoparticle coating, transfection vector, and nucleic acid cargo) can be individually optimized to tailor to different specific applications (Scherer et al. 2002).

For magnetic nanoparticles (MNPs) to be deemed applicable for magnetofection, they must possess four key characteristics: (1) the ability to interact with nucleic acid or with nucleic acid vector, (2) sufficient magnetic properties to drive rapid sedimentation by affordable and relatively mobile magnetic devices, (3) high stability for long-term storage, and (4) biocompatibility for application in living cells/tissues (Mykhaylyk et al. 2007). Typically, superparamagnetic iron oxide nanoparticles (SPIONs) are employed as the magnetic core due to their superparamagnetic properties, high biocompatibility, low toxicity, biodegradability in vivo, and concurrent applications in magnetic resonance imaging (MRI) (Plank et al. 2011; Kami et al. 2011). SPIONs are typically coated by different materials including carbohydrates (dextran, chitosan, heparin sulfate), synthetic polymers [polyethylenimine (PEI), polyethylene glycol (PEG), dendrimers, poly(lactic-co-glycolic acid) (PLGA)], proteins (albumin, streptavidin), cationic cell-penetrating peptides (e.g., TAT peptide), surfactants, phospholipids, silica, and gold (Kami et al. 2011; Wahajuddin 2012). The surface coating reduces MNP aggregation, prevents oxidation, lowers cytotoxicity, and enables functionalization of MNPs. SPIONs are commonly coated with PEI, well known as a gene carrier with high transfection efficiency but also associated with high cytotoxicity. PEI-coated MNPs are commonly used in transfection studies as evidenced by several commercially available products (transMAG^PEI, PolyMAG, CombiMAG). Lipid-coated MNPs, or magnetoliposomes, can be generated by encasing the magnetic core with a cationic shell composed of various mixtures of cationic and neutral lipids/phospholipids (Yang et al. 2008; Namiki et al. 2009).

The formation of non-viral magnetic vectors involves the association of MNPs with lipid- or polymer-based transfection vector, predominantly through electrostatic interaction (Mykhaylyk et al. 2007). As MNPs are commonly coated with materials pertinent to transfection (e.g., MNP^PEI and cationic lipids), magnetic vectors can be assembled through direct interaction between coated MNPs and nucleic acid. By themselves, MNP/PEI/DNA or MNP^PEI/DNA magnetoplexes and magnetoliposomes can self-assemble into effective magnetic vectors (Fig. 2.2); however, additional components are usually incorporated to assist with the transfection process. The incorporation of PEG is advantageous in in vivo applications as it reduces serum protein interactions, thereby prolonging vector half-life; however, addition of PEG has been shown to dampen gene delivery possibly due to interference with intracellular trafficking (Cool et al. 2013). The incorporation of TAT peptides permits efficient intracellular trafficking resulting in further enhancements to transfection efficiencies. Subsequent fusion of histidine residues to TAT peptide permits effective endosomal escape through the proton sponge effect as histidine residues adsorb protons within the acidic environments of endolysosomes. Using histidine-rich TAT peptides, Song et al. (2010) demonstrate in vitro gene delivery of self-assembled MNP^PEI(PolyMAG)/DNA/TAT (histidine-rich) ternary complexes into human astrocytoma cells, yielding transfection efficiencies as high as 60% at 12:1 TAT/DNA ratio. Subsequent in vivo delivery of the ternary complexes into a rat spinal cord by lumbar injection resulted in a twofold increase in transgene expression when compared to MNP^PEI(PolyMAG)/DNA (Song et al. 2010). As cellular uptake of magnetic vectors is predominantly through endocytosis (Scherer et al. 2002; Plank et al. 2011; Sauer et al. 2009), functionalizing magnetic vectors with ligands allows for targeted delivery into specific cells. These ligands can be incorporated in several different methods: (1) electrostatic interactions, (2) conjugation to a preexisting component to be incorporated into the vector (e.g., PEG), or (3) direct conjugation to MNPs by a biotin-streptavidin bridge. Surface characteristics of magnetic vectors heavily influence serum protein interactions, aggregation behavior, cellular uptake, and intracellular trafficking (Fig. 2.3) (Arsianti et al. 2010). Since cellular uptake of magnetic vectors is mediated by endocytosis, the size of the vector plays an important factor in transfection efficiency as it affects the extent of uptake and intracellular trafficking (Plank et al. 2011; Steitz et al. 2007; Park et al. 2011). Both size and surface characteristics must be taken into consideration in order to achieve high transfection efficiencies. Therefore, controlling the ratios of each magnetic vector component must be optimized.

Since magnetic vectors are assembled in a combinatorial manner, formation of lipoplexes and polyplexes can precede the addition of MNPs during vector assembly. By mixing the order of the different vector components, the vector’s biophysical properties are affected, thus impacting transfection efficiency. With respect to PEI, DNA, and MNP^PEI compositions, vectors with surface PEI (positive zeta potential) achieve effective gene delivery through rapid cellular uptake and release from endolysosomes; on the other hand, vectors with surface DNA (negative zeta potential) have inefficient gene delivery due to inadequate cellular uptake and endolysosomal release (Arsianti et al. 2010). Formation of lipoplex/polyplex prior to MNP addition may ensure optimal surface characteristics for better transfection efficiency. A study by Ma et al. (2011) indicates that only free PEI/pDNA polyplexes may enter the nucleus, while MNP^PEI remains within the cytoplasm. In addition, transfection of SPC-A1 and COS-7 cells is hindered by the absence of free PEI/pDNA polyplexes; this implicates the requirement of PEI/DNA polyplexes to dissociate from MNP for nuclear entry (Ma et al. 2011). Therefore, it may be beneficial to form lipoplexes/polyplexes prior to MNP addition during vector assembly. Overall, component selection and assembly method are both critical to the design of a magnetic vector.

Early literature results suggested that magnetofection has no influence on the uptake and subsequent intracellular trafficking of magnetic vectors. Further analysis by Sauer et al. (2009) confirmed this with single-particle analysis; assessing the intracellular fate of magnetic lipoplexes indicated the same three distinct phases present in both magnetic and nonmagnetic PEI/DNA polyplexes. Upon rapid sedimentation, Sauer et al. observed (1) cellular uptake by endocytosis, (2) confined diffusion in cytoplasm, and (3) microtubule-mediated active transport of vesicle-bound magnetic lipoplexes. The observation of a perinuclear ring confirms the intracellular trafficking of the vesicle-bound magnetic lipoplexes into the perinuclear region. However, endosomal escape and MNP dissociation remained problematic as the majority of magnetic lipoplexes remained vesicle-bound after 24 h with 90% of the complexes showing no separation between pDNA and MNPs (Sauer et al. 2009). In contrast, Xie et al. (2015) more recently examined transfection of HepG2 cells with magnetic PEI/DNA complexes (MPD-cc) using fluorescent imaging. Although transfection efficiency was overall low (11.3%), transfection with application of a magnetic field was far superior than without (0.7%). More importantly, visualization of Cy5-labeled DNA and FITC-labeled PEI with confocal microscopy demonstrated destruction of the endolysosomal compartment after 24 h, indicating improvement in endosome escape using magnetofection. Xie et al. concluded the increased transfection efficiency could be attributed to lower protein adsorption when within a magnetic field (Xie et al. 2015). These results in contrast to previous results indicate that more factors could be in play, including target cell type, cargo DNA type, degree of magnetic field, and the composition of the magnetic complexes.

Magnetic particle size, composition and concentration, magnetic field strength, and exposure times are key parameters that need to be optimized for enhanced transfection efficiency (Stride et al. 2009). Magnetofection generally involves the application of a permanent (static) magnetic field with a field strength ranging from 70 to 250 mT and a field gradient ranging from 50 to 130 T m⁻¹ (Mykhaylyk et al. 2007). It has been proposed that pulsating/oscillating (dynamic) magnetic fields introduce lateral, rolling motion of the magnetic vectors upon rapid sedimentation, and such motions may prevent large aggregate formation, enable adequate distribution of sedimenting magnetic vectors, induce receptor clustering, and apply mechanical stimuli facilitating endocytosis (McBain et al. 2008; Vainauska et al. 2012; Kamau et al. 2006). At optimal conditions, the application of pulsating/oscillating magnetic fields, alone or in combination with permanent magnetic fields, has shown to improve transfection efficiencies when compared to magnetofection solely by permanent magnetic fields (Steitz et al. 2007; McBain et al. 2008; Vainauska et al. 2012; Kamau et al. 2006).

Alternatively, magnetofection has been conducted in conjunction with magnetic cell sorting through a process referred to as “magselectofection” (Fig. 2.4) (Plank et al. 2011). A high-gradient magnetic field cell separation column is used as the magnetic device for simultaneous cell sorting and magnetofection. Magnetic vectors and a mixed cell population, composed of both magnetically labeled and unlabeled cells, are added and passed through the separation column. The induced magnetic fields promote close interactions between magnetic vectors and magnetically labeled cells which results in high transfection efficiencies (Plank et al. 2011). Cell separation is achieved as magnetically labeled cells remain within the column while non-labeled cells pass through the column.

Studies have indicated that, in the presence of a magnetic field, MNP/PEI/DNA- or MNP^PEI/DNA-based magnetoplexes can induce rapid internalization and enhance transfection, yielding efficiencies greater than transfections mediated by Lipofectamine (Song et al. 2010; Vainauska et al. 2012; Miao et al. 2013; Al-Deen et al. 2011; Ensenauer et al. 2011; Kong et al. 2012a; Wang et al. 2011) and PEI (Park et al. 2011; Miao et al. 2013; Namgung et al. 2010; Prijic et al. 2012) while exhibiting lower cytotoxicity than PEI (Steitz et al. 2007; Song et al. 2010; Sun et al. 2012; Park et al. 2011; Al-Deen et al. 2011; Namgung et al. 2010). As expected, transfection efficiency is also shown to be enhanced by the implementation of magnetoliposomes over traditional liposomes through magnetofection (Yang et al. 2008; Prijic et al. 2012). For traditional transfection agents (e.g., PEI and cationic lipids), higher transfection efficiencies are generally associated with greater cytotoxicity from increasing concentrations. Higher concentrations are required such that a sufficient N/P ratio (molar ratio of the complexing cationic polymers to the anionic phosphate groups of the nucleic acid) could be attained for effective compaction and encapsulation of nucleic acid. With respect to MNP-PEI-DNA-based magnetoplexes, and magnetic vectors in general, lower cytotoxicity is mainly attributed to effective internalization induced by rapid sedimentation. Through magnetofection, lower nucleic acid dose is required for effective transfection efficiencies; this corresponds to lower concentrations of the transfection agent and lower cytotoxicity. The benefits of magnetofection are clearly highlighted by enhanced transfection efficiencies in the presence of serum proteins as rapid sedimentation limited nonspecific interactions between magnetic vectors and serum proteins (Liu et al. 2011a; Sun et al. 2012; Al-Deen et al. 2011; Ensenauer et al. 2011). In the presence of serum proteins, interactions between negatively charged serum protein and surface PEI induce surface charge reversal and disruption of the magnetic vector (Arsianti et al. 2010). Nonspecific serum protein interactions present an important roadblock for successful translation of in vitro findings to in vivo settings. This imposes significant implications to in vivo applications of magnetic vectors.

Magnetofections have been employed in various different applications of gene delivery in vitro including transfection of primary cells (epithelial, endothelial, cardiomyocytes, myoblasts, chondrocytes, fibroblast, synoviocytes, osteoblasts, melanocytes, and mouse embryonic stem cells) (Plank et al. 2011), transfection of neural precursor/stem cells (Sapet et al. 2011; Jenkins et al. 2011; Pickard et al. 2011; Fallini et al. 2010), and induced pluripotent stem cell generation (Park et al. 2012). Magnetofections have been applied to in vivo applications, particularly in cancer gene therapy due to tumors’ leaky vasculature permitting both passive targeting, by enhanced permeability and retention effect, and active targeting by magnetic force and ligand-receptor interactions (Plank et al. 2011). Magnetofection studies involving the delivery of plasmids coding for tumor necrosis factor-related apoptosis ligand (TRAIL) (Miao et al. 2013) and murine interleukin 12 (IL-12) (Prijic et al. 2012) both report significant tumor reductions and antitumor effect in mice. Similarly, tumor growth rate in mice has been significantly lowered as a result of downregulated expression of type 1 insulin-like growth factor receptor (IGF-1R) through magnetofection of IGF-1R siRNA (Kong et al. 2012a). Through magnetofection, Zhou et al. (2007) were able to demonstrate enhanced immune response to HIV-DNA vaccine in mice through the delivery of HIV-gag encoding magnetic vectors by tibialis anterior injections and subsequent magnetofection. Using magnetofection, immune response is activated at low doses corresponding to 1000-fold increase in DNA vaccine potency. A single immunization by magnetofection was able to elicit 100% seroconversion after 1 week in contrast to 5 weeks after three separate immunizations of naked DNA; this highlighted the ability of magnetofection to accelerate humoral immune response. Subsequent in vitro analysis confirms the potential of cytotoxic T-lymphocyte response upon immunization (Zhou et al. 2007). In addition to enhancing transfection efficiency by magnetofection, the application of an external magnetic field allows for localized and guided targeting in regions specifically exposed to the magnetic field. This feature was highlighted by Song et al. (2010) where moving the external magnetic field down the spinal cord changed the regional distribution pattern of transgene expression.

Magnetofection presents a highly promising avenue of increased vector accumulation to targeted cells. However, despite the advantages of magnetofection in vitro and in vivo, the technique is still limited by the low penetration depth of the external magnetic field, high field strengths required to overcome hydrodynamic forces for MNP localization, and insufficient biodistribution analysis demonstrating the effects of magnetic targeting (Plank et al. 2011; Owen et al. 2012). In order to truly realize the benefits of in vivo magnetofection, these obstacles must be overcome for deeper tissue penetration.

2.5 Sonoporation: Nano-/Microbubbles and Bubble Liposomes

Ultrasound consists of acoustic waves that propagate at frequencies beyond the human auditory range. For medical applications, ultrasound has been commonly used for ultrasonography, a noninvasive diagnostic medical imaging technique capable of visualizing internal structures and organs in real time. In comparison to other imaging techniques—computed tomography (CT), positron emission tomography (PET), and magnetic resonance imaging (MRI)—ultrasound is advantageous in terms of its low costs, portability, ease of use, and suitability for repetitive administrations in the absence of ionizing radiation. The low power levels involved in ultrasound imaging limit the risk of any significant adverse effects on tissues. Conventional ultrasound imaging employs sound wave pulses, transmitted from a probe (transducer) at a specific frequency. These sound waves become reflected and scattered as they pass through tissues. Scattering is caused by variations in density and compressibility of the different tissue components; this results in contrast and the generation of an ultrasound image (Alzaraa et al. 2012). Images can be refined by the addition of contrast agents that serve to increase scattering and reflection of ultrasound signals (Böhmer et al. 2009).

In addition to diagnostic imaging, ultrasound can be applied in therapeutics through thermal and nonthermal (pressure) effects associated with sound waves. Thermal effects can be induced upon the application of high-intensity focused ultrasound (HIFU) for the absorption of acoustic energy by fluids and tissues, resulting in an increase in localized temperature and tissue damage (Alzaraa et al. 2012; Nomikou and McHale 2012; Husseini and Pitt 2008; Feril and Tachibana 2012). Thermal effects associated with HIFU are the basis of hyperthermic ablation therapy, a type of cancer therapeutic administered in conjunction with systemic chemotherapy and radiotherapy (Alzaraa et al. 2012; Nomikou and McHale 2012; Sanches et al. 2011). Temporal induction of localized hyperthermia has also been used in drug/gene delivery by means of temperature controlled release and subsequent induction of elevated transgene expression (Böhmer et al. 2009; Feril and Tachibana 2012). Nonthermal effects of ultrasound are associated with the acoustic pressure waves generated from low-intensity ultrasound, and it is the basis of thrombolysis therapy (Kiessling et al. 2012) and sonoporation.

Sonoporation is a physical method of non-viral gene delivery through modifications to the endothelial layer and the cell membrane permeability as a result of the application of low-intensity ultrasound (Fig. 2.7); thus, sonoporation facilitates gene delivery by enhancing extravasation and internalization. Transient pore formation and upregulation of endocytosis are speculated to be the key mechanisms of gene delivery by sonoporation (Delalande et al. 2012). Transient pore formation is proposed to occur by several different ways: (1) mechanical stimulation (pushing and pulling effects on cell membrane), (2) shear stress and microstreams, (3) microjets/jetting, (4) sonophore formation, and (5) microbubble translation (Delalande et al. 2012). Pore formation may permit direct entry of nucleic acids into the cytoplasm and has been postulated to cause transient intracellular Ca⁺² influxes that result in upregulation of endocytosis (Delalande et al. 2012). Other associated effects resulting from transient pore formation include increases in membrane permeability by mechanical stimulation and reactive oxygen species as well as increases in membrane fluidity by minor elevations in temperature (Alzaraa et al. 2012; Husseini and Pitt 2008; Feril and Tachibana 2012; Kiessling et al. 2012; Delalande et al. 2012; Yoon and Park 2010).

At very low ultrasound intensities, the acoustic pressures generated by ultrasound waves induce certain membrane effects that may lead to an increase in permeability and pore formation. The bilayer sonophore model, introduced by Krasovitski et al. (2011), suggests that cell membranes are capable of converting acoustic pressure wave into mechanical forces that act to cause expansions and compressions of the intramembrane space; transient pore formation can result from the growing tension on the membrane leaflets. Sonophores have been shown to arise from membranes of subcellular compartments (Krasovitski et al. 2011). However, sonophores alone may not be able to cause sufficient membrane permeability/disruption required for effective gene delivery. The addition of microbubbles, contrast agents used in ultrasound imaging, effectively introduces and amplifies linear/nonlinear oscillations for effective ultrasound-mediated gene delivery. Microbubbles have been observed to undergo oscillations causing local deformation and transient pore formation at low ultrasound intensities (Moosavi Nejad et al. 2011). Through observations with fluorescently labeling microbubbles, Delalande et al. (2011) proposed a new mechanism involving the direct translation of microbubbles across cell membranes; microbubbles have been shown to enter cells and rapidly dissolve, losing a portion of its shell during the process.

The extent of microbubble oscillations is under the influence of applied acoustic pressures, which is proportional to ultrasound intensity (Fig. 2.5) (Nomikou and McHale 2012). At higher acoustic pressures, microbubbles introduce and amplify nonlinear oscillations that give rise to cavitation. Depending on the ultrasound intensity, cavitation is subdivided into two categories: stable and inertial cavitation (Fig. 2.6). Stable cavitation occurs at intensities where continuous compression and expansion of microbubbles give rise to nonlinear oscillations leading to the formation of microstreams. Microstreams in the surrounding region induce shear stress that creates transient pores, increases membrane permeability, and upregulates pathways associated with stress response (Cool et al. 2013; Nomikou and McHale 2012; Husseini and Pitt 2008; Delalande et al. 2012; Suzuki et al. 2011; Qiu et al. 2012). The microbubbles at this point remain stable and grow in size by accumulating dissolved gases. They continue to grow until they reach resonant size, resulting in stable oscillations (Husseini and Pitt 2008; Suzuki et al. 2011). At higher intensities, forced expansion and compression of microbubbles cause bubbles to collapse and inertial cavitation. Inertial cavitation creates substantial nonlinear oscillations and leads to the formation of shockwaves and microjets that give rise to transient pores on cell membranes (Cool et al. 2013; Nomikou and McHale 2012; Husseini and Pitt 2008; Suzuki et al. 2011; Qiu et al. 2012). Cell membrane deformations caused by oscillating microbubbles has been previously captured by ultrafast camera; thus, the process of transient pore formation by stable and inertial cavitation has been demonstrated (Böhmer et al. 2009). Studies by Qiu et al. (2010, 2012) have indicated the influence of inertial cavitation in sonoporation from evidence indicating that there is a high correlation between inertial cavitation, transfection efficiency, and cell toxicity. Initially, inertial cavitation is linearly correlated with transfection efficiency but becomes saturated at a certain point; this is an indication of excessive inertial cavitation leading to permanent membrane damage and cell death. Increasing acoustic pressure or duration has been correlated to increases in pore size and higher transfection at a cost of increasing cytotoxicity (Qiu et al. 2012). Hence, ultrasound parameters need to be optimized such that sufficient ultrasound exposure is applied to enhance transfection efficiency without compromising cell viability.

Frequency, intensity, duration, microbubble composition, microbubble and nucleic acid concentrations, and cell type are key parameters associated with ultrasound-mediated gene delivery (Fig. 2.7) (Delalande et al. 2012; Yoon and Park 2010; Suzuki et al. 2011; Figueiredo and Esenaliev 2012). Generally, high ultrasound frequencies, ranging from 3 to 10 MHz, are used for clinical imaging, while lower ultrasound frequencies, ranging from 1 to 2 MHz, are used for conventional therapeutic ultrasound (Yoon and Park 2010). The most optimal frequency for gene delivery is dependent on optimal microbubble interactions and propagation path. The highest level of oscillation occurs when the applied frequency matches the resonant frequency of the microbubble (Böhmer et al. 2009; Husseini and Pitt 2008). The low frequencies are generally employed for ultrasound-mediated gene delivery because greater tissue penetration is attained with lower frequencies. It has been shown that microbubbles oscillate steadily, with maximized tissue penetration at approximately 1 MHz (Nomikou and McHale 2012; Figueiredo and Esenaliev 2012). Ultrasound intensity (W/cm²) is proportional to acoustic pressure squared; since higher intensity constitutes higher energy deposition, high-intensity ultrasounds (>5 W/cm²) are suitable to induce thermal effects, while low-intensity ultrasounds (0.125–3 W/cm²) are suitable to induce nonthermal effects (sonoporation) (Yoon and Park 2010). The mechanical index (MI), a measure of the likelihood of inertial cavitation, is defined as \( \mathrm{MI}=\frac{P^{-}}{\sqrt{f\ \left(\mathrm{MHz}\right)}} \) where P⁻ is the peak rare-fractional pressure and f is the frequency (Husseini and Pitt 2008; Yoon and Park 2010). The MI must be carefully noted as it provides an indication of microbubble activity upon ultrasound exposure: MI < 0.05–0.1 (linear oscillations), 0.1 < MI < 0.3 (stable cavitation), MI > 0.3–0.6 (inertial cavitation), and MI > 1 (detrimental biological effects) (Husseini and Pitt 2008; Suzuki et al. 2011). Correlations between studies are difficult due to the variations in each parameter. Numerous in vivo studies have been conducted using a wide range of values for each parameter: frequency (1–4 MHz), intensity (0.5–5 W/cm²), pulse mode (10%—continuous wave), and stimulation time (10 s–30 min) (Delalande et al. 2012). Varying responses to microbubble composition significantly compound issues relating to cross-study comparisons, as a large number of different microbubble formulations have been presented in literature, each with its own inherent distinct properties. This is further compounded by variations in cell-type responses to different microbubbles and ultrasound conditions.

Microbubbles, with sizes ranging from 1 to 10 μm, can be classified into four different groups: coated microbubbles, multilayered microbubbles, phase-shift emulsion, and echogenic liposome. Coated microbubbles consist of a biodegradable shell coating, which in itself may be composed of surfactant, protein, biopolymer, galactose, and/or lipids (Alzaraa et al. 2012), encapsulating gases with air or gases. Coating composition and coating consistency determines the stability of the coated microbubble as well as its sensitivity to pressure changes induced by ultrasound exposure (Alzaraa et al. 2012; Yoon and Park 2010; Nomikou et al. 2012). First-generation coated microbubbles are composed of air encased with surfactant, galactose, or denatured albumin coating; they are ineffective due to their short half-life and high MI (Alzaraa et al. 2012). Second-generation coated microbubbles, composed of a stronger coating and hydrophobic heavy inert gases (perfluorocarbon, perfluoropropane, or sulfur hexafluoride), have greater stability (longer half-life) and can reach the main arteries by crossing the pulmonary vascular bed (Alzaraa et al. 2012; Suzuki et al. 2011). Optison (Molecular Biosystems Inc.), Sonovue (Bracco Diagnostics), and Definity (Bristol-Myers Squibb Medical Imaging) are three second-generation microbubbles approved by the FDA for human use. The aforementioned microbubbles are thin-shelled microbubbles (2–3 nm thick) that require hydrophobic fluorinated gases to prevent gas dissolution for better stability; on the other hand, polymer microbubbles are thick-shelled microbubbles (20–100 nm thick) that do not require hydrophobic fluorinated gas for stability but do require higher acoustic pressures for oscillation and cavitation (Böhmer et al. 2009; Sanches et al. 2011; Delalande et al. 2012). Multilayered microbubbles serve to increase the effective surface area of the microbubble for improved functionalization. Phase-shift emulsion consists of stabilized emulsions of volatile liquid droplets that become microbubbles upon injection and subsequent ultrasound exposure (Yoon and Park 2010). Echogenic liposomes are gas containing-liposomes composed of a phospholipid bilayer (Huang 2008).

Ultrasound-mediated gene delivery by sonoporation can be accomplished by co-administration of microbubbles and nucleic acid as separate entities (Nomikou and McHale 2012). As sonoporation induces transient pores that allow direct entry of nucleic acids into the cytoplasm, nucleic acids are not required to be packaged into a transfection vector. However, high nucleic acid concentrations are required to elicit therapeutic effect due to rapid degradation of naked nucleic acid (Cool et al. 2013). Hence, packaging nucleic acid into a transfection vector is still warranted for encapsulation, condensation, and protection. Alternatively, the attachment of nucleic acid to microbubble protects the nucleic acid from degradation, improves circulation profile, and increases delivery upon ultrasound exposure (Sirsi et al. 2012). Regardless, co-administration of microbubbles and nucleic acids/transfection vectors (lipoplex or polyplex) constitutes low in vivo transfection efficiency due to differing pharmacokinetic profiles of microbubbles and vectors that render in vivo distribution of both components difficult to control (Un et al. 2010). As a result of this, nucleic acid and/or transfection vectors are generally attached to microbubbles for transport to specific sites, followed by release and delivery upon microbubble cavitation and transient pore formation. There are numerous methods of incorporating nucleic acid cargo into microbubbles: (1) direct physical incorporation into microbubble shell during fabrication, (2) electrostatic interaction with cationic lipids incorporated into microbubble shell, (3) formation of a multilayered microbubble with a surface layer of cationic polymer for nucleic acid binding, (4) tethering transfection vectors by hybridization of complementary DNA oligonucleotides, (5) covalently linking transfection vectors to microbubbles, and (6) linking transfection vectors to microbubbles through a strong non-covalent biotin-avidin bridge (Fig. 2.8) (Cool et al. 2013; Sirsi et al. 2012). Similarly, ligands can also be linked to microbubbles with biotin-avidin linkage and peptide bonds or become incorporated during microbubble synthesis (Böhmer et al. 2009; Kiessling et al. 2012). The addition of ligands permits active targeting to specific sites, which enhances site specificity in conjunction with ultrasound-mediated delivery. Thus, lower toxicities can be attained from the reduction in transfection vectors necessary for enhanced transfection efficiency.

Sonoporation has been employed for in vivo gene delivery into various cells/tissues (endothelium, skeletal muscles, peritoneal macrophages, salivary gland, ocular ciliary muscle, myocardium, skin, carotid artery, and embryonic limb bud mesenchyme/ectoderm) (Nomikou and McHale 2012; Suzuki et al. 2011). siRNA delivery, limited by their inability to passively diffuse through cellular membranes, has been achieved through sonoporation in numerous documented studies (Cool et al. 2013; Böhmer et al. 2009; Yoon and Park 2010; Suzuki et al. 2011; Vandenbroucke et al. 2008). Similarly, shRNA has been documented to be delivered by sonoporation; this permits stable gene expression and prolong RNA interference (Chen et al. 2009a, b, 2010). In vivo studies in rabbits have demonstrated transient pore formation at the blood-brain barrier that persists for up to 6 h (Böhmer et al. 2009).

Focused ultrasound has been successfully employed in the noninvasive and reversible disruption of the blood-brain barrier (BBB), permitting drug or gene delivery to the brain (Timbie et al. 2015). MRI-guided ultrasound at low-pressure amplitude and low acoustic power has been shown to mediate transient pore formation at the BBB. Particles as large as 100 nm can then bypass the permeabilized BBB and circulate in the brain (Timbie et al. 2015). Nance et al. (2014) constructed stealth nanoparticles (PEGylated polystyrene particles) which they termed “brain-penetrating nanoparticles” (BPN) for delivery to the murine brain both ex vivo and in vivo. More recently, they demonstrated successful delivery of GDNF-expressing plasmid using BPN, eliciting robust and sustained gene expression over 10 weeks post treatment (Mead et al. 2017). These findings have serious implications on drug and gene delivery to the brain by focused ultrasound. Effective image-guided spatial and temporal control over ultrasound, by optimizing ultrasound parameters, enables site-specific gene delivery with negligible local and systemic toxicities (Sanches et al. 2011; Yoon and Park 2010; Sirsi et al. 2012). Hence, ultrasound is capable of increasing efficacy of non-viral vectors through targeted delivery into specific tissues. Repeat applications can occur without compromising cell/tissue viability since sonoporation is a transient process. Minimal invasiveness of ultrasound, high targeting capacity, and suitability for repeat administrations all contribute to numerous clinical applications as highlighted in reviews by Delalande et al. (2012), Geis et al. (2012) and Farhood et al. (1995).

Despite the fact that good acoustic responses can be attained with microbubbles, in vivo transfection efficiencies remain suboptimal due to their size limitations. The large size of microbubbles may cause them to be subjected to shear forces encountered in the bloodstream (Böhmer et al. 2009). Through stable cavitation, microbubbles can enhance extravasation; however, they maintain a relatively short half-life (20 min in humans; few minutes in mice or rats) (Böhmer et al. 2009) and are hampered by limited perfusion and gene delivery beyond their immediate vicinity. Generation of nanobubbles addresses the above issues as their smaller sizes allow for increased circulation times, extravasation to less accessible regions, and retention by the enhanced permeability and retention effect (Nomikou and McHale 2012).

Bubble liposomes are echogenic liposomes composed of 1,2-distearoyl-sn-glycero-3-phosphatidyl-ethanolamine-polyethyleneglycol (DSPE-PEG₂₀₀₀-OMe) and a phospholipid [1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC) or 1,2-distearoyl-sn-glycero-phosphatidylcholine (DSPC)]; they are filled with perfluoropropane gas and have an average size that is less than 1 μm (Fig. 2.9) (Un et al. 2010, 2011a, b; Sugano et al. 2012; Omata et al. 2011, 2012; Suzuki et al. 2010; Endo-Takahashi et al. 2012; Horie et al. 2010, 2011; Kodama et al. 2010). Nucleic acid is mixed with bubble liposomes to generate an acoustically active bubble lipoplex. These bubble lipoplexes can be further functionalized by attaching different ligands [laminin peptide (Negishi et al. 2010), mannose (Un et al. 2010, 2011a, b), and folate (Omata et al. 2012)] for active targeting. Bubble liposomes, and other echogenic liposomes, have several suitable characteristics for enhanced gene delivery: (1) high gene-loading capacities similar to conventional liposomes, (2) active targeting upon functionalization, (3) controllable release, and (4) suitable echogenic properties for sonoporation and image-guided delivery (Huang 2008). Studies by Negishi et al. (2010) and Omata et al. (2011, 2012) have demonstrated co-administration of a PEGylated transfection vector and bubble liposomes with ultrasound to significantly enhance transfection efficiencies when compared to the vector alone; induced oscillations and cavitation by bubble liposomes are postulated to permit endosomal escape, a process generally inhibited as a result of PEGylation (Fig. 2.10). Improved endosomal escape is attributed to transient influx of Ca²⁺ associated with pore formation. As endosomal acidification is influenced by Ca²⁺, influx of Ca²⁺ affects endosome functionality and contributes to endosomal escape (Omata et al. 2011). It is also feasible to suggest that ultrasound-induced sonophores may also play a role in endosomal escape. Since bubble lipoplexes are similar to conventional lipoplexes, they are capable of transfecting cells in the absence of ultrasound. Studies by Endo-Takahashi et al. (2012) and Un et al. (2011a) have shown that using bubble lipoplexes as both the acoustically active agent and transfection vector results in direct gene delivery through pores rather than through endocytosis and endosomal escape. Enhanced transfection efficiencies with bubble liposomes are observed in in vivo gene delivery into tongue tissue (Sugano et al. 2012), bladder (Horie et al. 2010), and skeletal muscle (Kodama et al. 2010) as well as different formulations of cancer gene therapeutics (Un et al. 2010, 2011b; Suzuki et al. 2010; Horie et al. 2011).

Poly(d,l-lactic-co-glycolide) (PLGA) polymers have been considered excellent materials for nanobubble synthesis as they are approved by the FDA. They possess the following properties which are amenable for drug delivery applications: biodegradability, drug biocompatibility, mechanical properties, and ease of processing (Figueiredo and Esenaliev 2012). In addition, solid polymeric nanoparticles are capable of lowering the cavitation threshold, in the absence of gas bubbles, upon ultrasound exposure. Various aspects of PLGA in ultrasound-mediated gene delivery are described in a detailed review by Figueiredo and Esenaliev (2012).

Overall, physical forces such as ultrasound pulsing have been shown to greatly improve cell uptake and tissue penetration of oligonucleotides as delivered with lipid or polymeric carriers. Although the deep tissue penetration of sonoporation or microbubbles remains limited, their more transient and noninvasive nature makes this a highly valuable gene delivery tool, particularly for the delivery of genetic material past difficult barriers such as the BBB.

2.6 Combination of Acoustic and Magnetically Active Nanoparticles

Magnetofection and sonoporation both have inherent properties that enhance transfection efficiency (see Sects. 2.4 “Magnetofection” and 2.5 “Sonoporation—Nano Microbubbles and Bubble Liposomes”). Briefly, magnetofection enables rapid sedimentation of magnetic vectors to cell surface; this improves transfection kinetics and promotes close interaction between vector and cell membrane. Sonoporation facilitates internalization through transient pore formation and upregulation of endocytosis. In in vivo settings, magnetic vectors can achieve efficient localization at target sites through the administration of an external magnetic field, while microbubbles promote site-specific delivery upon the administration of ultrasound waves. In addition, both techniques are (1) minimally invasive, (2) are suitable for repeat administrations, and (3) act as stand-alone image modalities enabling diagnosis and image-guided delivery. The effects of magnetofection and sonoporation act in perfect marriage, eliciting synergistic enhancements to transfection efficiency (Fig. 2.11).

Magnetic microbubbles are gene delivery vectors that can be remotely controlled by magnetic fields and ultrasound. The combined effects of magnetofection and sonoporation can be elicited through the co-administration of magnetic nanoparticles (MNP) and nonmagnetic microbubbles along with activation by the application of magnetic field and ultrasound. Alternatively, magnetic microbubbles can be generated by loading MNPs into microbubbles or by attaching MNPs onto the surface of the microbubble (Owen et al. 2012). Through co-administration of l-α-phosphatidylcholine phospholipid microbubbles (~6 μm) and magnetic micelles (~2 μm), Stride et al. (2009) were able to demonstrate increased transfection of naked plasmid DNA in vitro; this assumption is due to sensitization of cells from magnetic micelles and magnetic field exposure. Similarly, the incorporation of magnetic micelles into microbubbles enhances transfection as a result of magnetically induced close proximities between cells and microbubbles (Stride et al. 2009). Mannell et al. (2012) were able to demonstrate efficient binding of plasmid DNA to magnetic nanoparticle-coated DPPC/DPPE phospholipid microbubbles and subsequent enhanced transfection efficiencies in vitro and in vivo; site-specific localization and targeted delivery are confirmed to be mediated by the synergistic effects of magnetofection and sonoporation. The incorporation of all three components (microbubble, MNP, and nucleic acid) has positive implications in in vivo applications as it resolves the issue relating to the differing biodistributions of the three components during co-administration.

Vlaskou et al. (2010a) have demonstrated the generation of magnetic and acoustically active lipospheres (MAALs) comprised of soybean oil, stabilized by Metafectene/1,2-dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE), with different surfactant-coated MNPs and perfluoropropane gas (Tween60-MAG-AAL and PEI-MAG2-AAL). The sequential addition of MNP and nucleic acid results in greater microbubble aggregation than simultaneous mixing of MNP and nucleic acid (Vlaskou et al. 2010a). Simultaneous mixing of MNP and nucleic acid prevented aggregations such that the microbubbles are suitable for systemic delivery; however, this comes at a cost as MAAL-nucleic acid interactions are substantially hampered. Localization of MAALs into specific targeted regions has been confirmed upon application of ultrasound and magnetic field under in vivo conditions; high nucleic acid-binding capacity and magnetically induced close proximities between cells and microbubbles are observed in vitro as evidenced by enhanced transfection efficiency with magnetic field (Vlaskou et al. 2010a). However, MAALs do not possess sufficient acoustic activities to enhance transfection by sonoporation. This paves the way for the development of a second generation of magnetic microbubbles comprising of a DPPC/DSPE-PEG₂₀₀₀-biotin lipid bilayer similar to that of a bubble liposome (Vlaskou et al. 2010b). The incorporation of MNPs and nucleic acids, along with subsequent loading of perfluoropropane gas, results in magnetic microbubbles with sizes of 1–2 μm. Using the second-generation magnetic microbubbles, gene delivery occurs in a plasmid concentration-dependent manner, a feature not observed in nonmagnetic microbubbles (Vlaskou et al. 2010b). In addition, such enhanced gene delivery can only be observed upon ultrasound exposure; hence, these microbubbles are acoustically active and can induce sonoporation. The protocol for the synthesis of MAALs and its second-generation microbubbles is described in detail by Vlaskou et al. (2013).

The physical (size, size distribution, and concentration) and acoustic properties of magnetic microbubbles only vary slightly when compared to nonmagnetic microbubbles of the same formulation (Mulvana et al. 2012). The magnetic properties of each individual microbubble vary depending on the amount of MNPs bound to the microbubble. Application of a magnetic field induces an attractive force that can overcome the shear forces of blood flow and localizes magnetic microbubbles in a confined area (Mannell et al. 2012). Subsequent microbubble disruption by ultrasound increases endothelial permeability, increases extravasation, and enhances gene delivery. The concerted effects allow site-specific targeting with high precision; this reduces the likelihood of adverse effects. Thus, the use of magnetic microbubbles for gene delivery poses to be a novel approach that will be thoroughly explored in the coming years.

2.7 Optical Transfection

Optoinjection, the light-mediated transfer of membrane-impermeable substances, and optical transfection, the process of light-mediated nucleic acid delivery, stem from the induction of membrane permeations and the generation of small transient membrane pores (optoporation) through the administration of continuous wave (CW) or pulsed laser irradiation. There are several inherent advantages in using optical transfection for nucleic acid delivery. The noncontact nature of optical transfection permits an aseptic means of delivery in sterile environments (Praveen et al. 2011; Soman et al. 2011). High precision of laser irradiation enables selective delivery with high specificity and minimal collateral damage. Lastly, optical transfection can further be integrated with microscopy imaging, optical tweezers, and microfluidic systems (Soman et al. 2011; Antkowiak et al. 2010). The combination of optical transfection and microscopy imaging enables simultaneous visualization of nucleic acid delivery and subsequent intracellular trafficking processes. High selectivity and high throughput can be achieved through integration with optical tweezers and microfluidic systems (Kim et al. 2011; Rendall et al. 2012). However, such advantages are overshadowed by insufficient penetration depth, approximately only 2 mm using infrared wavelengths within the optical window, which severely limits in vivo applications of optical transfection (Stevenson et al. 2010).

The proposed mechanism of membrane permeation varies with respect to laser source, wavelength, and pulse duration (Antkowiak et al. 2010; Kim et al. 2011; Stevenson et al. 2010; Yao et al. 2008). The earliest works have employed ultraviolet (UV) light to induce optoporation on the basis of light absorption by membrane components; however, the application of UV light raised concerns relating to irreversible membrane damage, target cell death, as well as potential damage to the administered exogenous DNA (Yao et al. 2008). The implementation of blue CW lasers (488 nm) has been shown to induce transient membrane permeability for the delivery of plasmid DNA through localized heating from light absorption by phenol red, a cell medium component (Kim et al. 2011; Stevenson et al. 2010; Yao et al. 2008). Using this technique, high postirradiation viability and successful stable/transient transfections have been reported, albeit at low transfection efficiencies (Stevenson et al. 2010). Although the utilization of blue light reduces significant detrimental effects, potential cellular alterations and deleterious effects may still occur from absorption by various intracellular constituents.

Alternatively, there has been the development of pulsed lasers, whereby the mechanism of membrane permeation and their resultant effects are dependent on pulse duration. The application of nanosecond-pulsed laser generates photomechanical stress waves responsible for transfection by laser-induced stress waves (LISW) (Fig. 2.12). The generation of stress waves induce shear forces that cause transient increase in membrane permeability without affecting viability so long as the applied stress waves are maintained below the damage threshold; this allows entry of exogenous DNA by passive diffusion (Yao et al. 2008). As such, cell viability and transfection efficiencies greatly depend on the distance between cells and the propagation of the laser-induced shockwave. LISW in gene therapy has been effectively applied, in vitro and in vivo, through the implementation of experimental configurations highlighted in Fig. 2.12. Irradiation with 552-nm nanosecond laser pulses from a Q-switched Nd:YAG laser, targeting an elastic natural rubber disk (laser target), creates stress waves resulting from the optical breakdown, ablation, or rapid heating of the laser target; stress wave propagation through cells and/or skin on the opposite side induces transient increase in membrane permeability suitable for gene delivery (Yao et al. 2008; Terakawa et al. 2004).

Using this configuration, with careful consideration to several critical parameters (e.g., impulse and number of LISWs), transfection of plasmids encoding luciferase, enhanced green fluorescent protein (EGFP), and/or β-galactosidase have been demonstrated in NIH3T3 murine fibroblast (Terakawa et al. 2004) in vitro as well as rat skin/muscle (Kurita et al. 2011; Ogura et al. 2004) and mouse central nervous system (CNS) (Satoh et al. 2005) in vivo (Fig. 2.13). Irradiation following intradermal injection of plasmid DNA has shown to induce selective expression in rat epidermal cells in vivo (Ogura et al. 2004). In addition, a positive correlation between transgene expression and number of administered pulses (up to three pulses) has been observed; administration of three pulses constitutes an increase in transgene expression corresponding to a 120-fold increase in luciferase activity along with sustained luciferase expression for up to 5 days post-transfection (Ogura et al. 2004). The application of LISW in combination with delivery vectors composed of PEI/DNA polyplexes has yielded efficient and widespread EGFP expression in mouse CNS, achieving penetration depths of 1.5 mm and 3.5 mm into the cortex surface of adult and newborn mice, respectively (Satoh et al. 2005). Similarly, the combination of LISW and PEI/DNA vectors have resulted in significant increases in transgene (lacZ for expression of β-galactosidase) expression as demonstrated by greater β-galactosidase activity in rat facial muscles; β-galactosidase activity has been shown to sustain for up to 14 days with peak activity 7 days post-transfection (Kurita et al. 2011). Gene delivery by LISW therefore enables simultaneous treatment of a large number of cells while achieving targeted delivery through controlled administration of stress waves. In addition, LISW permits deep tissue treatments as stress wave propagation exceeds the limited penetration depth attained from direct irradiation (Terakawa et al. 2004). Lastly, LISW by Q-switched Nd:YAG laser offers additional advantages as they are commercially available, cost-effective, easy to operate, safe, and approved for use in head and neck surgeries (Kurita et al. 2011).

In contrast to nanosecond-pulsed lasers, the administration of femtosecond-pulsed lasers, involving near-infrared (NIR) (700–1100 nm) multiphoton processes, has been generally applied for single cell transfections yielding high transfection efficiency and low cytotoxicity (Praveen et al. 2011; Antkowiak et al. 2010; Stevenson et al. 2010; Yao et al. 2008). Direct single target cell irradiation leads to photon absorption and free electron excitation, which contributes to plasma formation and subsequent localized changes in membrane permeability including the formation of cavitation bubbles and submicron transient pores (Praveen et al. 2011; Stevenson et al. 2010; He et al. 2011; Baumgart et al. 2012). Transient pore formation permits the delivery of exogenous DNA in the absence of substantial thermal effects due to ultrashort pulse duration and high repetition rates. Ultrashort pulses prevent heat accumulation, while high repetition rates, greater than the critical repetition rate of 1 MHz, enable the diffusion of excess thermal energy (He et al. 2011). Thermal effects can also be mitigated by using objectives with high numerical aperture (NA). Higher NA objectives are commonly used to focus the laser beam in order to reach high photon density; the resulting small focal volumes permit adequate diffusion of thermal energy into the surroundings. Objectives with NA greater than 1 will exhibit low heat accumulation, whereas objectives with NA of 0.6 or less will exhibit substantial heat accumulation and temperature increase (He et al. 2011).

Femtosecond-pulsed laser-mediated transfections are at a disadvantage due to low throughput, limited loading concentrations, and the requirement for precise instrumentation to achieve direct laser alignment (Yao et al. 2008). Such issues may be mitigated with the implementation of Bessel beams and spatial light modulator (SLM). Laser alignment is crucial for successful transfection as direct membrane interaction can only be achieved upon precise positioning of the beam on the membrane. The use of a Gaussian beam demands precise laser alignment due to a very short effective focal length contributing to a confined axial therapeutic range; such properties raise difficulties in focusing the laser beam on the membrane as any minute misalignments, as little as 3 μm, may result in over 50% reduction in transfection efficiency (Stevenson et al. 2010; He et al. 2011; Tsampoula et al. 2007). Bessel beams, characterized as non-diffracting rods of light, possess an axial therapeutic range that is 20 times greater than that of Gaussian beams, thus overcoming issues relating to misalignment and the requirement for precise beam focus (He et al. 2011; Tsampoula et al. 2007). Since only the central core has sufficient energy to induce free electron excitation and plasma formation, Bessel beams introduce less damaging effects in comparison to out of focus CW light (Stevenson et al. 2010). In combination with SLM for steering and emission of multiple arrays of Bessel beams, the potential for “point and shoot” capabilities may enable simultaneous targeted transfection of preselected groups of cells (Stevenson et al. 2010; Cizmár et al. 2008). As demonstrated by Antkowiak et al. (2010), SLM enables submicron three-dimensional positioning that results in consistent, repeatable poration and time-sequenced irradiation at multiple axial and lateral positions (Fig. 2.14). The application of “point and shoot” targeted transfection has permitted selective co-transfections which resulted in co-expression green fluorescent protein (GFP) and Mito-DsRed (Fig. 2.14). Using unassisted raster scan irradiation, the authors are able to achieve high-throughput transfections at an average rate of one cell per second (Antkowiak et al. 2010).

Previous studies have shown successful in vitro optical transfection of various cell lines with varying transfection efficiencies: Chinese hamster ovary (CHO, 63%), human embryonic kidney (HEK, 52%), mouse neuroblastoma-rat glioma hybrid (NG108-15, 40%), and human neuroblastoma (SK-N-SH, 45%) cells (Dholakia 2010). In contrast, Tirlapur and König reported transfection efficiencies of 100% in CHO and rat-kangaroo epithelial cells (PtK2) in 2002 (Tirlapur and König 2002). The differences in transfection efficiency may lie in the laser configuration and choice of plasmid construct. Of significance, optical transfection has yielded successful delivery into zebrafish embryos and mouse/human embryonic stem cells. Through optical transfection, the delivery of Simian-CMV-EGFP plasmids into blastomeres of early to mid-cleavage stage zebrafish embryos results in wide distribution of EGFP expression along yolk extension, notochord, floor plate, and somites (Kohli et al. 2007). Application of a sub-20 femtosecond infrared laser pulse constitutes 70–80% transfection efficiency for human salivary gland and human pancreatic stem cells with 100% viability and cell division for some of the transfected cells (Uchugonova et al. 2008). Optical transfection of mouse embryonic stem cells (E14g2a) with plasmid expressing Mito-DsRed (pDsRed2-Mito) has shown to result in 25% transfection efficiency with demonstrated differentiation potential upon the delivery of endoderm-associated transcription factor Gata-6 gene; development of spindle- and stellate-shaped morphology along with expected up-/downregulation of transcriptional factors, characteristic of extraembryonic endoderm tissue, confirms Gata-6-induced stem cell differentiation (Dholakia 2010). Ishii et al. (2017) recently tracked the intracellular localization of fluorescently labeled plasmid DNA-liposomal complexes (Lipofectamine 3000) as delivered through femtosecond laser pulsation. While DNA uptake occurred regardless of laser irradiation, gene expression was only observed in the event of laser irradiation. They hypothesized that laser irradiation may enhance persistence of the DNA/liposome complex in the endosomal compartment, reducing its degradation in the lysosomal pathway. They further observed increased transport of DNA into the nucleus upon laser irradiation, which they attributed to perforation of the nuclear membrane and subsequent diffusion of DNA/liposome complexes toward the nuclear region. These highly promising results further elucidate the mechanisms by which laser-aided transfections can improve non-viral gene delivery.

Femtosecond-pulsed laser transfections have been applied in vivo for intradermal gene delivery, implicating the potential for long-term gene expression and genetic immunization. As demonstrated by Zeira et al. (2007), intradermal injection of plasmid into mice, conferring luciferase reporter expression, constitutes a 625-fold increase in luciferase expression 43 days after 2-minute irradiation under optimized parameters (780 nm, 200 fs pulses, 76 MHz); luciferase expression remained stable for over 7 months in contrast to a lack of transgene expression for nonirradiated mice. Vaccination by intradermal injection of hepatitis B virus surface antigen (HBsAg) expressing plasmids and femtosecond-pulsed irradiation induce humoral and cellular immune activation, denoted by elevated levels of IgG1, IgG2a, IFN-γ, and IL-4, along with increases in HBsAg-specific titers that contributed to a 215-fold higher titer on day 93 and 22-fold higher titer on day 210. Vaccinated mice have been shown to be immunized against HBsAg-transfected tumor cells with significant reduction in tumor growth over mice without irradiation (Zeira et al. 2007). Immunity is confirmed to be tumor specific as tumor growth retardation is not observed after implantation of untransfected tumor cells.

In conjunction with optical transfection through direct laser-membrane interactions, optical transfections have been achieved with light-absorbing particles including gold and carbon nanoparticles. As demonstrated by Felgner et al. (1987), the application of 100 nm off-resonance spherical gold nanoparticles and low fluence, femtosecond-pulsed irradiation enables plasma formation, nanocavitation, and membrane permeabilization while preventing nanoparticle fragmentation through limited photon absorption. Low fluence laser irradiation ensures nanoparticle-dependent optical transfection, achieving localized gene delivery into WM278 human melanoma cells with 23% transfection efficiency and low gold nanoparticle cytotoxicity (<1%). Similarly, Wu and Wu (1987) achieved a 17-fold increase in luciferase reporter gene expression after gene delivery into DU145 human prostate cancer cells upon femtosecond irradiation of carbon black nanoparticles. However, the mechanism of membrane permeabilization is attributed to a photoacoustic effect producing cavitational shockwaves through light-mediated carbon steam reaction.

In addition to gene delivery by optical transfection, light has been implemented for endosomal escape and activation of RNA interference (RNAi) upon successful internalization of small interfering RNA (siRNA) or small hairpin RNA (shRNA) (Fig. 2.15). Photochemical internalization (PCI) employs light activation of internalized photosensitizers (PS) for reactive oxygen species (ROS) generation that facilitates photochemically induced endosomal escape (Yao et al. 2008; Matsushita-Ishiodori and Ohtsuki 2012). Endocytosis of siRNA complexes with PS permitted cytosolic release for siRNA/shRNA; however, RNAi by PCI must be finely tuned in order to circumvent issues relating to PS toxicity and potential photodynamic effect on other photodynamic therapy (PDT) targets including nuclei, mitochondria, lysosomes, and plasma membrane. CPP-linked RBP-mediated RNA internalization and photoinduced RNAi (CLIP-RNAi) served as an improvement to PCI as the PS is directly linked to shRNA through a photosensitizing carrier consisting of a cell-penetrating peptide, an RNA-binding protein, and a photosensitive fluorescent dye; application of CLIP-RNAi has shown a RNAi efficiency of up to approximately 80% while preventing any un-induced leakage activity observed in PCI (Matsushita-Ishiodori and Ohtsuki 2012; Matsushita-Ishiodori et al. 2011). Gold nanostructures, with strong surface plasmon resonance absorption at NIR wavelengths and acceptable in vitro/in vivo toxicities, enable temporal- and spatial-controlled RNAi by photothermal transfection (Matsushita-Ishiodori and Ohtsuki 2012; Braun et al. 2009; Lu et al. 2010a).

As demonstrated by Braun et al. (2009), application of a TAT-lipid-coated gold nanoshell-siRNA conjugate permitted efficient internalization by endocytosis and subsequent endosomal escape by heating, cavitation, and endosome leakage upon femtosecond-pulsed irradiation of gold nanoshells. At the same time, irradiation causes Au plasmon-induced destabilization and cleavage of polyethylene-glycol-thiol bonds for the dissociation of surface-bound anti-EGFP siRNA from gold nanoshell structures, thus achieving cytosolic release and RNAi. Photoinduced dissociation of the siRNA was denoted by controllable increases in fluorescence of siRNA-bound Cy3, a fluorescent molecule that was actively quenched by gold. Through photothermal transfection, the authors are able to induce ~80% reduction in GFP; a lack of cytosolic release of nanoshells corroborates with the lack of GFP downregulation in absence of irradiation (Braun et al. 2009). Using folate functionalized, anti-NF-ΚB p65 siRNA-bound, hollow gold nanospheres (F-PEG-HAuNS-siRNA), Lu et al. (2010a) achieved selective uptake by HeLa cells overexpressing folate receptors and up to 95% p65 downregulation 72 h after photothermal transfection. The authors attributed the cytosolic release of siRNA to structural transformations of nanospheres, as confirmed by TEM, accompanying siRNA dissociation. In vivo studies, with mice harboring HeLa cervical cancer xenografts, indicate significant uptake of folate functionalized nanospheres and significant downregulation of NF-ΚB p65 after NIR irradiation. Enhanced apoptotic response, achieved through the combined treatment of p65 siRNA photothermal transfection and irinotecan, contributes to a significant delay in tumor growth (Lu et al. 2010a). Building on the concept of RNAi by photothermal transfection, an elegant system of photonic gene circuits for bimodal photothermal regulation of NF-ΚB p65 expression has been created; the development of such an intricate system is described in detail by Lee et al. (2012).

2.8 Gold Nanostructures

Gold nanoparticles (GNPs) are attractive for use in drug and gene delivery due to a number of favorable properties such as inertness and low toxicity, straightforward synthesis with low size dispersion, easy functionalization (usually through thiol linkages), and photophysical properties (Ghosh et al. 2008). Depending on the synthetic method used, gold nanoparticle core sizes have been known to range from 1 to 150 nm (Ghosh et al. 2008). In the following section, the use of GNPs to gene delivery will be discussed, with focus on the properties of GNPs that uniquely facilitate such an application.

2.8.1 Gold Nanoparticles as Oligonucleotide Carriers

When used as non-viral gene delivery vectors, strategies to achieve efficient gene delivery often involve chemically functionalizing the gold nanoparticles either with the therapeutic oligonucleotides themselves or with other existing transfection agents to aid in complex formation and delivery (Barhoumi et al. 2009). When DNA is bound to gold nanostructures, steric hindrance increases the DNA half-life from minutes to hours against digestion from nucleases (Barhoumi et al. 2009). Additionally, electrostatic repulsion between polyvalent cations near the gold surface and dications within the nucleases contributes to oligonucleotide stability (Barhoumi et al. 2009).

Chemical conjugation of cationic molecules such as polyethyleneimine (PEI), quaternary ammonium chains, cationic lipids, or cationic amino acids has been used to improve electrostatic interaction with the nucleic acid cargo, thus enabling efficient gene delivery (Cebrián et al. 2011). For example, thiol-modified 2 kDa PEI has been conjugated to gold nanoparticles and used to deliver green fluorescent protein (GFP)-encoding plasmid DNA in vitro (Sharma et al. 2011). Here, GFP expression was seen in nearly 50% of human corneal fibroblasts within 12 h following transfection and cellular phenotype, and viability appeared unaltered with nitrogen/phosphorous ratios up to 180 (Sharma et al. 2011). Topical application of PEI2-GNPs to rabbit corneas in vivo resulted in PEI2-GNPs uptake 12 h later, with an indication of slow clearing from the cornea and no significant immune response compared to untreated corneas (Sharma et al. 2011). Furthermore, transmission electron microscopy data of treated rabbit corneal keratocytes showed PEI2-GNPs within the extracellular matrix and endosomes near the cell surface, as well as evidence of endosomal rupture in corneal keratocytes, suggesting the nanoparticles are taken up by endocytosis and that PEI in PEI-conjugated gold nanoparticles likely continues to facilitate cargo release from the endosome through the “proton sponge” effect seen in other PEI-containing systems (Sharma et al. 2011).

Similarly, gold nanoparticles may be encapsulated by a polymer to improve DNA binding and delivery. GNPs entrapped by amine-terminated generation 5-poly(amidoamine) (PAMAM) dendrimers (G5.NH₂) formed smaller complexes with plasmid DNA (pDNA) than G5.NH₂ alone and demonstrated lower cytotoxicity and 100 times greater transfection efficiency (Shan et al. 2012). Shan et al. claimed that the entrapped gold nanoparticle helped preserve the 3D spherical morphology of the dendrimer, allowing for more binding sites with DNA (compared to dendrimers without a gold core), higher compaction of the pDNA, smaller particle sizes, and enhanced gene delivery (Shan et al. 2012).

In another scenario, methylated N-(4-pyridinylmethyl) chitosan chloride (MPyMeChC) and PEI have also been used to generate multilayer gold-polymer-DNA nanoparticles (Tencomnao et al. 2011). Figure 2.16 shows an illustration of the nanoparticles used by Tencomnao et al. in 2011. Use of the gold nanoparticle scaffold in this study enhanced transfection efficiencies compared to the C/D and P/D polyplex systems (Tencomnao et al. 2011). Furthermore, combining gold and PEI resulted in transfection efficiency 10 times greater than that of DNA with PEI (P/D) alone; G/C/D/P and G/P/D complexes produced results comparable to Lipofectamine 2000 in HeLa and A549 cell lines (Tencomnao et al. 2011). PEI-containing multilayer nanoparticles (G/C/D/P and G/P/D) showed higher efficiency than that of G/C/D systems, seemingly due to the proton sponge effect brought along by the PEI content (Tencomnao et al. 2011). Additionally, the investigators claimed that the use of colloidal gold improved the monodispersity of the carrier samples as well as facilitated easier release of the DNA after cell internalization compared to polyplex systems where DNA is tightly condensed (Tencomnao et al. 2011). In cell viability assays, toxicity increased with the introduction of gold (60–80% viability) versus C/D and Lipofectamine (80%); however, the toxicity remained lower than that of P/D complexes (~50% viability) (Tencomnao et al. 2011).

2.8.2 Photophysical Properties of Gold Nanoparticles and Gene Delivery

Alternatively, a number of groups have exploited the photophysical properties of gold nanoparticles and nanorods, namely, the surface plasmon resonance, to achieve remotely triggered gene delivery to target cells. In brief, exposure to an electromagnetic frequency prompts a collective oscillation of the valence electrons at the surface of spherical nanoparticles with a diameter much smaller than the light wavelength (Boisselier and Astruc 2009). This oscillation is known as the surface plasmon resonance, which is strongest for noble metals (gold, silver, and copper) and is shifted into the visible portion of the electromagnetic spectrum for these metals (Boisselier and Astruc 2009; Link and El-Sayed 2003). The width of the surface plasmon absorption and frequency depend on the nanoparticle size and shape, as well as the dielectric constant of the metal and the surrounding medium (Link and El-Sayed 2003). Small nanoparticles (<100 nm) absorb light and display this surface plasmon resonance; meanwhile as the particle size increases to the micron scale, the particle becomes a better light scatterer than light absorber (Barhoumi et al. 2009; Boisselier and Astruc 2009). Therefore manipulation of the nanoparticle shape and size allows its resonance to be tuned to a desired resonant frequency (Barhoumi et al. 2009; Boisselier and Astruc 2009). For example, nanoparticles with optical resonances in the near-infrared (NIR) range, where blood and tissue are nearly transparent, are especially useful for biomedical applications (Barhoumi et al. 2009; Boisselier and Astruc 2009). Furthermore, upon illumination, plasmon-resonant nanoparticles dissipate energy through their lattice phonons, converting light to heat (Barhoumi et al. 2009). Thus, the nanoparticle’s photothermal response may be used for applications such as photothermal tumor ablation or light-assisted biomolecule delivery like those discussed below (Barhoumi et al. 2009).

Light-induced, specifically NIR-induced, release of DNA or RNA cargo is a promising method of achieving controlled release for gene delivery. In 2009, nanoshells consisting of a spherical silica core and a gold shell with a plasmon frequency determined by its inner and outer radii were used to demonstrate light-induced release of single-stranded DNA (ssDNA) for antisense therapy (Fig. 2.17) (Barhoumi et al. 2009). In this case, double-stranded DNA (dsDNA) was covalently bound to the gold shell via a 5′-thiol group present only on the sense strand, leaving the complementary strand (the therapeutic antisense oligonucleotide) free to be released from the nanostructure upon thermally induced or NIR light-induced DNA melting (Barhoumi et al. 2009). Light-induced DNA release occurred near-ambient temperature in solution (minimal temperature increase was observed) and appeared to be independent of length for oligonucleotides in the range of 20–70 base pairs (Barhoumi et al. 2009; Huschka et al. 2011). The mechanism for dehybridization under these conditions appears to be due to hot electrons (from plasmon decay) transferred to the adsorbed DNA leading to dehybridization due to repulsion between the two strands (determined in a subsequent study) (Barhoumi et al. 2009; Huschka et al. 2011). On the other hand, thermally induced release was attained at the DNA melting temperature, which is influenced by the nucleotide sequence, type of linkage to the nanoparticle surface, and solution properties such as ion concentration (Barhoumi et al. 2009; Huschka et al. 2011). There was also a notable difference in release efficiency between the two methods; about 50% of the available ssDNA was released with light, but virtually all of the ssDNA was released using the thermal method (Barhoumi et al. 2009; Huschka et al. 2011).

The nanoshell-dsDNA complex was also used for NIR light-triggered release of 4′,6-diamidino-2-phenylindole (DAPI) in H1299 lung cancer cells in vitro (Huschka et al. 2010). Here it was demonstrated that intracellular light-triggered release can be achieved and visualized with a DNA-binding fluorescent molecule such as DAPI and that the complex and the laser power and irradiation time does not adversely affect cell viability (Huschka et al. 2010).

On the other hand, Lukianova-Hleb et al. took a different approach in using the light-to-heat converting property of plasmon-resonant GNPs to achieve cell-specific gene delivery within heterogeneous cell samples. Gold spheres 60 nm in diameter were conjugated with target-specific antibodies and added to the cells, accumulating in target cells more than nontarget cells through receptor-mediated endocytosis (Lukianova-Hleb et al. 2012). Cell-specific delivery was described as follows (Fig. 2.18). When exposed to a laser pulse, the optical energy was converted to thermal energy through plasmon resonance of the gold nanoparticle clusters. Local heat released by the GNPs evaporated the liquid nanoenvironment, resulting in the generation of transient vapor nanobubbles, known as plasmonic nanobubbles (PNBs). In this case, only minimal optical energy was supplied by the pump laser pulse so that only large clusters of GNPs (mostly in target cells) are capable of forming PNBs at such a low energy; PNB generation was therefore prevented around the smaller clusters and single nanoparticles present in nontarget cells (Lukianova-Hleb et al. 2012). The expanding PNBs formed a transient and reversible nanohole in the cell membrane, and upon PNB collapse, an in-bound jet injected the extracellular matrix including the cargo into the cytoplasm. The investigators reported large nanoparticle clusters in 99% of target cells and successful (96%) target-specific transfection with a GFP-encoding plasmid in the heterogeneous cell sample of CD3-positive human T cells (target) and CD3-negative cells (nontarget) after a single pulse flow treatment along with 75% viability in treated target cells (Lukianova-Hleb et al. 2012).

The reader is referred to the following materials for examples of similar techniques in gold nanoparticle-mediated delivery of other nucleic acid structures such as small interfering RNA (siRNA), microRNA, and short hairpin RNA (shRNA): Lu et al. (2010a), Ghosh et al. (2013), Kong et al. (2012b), Conde et al. (2012), Ryou et al. (2011), Zheng et al. (2012).

A final application of gold nanoparticles is their use as templates in the development of spherical nucleic acid structures which are capable of cellular internalization and gene regulation (Young et al. 2012; Cutler et al. 2012). These structures traditionally consist of gold cores with DNA shells, but investigations have shown that cross-linking the DNA at the base and subsequently dissolving the gold core produces a hollow structure with many of the key properties of the gold nanoparticle conjugates without concerns of toxicity from the gold core (Young et al. 2012; Cutler et al. 2012).

2.8.3 Gold Nanorods

Similarly, gold nanorods have attracted a considerable amount of attention in gene delivery among other applications such as photothermal therapy, biosensing, and imaging, due to their near-infrared light-enhanced absorption and plasmon resonance (Cui et al. 2011). Owing to the surface plasmon resonance of gold nanorods being easily tunable through alterations in their aspect ratio, gold nanorods can be used for selective and/or controlled release of DNA molecules attached to their surface in the same manner described above for gold nanoshells.

For example, by exploiting the association between gold nanorod aspect ratios and their corresponding surface plasmon resonance, one group was able to demonstrate selective release of two differentially labeled “thiolated DNA 40mers” in a mixed sample of DNA-conjugated gold nanorods (Wijaya et al. 2009). Thiolated 6-carboxyfluorescein-labeled oligomers were conjugated to short, “nanocapsule,” nanorods that had a longitudinal surface plasmon resonance at 800 nm; meanwhile, thiolated tetramethylrhodamine-labeled oligomers were conjugated to long, bone-shaped nanorods with a longitudinal surface plasmon resonance at 1100 nm (Wijaya et al. 2009). With this method, when a mixture of the purified, conjugated nanorods was exposed to laser irradiation at one of the surface plasmon resonance wavelengths, only the corresponding nanorod was excited and released its DNA cargo from the surface (50–80% DNA release), leaving the other essentially unchanged (less than 10% DNA release) and the DNA still functional (Wijaya et al. 2009). The release was also tunable by laser fluence (Wijaya et al. 2009). This type of system has potential for multidrug/gene delivery.

On the other hand, in a 2011 study comparing thermal and light-mediated release of ssDNA from a gold nanorod ([w,l] = [13,47]) surface, no light-induced release was achieved upon excitation of either the transverse or the longitudinal plasmon resonance (Huschka et al. 2011). This was evidenced by the resultant ssDNA release curves, which were very similar to the thermal release curves and showed no significant increases in release at temperatures below the DNA melting temperature (Huschka et al. 2011). Huschka et al. suggested that light-triggered release occurs through a nonthermal mechanism where hot electrons generated by plasmon excitation of the gold nanostructure are transferred to the adsorbed cargo, thus facilitating DNA dehybridization well below its melting temperature (Huschka et al. 2011). Therefore, the capacity for light-triggered release is affected by the absorption cross section and DNA densities on the nanorod surfaces (Huschka et al. 2011). The calculated near-field intensity enhancements were large for nanorod longitudinal excitation, but this local field would be confined to the tips of the rod and decay rapidly with increasing distance from the gold surface (Huschka et al. 2011). Consequently, if charge transfer is higher in regions of large local field, it is possible that an insufficient amount of DNA was susceptible to charge transfer, resulting in a negative result under the experimental conditions (Huschka et al. 2011).

However, it should be noted that the methods discussed in this section have yet to be successfully demonstrated to any large extent in biological systems. The biosafety of gold nanoparticles has already been demonstrated elsewhere, most notably in clinical trials against prostate cancer (Stern et al. 2016). A recent review by Pedrosa et al. (2015) further highlights investigations into GNPs in preclinical and clinical phases. Additionally, a preliminary study of layered polyelectrolyte-conjugated gold nanoparticles demonstrated encouraging results in cultured prostate cancer cells (Huang et al. 2009). Oligonucleotide-conjugated nanorods were used for optically activated release of antisense oligonucleotides for mRNA interference in breast carcinoma cells in vitro as well (Lee et al. 2009). More recently, a series of transfection studies was carried out examining PEGylated PEI-GNP complexes and their ability to transfect HeLa and HEK 293T cell lines with plasmids of different sizes (3000 bp to 40 kbp) (Mar Encabo-Berzosa et al. 2017). In comparison to a standard transfection reagent (Lipofectamine 2000), their PEI-GNP complex performed admirably; transfection efficiencies from gold-mediated gene delivery even surpassed that of Lipofectamine for the larger DNA (40 kbp). Their studies support the combinatorial approach of polymeric or liposomal gene carriers in conjunction with gold nanoparticles as a viable gene delivery vehicle.

2.9 Carbon Nanostructures for Gene Delivery

Carbon nanostructures hold great promise in their potential as nanoscale biomaterials. Some carbon nanostructures have been deemed useful in gene delivery applications, each with their own favorable properties. The following discussion will be organized by structure type, highlighting particular structures found favorable for gene and drug delivery.

2.9.1 Carbon Nanotubes

Carbon nanotubes (CNTs) are well suited for use in the field of nanomedicine given their high aspect ratios, electronic and optical properties, and large surface area, which can be functionalized and loaded with nucleic acid, drug, or protein therapeutics (Cheung et al. 2010; Varkouhi et al. 2011). Moreover, there have been claims that CNT systems enter cells through a nanoneedle, endocytosis-independent mechanism, possibly contributing to the enhanced transfection results seen with some CNT-mediated delivery systems. However, a definite conclusion regarding the cellular uptake mechanism of CNTs has not been made since groups in the field have reported varying results. In one corner lies the work of Dai et al. (Shi Kam et al. 2004; Kam and Dai 2005; Kam et al. 2006) who, through systematic uptake inhibition assays, demonstrated that energy-dependent endocytosis, specifically through clathrin-coated pits, is essential for cellular internalization of short, well-dispersed protein- or DNA-bound single-walled CNTs. Investigations using the near-infrared fluorescence and Raman properties of single-walled CNTs by Jin et al. (2009) also support this view. In contrast, results from Kostarelos et al. (2007), Lacerda et al. (2012) and Pantarotto et al. (2004) lie in the opposing corner, where CNTs are thought to enter cells independent of energy-dependent endocytosis. A series of CNTs functionalized with various small molecules appeared to enter cells regardless of the functional group, cell type and ability for endocytosis, and presence of endocytosis inhibitors. Given these observations, the cylindrical shape of CNTs and their high aspect ratio, some groups (Cai et al. 2005; Rojas-Chapana et al. 2005; Kostarelos et al. 2007) proposed that CNTs penetrate the cell membrane, like a “nanosyringe.” Such notable discrepancies between groups may be due to significant differences in the CNT characteristics: functionalization with macromolecules such as proteins and DNA versus smaller molecules, as well as CNT length (Cheung et al. 2010; Kang et al. 2010). Perhaps different cellular entry mechanisms will dominate depending on the physical properties of the CNTs in question.

For a summary of CNTs for gene delivery to date, the reader is referred to some excellent reviews elsewhere (Cheung et al. 2010; Marchesan et al. 2015). Similar to gold nanostructure delivery systems, investigations have involved surface-modified CNTs for delivery of nucleic acid cargo to cells in vitro. When left “as-is” CNTs are water-insoluble; therefore, chemical modification is necessary for their use in gene delivery applications (Liu et al. 2011b). This not only allows for greater solubility and individualization in hydrophilic environments, but the cationic surface modifications also provide a necessary positive surface charge for interaction with cellular membranes and nucleic acid binding. Binding can occur either through electrostatic interactions or chemical conjugation, depending on the functionalization used.

A number of functionalized CNTs have been designed for nucleic acid delivery, including those functionalized with dendrimer polymers. (Poly(Lys:Phe))-coated double-walled CNTs (DWCNTs) were conjugated to thiol-modified siRNA targeted to survivin as a demonstration of an anticancer non-viral siRNA strategy (Neves et al. 2012). CNT-delivered siRNA demonstrated successful transfection in vitro and exhibited similar levels of survivin knockdown and apoptosis as compared with DharmaFECT delivery of siRNA. Furthermore, the DWCNTs do not show any accumulation within cells after transfection and do not induce cellular stress as measured by mitogen-activated protein kinase phosphorylation. Work by Liu et al. using multi-walled carbon nanotubes (MWCNTs) conjugated to linear polyamidoamine demonstrated lower cytotoxicity and produced in vitro transfection results comparable or higher than that of the polyamidoamine and branched PEI (25 kDa) controls, with little serum sensitivity (Liu et al. 2011b). Intracellular trafficking results with Cy3-labeled pGL3 indicated greater nuclear localization after 24 h. A study reported by Podesta et al. employed amino-functionalized MWCNTs to deliver a toxic siRNA sequence to cells in vitro and in vivo through intratumoral injection into xenograft tumors (Podesta et al. 2009). In vitro transfection with the functionalized MWCNTs (compared to 72 h with the liposomal control) and significant tumor reduction was observed in vivo, corresponding to prolonged animal survival. More recently, MWCNTs were modified with amine groups to deliver siRNA against PLK1 (anti-lung cancer target) in tumor xenografts against a liposomal control, demonstrating significant siRNA retention and uptake in tumors after local administration (Guo et al. 2015).

Alternatively, carbon nanotubes might also be magnetized for improved delivery. Cai et al. demonstrated delivery of and expression of plasmid DNA via magnetically directed CNT penetration (“nanotube spearing”) into target cells (Cai et al. 2005). Similar to physical methods of gene delivery, the plasmid DNA bypasses the endocytic pathway and may even penetrate pass the nuclear membrane, potentially increasing transfection efficiency in slow- or nondividing cells. CNTs carrying ferromagnetic nickel particles were sensitized to magnetic agitation. Within a magnetic field, the CNTs complexed with pEGFP-c1 could then be oriented and directed to spear cell membranes in both dividing (Bal17) and nondividing (primary neuron) mammalian cells. Transfection efficiency, as observed by EGFP expression, was observed to be almost 100% after 24 h in rapidly dividing cells and approximately 80% after 48 h in nondividing cells (Cai et al. 2005). More recent work by Jie Chen and colleagues also demonstrated efficient uptake of fluorescein isothiocyanate (FITC)-labeled magnetic carbon nanotubes (mCNTs) into human monocytic leukemia THP-1 cells in vitro (Gul-Uludag et al. 2012). After addition to the cells in culture, samples were exposed to external rotating then static magnetic fields for 10 min. Compared to spherical magnetic nanoparticles, the use of mCNTs resulted in more rapid and efficient uptake. One hundred percent uptake was achieved after 1 h (determined through fluorescence-activated cell sorting) with no decrease in uptake efficiency after 24 and 48 h. Additionally, three-dimensional reconstruction images of confocal microscopy results showed nuclear localization 6 h after delivery with no change in cell viability. The cylindrical shape and magnetically functionalized tip of the mCNTs is thought to allow the mCNT to follow the guidance of an external magnetic field better than spherical magnetic nanoparticles. More favorable properties such as a smaller diameter (~1 nm versus 5–500 nm for nanoparticles) and elongated shape (thought to reduce drag forces) are also thought to enhance delivery. In the future, this sort of method might be used for ex vivo transfection and/or magnetic trafficking to some sites of interest in in vivo administration. For example, it may be possible to deliver mCNTs into monocytes followed by in vivo administration of the monocytes and then apply a clinically appropriate magnetic field to direct the monocytes to a tumor site, thereby using the genetically engineered monocytes as cellular delivery vehicles for cancer gene therapy.

2.9.2 Fullerene

The use of C₆₀ fullerene (Kroto et al. 1985), also known as buckminsterfullerene or buckyball, has resulted in successful delivery of exogenous nucleic acids to cells. Although other forms of fullerene exist, for example, C₇₀, fullerene-mediated gene delivery has focused on the C₆₀ allotrope. Following early investigations of its interactions with DNA pioneered by Nakamura and co-workers, including the ability of fullerene to oxidatively cleave the DNA duplex through photo-excitation (Tokuyama et al. 1993), studies have now demonstrated that fullerene derivatives are capable of oligonucleotide delivery into cells (Nakamura et al. 2000; Klumpp et al. 2007; Sigwalt et al. 2011). For more discussion of early studies regarding these initial fullerene derivatives and their structure-activity relationships for gene delivery, the reader is referred to reports and reviews by Nakamura et al. and other groups (Nakamura et al. 2000; Nakamura and Isobe 2003; Isobe et al. 2006a, b; Bakry et al. 2007).

Like the carbon nanotubes, fullerene requires hydrophilic surface modifications for solubility under physiological conditions as an alternative to the use of solvents which often result in cytotoxicity (Nakamura et al. 2000; Isobe et al. 2006b). The resulting amphiphilic fullerene derivatives possess the ability to self-assemble into vesicles, often referred to as “buckysomes” approximating anywhere from 10 to 200 μM in diameter (Partha et al. 2007), and have been used for drug delivery (Partha et al. 2008). Such modifications also allow for DNA complexation and compaction mainly due to electrostatic and hydrophobic forces (Sitharaman et al. 2008). It is notable, as with carbon nanotubes, that a positive charge is necessary for full DNA interaction and complexation.

A cationic functionalized fullerene, tetra(piperazino)fullerene epoxide (TPFE), shown in Fig. 2.19, formed stable complexes with DNA and demonstrated efficient gene delivery in vivo (Maeda-Mamiya et al. 2010). TPFE-DNA complexes remained under 100 nm in serum-containing buffer and achieved more efficient DNA delivery to the liver and spleen compared to Lipofectin without acute liver or kidney toxicity (Maeda-Mamiya et al. 2010). Previously, this system showed successful transfection in vitro and protection from nucleases and has the potential for DNA release through loss or neutralization of the DNA-binding amino groups (Isobe et al. 2006a; Maeda-Mamiya et al. 2010).

Likewise, water-soluble C₆₀ derivatives prepared via the Hirsch-Bingel reaction, a simple, high-yielding and versatile fullerene functionalization method, all showed significant in vitro plasmid DNA transfection compared to naked DNA (Sitharaman et al. 2008). However, only two cationic fullerene derivatives, octa-amino and dodeca-amino functionalized fullerenes, at their optimal fullerene-to-DNA-base pair ratio achieved transfection efficiencies comparable to the 25 kDa PEI and CytoPure reagent positive controls. In the presence of serum-free DMEM cell culture medium, the hydrodynamic radii (R_h) of both the octa-amino and dodeca-amino derivatives were below 200 nm, a diameter upper limit that has been associated with more efficient cellular uptake. This may speak to the increased capacity for DNA compaction due to an increased density of positive charges on the surface of these particular derivatives. These samples were shown to consist of polydisperse aggregate populations, which may not be correctly represented in the R_h data (Sitharaman et al. 2008). Additionally, these particular formulations resulted in approximately 30% cell viability 24 h post-transfection and cell morphological changes in culture (no negative control data shown). This cytotoxicity is thought to be largely due to the aggregation of the complexes over time and their subsequent precipitation onto the cell surface. The authors postulate that preventing or disassembling aggregates during complex preparation could also reduce this resultant toxicity. It is also worth noting that although such low cell viability appears highly unfavorable at first glance, these results may be acceptable as dependent on their intended use (e.g., cytotoxicity in a cancer cell population would be quite acceptable and, in fact, desirable) and when keeping in mind that it may not necessarily reflect systemic toxicity (Sitharaman et al. 2008).

Polycationic fullerene hexakis-adducts (C₆₀(NH₃⁺)₁₂) (Richardson et al. 2000) satisfy the criteria for DNA complexation and cell penetration as these amphiphilic structures possess both positive charges for DNA interaction and hydrophobic regions for cell membrane interaction. Though previously thought unsuitable for transfection due to the isotropic distribution of positive charges (Isobe et al. 2006b), a series of amphiphilic globular fullerene derivatives was demonstrated to have noteworthy gene delivery capabilities, achieving gene expression comparable to a commercial lipid control (JetSI-ENDO Polyplus-Transfection) (Sigwalt et al. 2011). Fullerene hexakis-adduct derivatives were mixed with DNA, and DNA condensation was allowed to occur in 5% glucose, resulting in compact “donut-like” structures less than 100 nm in diameter. The derivatives were similarly mixed in 150 mM NaCl, resulting in aggregated spherical complexes that were observed to mimic PEI-DNA complexes. Transfection efficiency was measured through luciferase expression from delivery of plasmid pCMV-Luc into HeLa cells. In addition to comparable transfection efficiency to the commercial control, cell toxicity was also low. These highly promising results indicate the potential of fullerene hexakis-adducts as nontoxic and efficient transfection agents (Sigwalt et al. 2011).

2.9.3 Carbon Nanohorns

A variant of single-wall carbon nanotubes, this carbon nanomaterial was first observed after CO₂ laser ablation of graphene sheets and characterized by Iijima et al. (1999) and Bandow et al. (2000). The single-wall carbon “nanohorn” was thusly named due to the resultant cone-shaped morphology of the resultant particles, which were also observed to assemble in larger “dahlia flower”-like aggregates (requiring purification into individual nanohorns thereafter). The dahlia flower aggregates themselves typically have diameters around 100 nm and are comprised of tightly clustered nanohorns bound by van der Waals forces. Interstitial spaces and the hollow inner “horn” spaces in dahlia aggregates can encapsulate therapeutic cargo as is. To demonstrate their capacity for cargo, dahlia structures were shown to carry and deliver vancomycin hydrochloride in a controlled release manner (Xu et al. 2008).

Nanohorns have great potential as drug carriers due to their ease of synthesis and overall high production yield. As they do not require metal catalysis (Iijima et al. 1999) for synthesis, carbon nanohorns were thought to be less toxic. The extensive surface area inside the “horn” facilitates adsorption of drug cargo and protection from degradation during delivery, allowing it to function much like a nanocage. Murata et al. induced the formation of micropores and mesopores in nanohorn structures with heat (>693 K) (Murata et al. 2001). The micro- or meso-porous structure can further improve the adsorption and release of cargo molecules. Nanohorns and functionalized nanohorns were quickly demonstrated to carry a range of drugs including the aforementioned vancomycin, dexamethasone (Murakami et al. 2004), doxorubicin (Murakami et al. 2006), and cisplatin (Ajima et al. 2008). Fan et al. observed localization of nanohorns within the perinuclear regions of HeLa cells 4–7 h after cell uptake, indicating an endocytic mechanism of cell entry (Fan et al. 2007). They failed to observe nanohorn entry into the nuclear compartment, indicating any nuclear-targeted cargo must dissociate from the nanohorn in order to bypass the nuclear membrane. Miyawaki et al. also investigated the toxicity of carbon nanohorns and concluded low acute toxicities over a 90-day period with rat lung tissue (Miyawaki et al. 2008).

The ability of carbon nanohorns to encapsulate and release drug cargo is very promising. Its intracellular localization without undue toxicity further underlines its prospective use as a gene carrier although investigations into its role as such remain scarce. Guerra et al. reported functionalization of carbon nanohorns with PAMAM dendrimers in order to enhance their solubility and cell penetration (Guerra et al. 2014). PAMAM dendrimers were coupled to the nanohorn via an amidation reaction and were marked with gold nanoparticles to enable tracking and visualization of the hybrid complex. Coupling with the carbon nanohorn enables dispersion of the positive charges across the PAMAM surface, which was thought to reduce its overall toxicity. PAMAM-nanohorn delivery of siRNA targeted against p42-MAPK and GADPH (anti-prostate cancer targets) was evaluated in prostate cancer cell lines (PC-3). The hybrid vectors were indeed able to abolish target gene expression without significant toxicity. The synthesis of these vectors lays the groundwork for future nanohorn-encapsulated oligonucleotide delivery applications, though it remains to be seen if dsDNA can also be delivered with similar success.

2.9.4 Graphene

Graphene is formed of two-dimensional planar sheets of carbon atoms arranged in a honeycomb structure. With its large surface area to facilitate efficient loading of therapeutic cargo and excellent mechanical and chemical properties (Ferrari et al. 2015), explorations into graphene biomedical applications have exploded in recent years. The reader is directed to an extensive review by Ferrari et al. for more information on graphene production and applications (Ferrari et al. 2015).

Graphene does not intrinsically bind nucleic acid as its overall negative charge repulses the negative charge of the phosphate backbone. In 2010, Lu et al. demonstrated the ability of graphene oxide (GO) to bind to and protect oligonucleotides from enzymatic degradation (Lu et al. 2010b). Wu et al. further investigated the extent of DNA adsorption and desorption on GO (Wu et al. 2011). They found that shorter oligonucleotide length, lower pH, higher ionic strength, and lower temperature benefited DNA adsorption, while the reverse was true for desorption. Furthermore, single-stranded DNA bound reversibly to GO and could exchange with free DNA in solution. The strong affinity of GO to short oligonucleotides makes it particularly promising for the delivery of DNA/RNA aptamers or probes (Wang et al. 2010, 2013). Graphene derivatives seem able to bind effectively to single-stranded nucleic acids through π-π interactions but tend to demonstrate weaker dsDNA interactions (Tang et al. 2010a, b; Wen et al. 2010). The binding affinity of functionalized graphene for oligonucleotides can be exploited for biosensor and imaging applications.

In terms of graphene-mediated transfection, Feng et al. constructed hybrid polymers consisting of the gold standard PEI and GO (Feng et al. 2011). Functionalization of GO with PEI enriches the complex in positive charges allowing loading of dsDNA. The GO-PEI complex exhibited greatly reduced cytotoxicity in comparison with bare PEI polymers, even with very high molecular weight PEI (10 kDa), and did not decrease transfection efficiency. Addition of PEG to PEI-GO hybrid vectors further improved transfection efficiency (Feng et al. 2013). PEI-GO conjugates were also further transfected into zebrafish embryos with sustained gene transfection up to 72 h (Zhou et al. 2012). Zhou et al. concluded that reduced free PEI and smaller GO loads (“ultrasmall” GO with diameters approximately 10–20 nm) contributed to the high transfection efficiency. Cationic graphene oxide complexes could very well lead the way to an overall safer and more efficient non-viral gene therapy.

2.9.5 Conclusions

The family of carbon-based nanomaterials far exceeds what has been described in this chapter. However, extensive investigation into toxicity and bioaccumulation effects is necessary before their use in clinical gene therapy studies. Beyond their utility for gene delivery, carbon nanostructures have potential for innumerable applications as the building blocks for nanodevices and other applications. Overall, carbon structures require functionalization with charged constituents for solubilization in aqueous solution. Aggregation with serum components will be an obstacle to delivery in the future, though it may be mitigated with further complexation with PEG. Thus far, carbon nanotubes, fullerenes, and graphene functionalized with cationic polymers or lipids show great promise for oligonucleotide delivery.

2.10 Localized Hyperthermia

Localized hyperthermia has long served as an effective cancer therapeutic by acting synergistically with chemotherapy and radiotherapy; it has extended the long-term survival of patients suffering from soft tissue sarcoma, melanoma, and recurrent breast/cervical cancers among others (Sanches et al. 2011; Mallory et al. 2016). Localized hyperthermia induces a number of desirable therapeutic effects: direct thermal toxicity, antitumor inflammatory responses, and site-specific, enhanced drug/gene delivery. The application of mild hyperthermia increases blood flow and enhanced vascular permeability for effective delivery of chemotherapeutic agents. In addition, mild hyperthermia may serve to induce antitumor effects upon heat activation of gene products after successful delivery of a gene therapeutic.

Through in vivo delivery of the gene encoding heat shock protein (HSP) HSP70 into B16 melanoma nodules, Ito et al. (2003) have observed upregulated expression of HSP70 that led to antitumor effects after induced hyperthermia. Significant tumor arrest contributed to elevated cytotoxic T-lymphocyte (CTL) response against B16 melanoma in conjunction with prolonged survival/complete tumor regression in three of the ten mice. The combination of HSP70 gene therapy and hyperthermia could be an effective cancer therapeutic, as HSP70 may induce innate immunity by acting as a proinflammatory cytokine and induce adaptive immunity by T cell activation (Ito et al. 2003).

In addition to cancer therapy, the favorable effects of hyperthermia can be applied for gene delivery. Hyperthermia-induced downregulation of VE-cadherin is thought to contribute to rapid and reversible changes in endothelial cell permeability that can therefore enhance extravasation and promote drug/gene delivery (Böhmer et al. 2009; Yudina and Moonen 2012). Hyperthermia can initiate controlled release when used in conjunction with temperature-sensitive drug carriers. Yudina et al. (2012) demonstrated the successful delivery of cell-impermeable compounds in a two-step approach combining sonoporation and hyperthermic controlled release (Fig. 2.20). Encapsulation of cell-impermeable TO-PRO-3 fluorescent dye in a temperature-sensitive liposome (TSL) increased the dye’s half-life and protected it from plasma degradation. Using a two-step mechanism, site-specific delivery was achieved by hyperthermic destabilization of TSL followed by ultrasound-mediated delivery through sonoporation (Fig. 2.20) (Yudina et al. 2012). As ultrasound-induced permeability lasts for prolonged periods, it is feasible to induce sonoporation prior to hyperthermic dye release; this was confirmed in a separate study by Yudina et al. (Yudina et al. 2011) where in vivo administration of TSLs to mice induced a 2.4-fold increase in fluorescence through sequential applications of sonoporation and hyperthermic dye release. Such results hold promise for hyperthermia-induced gene delivery.

Hyperthermic conditions pertinent to temperature-sensitive delivery can be achieved by radiofrequency, thermal effects of ultrasound, magnetic particles, and infrared light (Böhmer et al. 2009; Walther and Stein 2009). Regardless of the technique, the heat source must stably reach the desired temperatures in a temporally and spatially controlled manner for activation of temperature-sensitive carriers (Kim et al. 2012). Focused ultrasound and magnetic nanoparticles are commonly employed for temporal and spatial control of heat-responsive gene therapy. Focused ultrasound is generally combined with MRI for real-time monitoring and feedback control of applied temperatures (Böhmer et al. 2009; Sanches et al. 2011; Yudina and Moonen 2012; Walther and Stein 2009). Magnetic nanoparticles, incorporated into temperature-sensitive carriers, are employed for magnetically induced hyperthermia by an alternating magnetic field; these magnetic nanoparticles continuously emit heat through the Neel and Brownian relaxation pathways (Yoo et al. 2011). Focused ultrasound and magnetic nanoparticles are effective means of inducing hyperthermia, but they must be closely regulated such that hyperthermia can be induced without damaging surrounding tissues.

Temperature-sensitive delivery is achieved through the administration of TSLs or polymers. These liposomes/polymers must be optimized such that effective encapsulation can occur at physiological temperatures while achieving rapid release under hyperthermic conditions. TSLs are liposomes comprised of lipids that exhibit a gel-to-liquid crystalline phase transition at hyperthermic temperatures between 40 and 45 °C; phase transition causes transient pore formation that leads to leaky bilayer and rapid release (Böhmer et al. 2009; Sanches et al. 2011; Shao et al. 2011). Newer formulations, comprised of lysolipids and PEGylated lipids, allow for enhanced circulation and stability at physiological temperatures while exhibiting hyperthermia-induced rapid release. PEGylated lipids provide TSLs with better pharmacokinetic properties that enhance stability and circulation times in vivo, while lysolipids promote transient pore formation for rapid release (Sanches et al. 2011). TSLs are generally employed for effective drug delivery, but they have the potential for gene delivery. Temperature-sensitive polymers (TSP) are polymers that undergo phase transition at temperatures above their lower critical solution temperature (LCST). Below LCST, these polymers are good solvents and stably assemble into micelles. At temperatures above LCST, the polymers undergo changes in conformation, solubility, and hydrophobic/hydrophilic balance; these changes lead to collapse and aggregation of polymers that result in enhanced gene delivery (Böhmer et al. 2009; Shao et al. 2011). In a study by Schwerdt et al. (2008), DNA polyplexes were constructed using PEI covalently bound to temperature-sensitive vinylpyrrolidone and poly(N-isopropylacrylamide) copolymer, and they elicited a 100-fold increase in transfection efficiencies in vitro and a 10-fold increased DNA deposition in vivo under hyperthermic conditions. The increased DNA deposition constitutes significant selective transgene expression. The enhanced transfection efficiencies are the result of phase transition of the copolymers causing polymer aggregation and accumulation of polyplexes at target site. TSPs could be attached to liposomes for the generation of thermosensitive polymer-modified liposomes; such formulations possess improved stability and release kinetics through liposomal destabilization as a result of polymer contraction (Shao et al. 2011). Kaiden et al. (2011) generated a pH and temperature-sensitive polymer-modified lipoplex that destabilized and released its encapsulated nucleic acid cargo under mildly hyperthermic and mildly acidic conditions (Fig. 2.21). Synthesis of these thermosensitive polymer-modified liposomes involved the attachment of hyperbranched poly(glycidols), possessing temperature-sensitive N-isopropylamide and pH-sensitive succinylate (Suc) groups, to stabilize the egg yolk phosphatidylcholine liposomes. The dual-responsive lipoplexes could then undergo heat-induced phase transitions under mildly acidic conditions within endosomes; this resulted in destabilization of the acidic endosome and subsequent release of nucleic acid into the cytosol. Undesirable release into extracellular space was avoided as temperature-induced destabilization was confined within endosomes after endocytosis (Fig. 2.21) (Kaiden et al. 2011). Hence, this has the potential of overcoming endosomal escape, a significant barrier to gene delivery. TSPs exhibit high versatility as their LCST can be altered depending on the ratio of hydrophilic and hydrophobic monomers in its composition (Shao et al. 2011); however, the phase transition of TSPs is difficult to control as LCST differs under in vitro and in vivo settings (Böhmer et al. 2009).

In addition to delivery, hyperthermia has been employed to induce transgene expression through the use of heat-responsive promoters. Heat-induced expression is mediated by heat-inducible promoters possessing heat shock elements and the subsequent binding of heat shock factors to those elements (Walther and Stein 2009). The use of endogenous pathways and transcription factors constitutes a one-component inducible system that overcomes several limitations (use of modified transcription factors and slow diffusion rate of inducers) that are present in two-component systems (Ortner et al. 2012). The magnitude of induced expression is dependent on temperature, duration, and cell type. Controllable transgene expression by heat-responsive promoters and hyperthermia can promote synergistic effect as seen in combinatorial cancer therapy through thermal ablation and induced transgene expression of suicide genes/cytokines. However, it is critical to define the correlation between temperature increase and expression such that transgene expression can be effectively controlled. Also, unintentional activation of heat-responsive promoters must be prevented so that induced expression can be deemed safe for implementation in gene therapy.

Heat shock protein 70B (HSP70B), growth arrest and DNA damage 153 gene (GADD153), human multidrug resistance gene 1 (MDR1/ABCB1), and cytomegalovirus immediate early (CMV-IE) are four heat-inducible promoters used in research and clinical applications, with HSP70B promoters as the most commonly used due to its low leakiness and high level of heat-induced transgene expression (Walther and Stein 2009). The addition of consensus heat shock factor protein 1 (HSF-1) binding sequences can optimize induced expression, while the use of a minimal HSP70B cassette ensures heat-specific activation as wild-type HSP70B are responsive to both heat and stress activation (Walther and Stein 2009). HSP70B has been used for induced suicide or cytokine (TNFα or IL-12) expression in cancer therapeutics. Tang et al. observed a 500-fold increase in transgene expression in vitro and higher concentration of gene product after 24 h in vivo from a combination of HSP70B and hyperthermia (Tang et al. 2008). GADD153 promoters are inducible by a variety of stressors including reactive oxygen species (ROS), DNA damage, cytotoxic drug effects, and heat. Through GADD153 promoters, Ito et al. (Ito et al. 2001) demonstrated that combinatorial therapy of hyperthermic ablation and gene delivery has promise as a potential viable cancer therapy; a threefold increase in TNFα was induced in the peripheral areas of a heated tumor environment where hyperthermic conditions were insufficient for direct cell death. Similar to GADD153, MDR1/ABCB1 promoters are stress-inducible promoters possessing heat shock elements that mediate heat-induced expression. CMV-IE promoters have been previously suggested to elevate transgene expression upon heat induction despite the fact that they do not possess any heat shock elements (Walther and Stein 2009). Through CMV-IE promoter activation, Kobelt et al. (2010) have shown the upregulated gene expression corresponding to a 25-fold increase in vitro and 2.5-fold in vivo. The effects of heat-induced activation of CMV-IE were presumed to be due to the activation of specific cellular pathways and transcription factors at a later time after induced hyperthermia (Kobelt et al. 2010).

The use of HSP70 promoter for controlled gene expression was further refined using MRI-guided high-intensity focused ultrasound (HIFU) for noninvasive controlled gene delivery. Deckers et al. (2009) demonstrated this spatiotemporal control of gene delivery and expression in a transgenic mouse model, observing transient gene expression 6–8 h after short thermal induction (approximately 2 min) without significant tissue damage. Increased duration (8 min) of hyperthermia did result in muscle damage, but such long duration was not necessary for gene delivery (Deckers et al. 2009). They later demonstrated this system again (Eker et al. 2011), whereby the gene encoding luciferase (lucF) under control of HSP70B was transfected into C6 cells, which were then injected and implanted into rat kidneys. Gene expression was then activated noninvasively using focused ultrasound, guided using MRI. Again, transient in vivo luciferase expression was observed 6–8 h after heat induction. More recently, the group turned to magneto-activatable thermogenic nanoparticles (MTN) (Sandre et al. 2017), which would allow for magnetic induction of hyperthermia as opposed to ultrasound. This could permit deeper penetration into tissues, such as poorly accessible tumors. MTNs were tested upon application to murine skin or were injected into tumor space for hyperthermia induction; however, in this study, activation of the HSP70 promoter was not stable, or consistent temperature increase was not detectable. This intriguing and innovative combination of technologies has many attractive assets but will require more investigation. Overall, the use of focused ultrasound for spatially targeted hyperthermia serves as a highly useful tool for controlled gene therapy.

HSP70B promoters can also be combined with a modified tetracycline (Tet) controlled expression system to generate a one-component, heat-inducible transgene expression system possessing a positive feedback loop for transcriptional amplification. This modified system, developed by Yamaguichi et al. (2012), is highly suited for hyperthermia-induced gene therapy as it merges the benefits of a tightly controlled, one-component heat-inducible system in HSP70B with transcriptional amplification from the Tet-OFF system (Figs. 2.22 and 2.23). In this system, heat activation of HSP70B promoter drives the initial expression of the Tet-responsive transcriptional activators (tTA); the subsequent binding of tTA to the TRE/PCMVmin promoter, situated downstream of the HSP70B promoter, constitutes a positive feedback loop for amplification of tTA expression. tTA serves to activate and amplify therapeutic gene expression downstream of a TRE/PCMVmin promoter (see Fig. 2.22) (Yamaguchi et al. 2012). Using this system, a 3000- and 6500-fold increase in expression was observed 48 h after heat induction at 43 °C and 45 °C, respectively; however, this required the successful co-transfection of both plasmids into the same cell in order for inducible expression to arise. To simplify the system, a one-pack plasmid was generated such that the expression of both tTA and therapeutic gene could be simultaneously amplified (see Fig. 2.23) (Yamaguchi et al. 2012). The one-pack plasmid was able to yield expression levels higher than the two-plasmid approach; this was mainly attributed to higher transfection efficiencies. The enhanced expression of the one-pack plasmid was confirmed to be due to the positive feedback loop as the addition of doxycycline, a tetracycline analog capable of inhibiting expression in traditional Tet-OFF systems, resulting in a sixfold reduction in expression (Yamaguchi et al. 2012). Findings from this study have serious in vivo implications as high transgene expression can occur at low transfection efficiencies. The authors have successfully applied the system for suicide gene therapy using herpes simplex virus thymidine kinase gene (HSV-tk) as the therapeutic gene; the high levels of cytotoxic effect observed is significantly greater than cytotoxic effects due to heat-induced apoptosis. Thus, the potential for effective combinatorial cancer therapy has been demonstrated.

The application of natural heat-inducible promoters for expression upregulation may be problematic as they are commonly integrated into other emergency pathways not mediated by heat shock factors (HSF-1), which could contribute to tissue-specific variations (Ortner et al. 2012). Such limitations can be bypassed through the generation of artificial heat-inducible promoters. In a study by Ortner et al. (2012), stable transfection and subsequent heat-induced gene expression was observed through the application of an artificial heat shock-inducible promoter consisting of two minimal CMV promoter elements flanking multiple high-affinity heat shock elements (Fig. 2.24). The artificial promoter elicited similar kinetics to natural heat shock promoters, and heat-induced expressions were highly regulated depending on the duration of induced hyperthermia. The use of magnetic nanoparticles and alternating magnetic field for heat generation allows for efficient modulation of gene expression. The authors were able to demonstrate the encapsulation of magnetic nanoparticles and stably transfected HEK293, possessing the aforementioned artificial promoter, into sodium cellulose sulfate (SCS) capsules. The encapsulation of SCS capsules had no effect on cell viability as cells continued to divide within the capsule. Exposure to alternating magnetic field resulted in 1700- and 950-fold increases in luciferase and GFP gene expression in vitro (Sanford et al. 1987). Such results signify the potential of encapsulation systems for cell therapy as they possess several advantages: (1) protection of genetically modified cells from immune activation, (2) cell selection and testing prior to encapsulation for optimized therapeutic gene expression, (3) inducible expression of therapeutic genes, and (4) remote control of induced expression by the application of alternating magnetic field (Ortner et al. 2012).

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Authors and Affiliations

University of Waterloo, School of Pharmacy, Waterloo, ON, Canada
Chi Hong Sum, Samantha Marisha Shortall, Shirley Wong & Shawn David Wettig

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Chi Hong Sum
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Samantha Marisha Shortall
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Shirley Wong
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Shawn David Wettig
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Correspondence to Shawn David Wettig .

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School of Pharmacy, University of Waterloo, Waterloo, Ontario, Canada
Roderick A. Slavcev
School of Pharmacy, University of Waterloo, Waterloo, Ontario, Canada
Shawn Wettig
School of Pharmacy, University of Waterloo, Waterloo, Ontario, Canada
Zhiheng Zeng

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Sum, C.H., Shortall, S.M., Wong, S., Wettig, S.D. (2018). Non-viral Gene Delivery. In: Slavcev, R., Wettig, S., Zeng, Z. (eds) Nanomedicine. Experientia Supplementum, vol 110. Springer, Cham. https://doi.org/10.1007/978-3-319-78259-1_2

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DOI: https://doi.org/10.1007/978-3-319-78259-1_2
Published: 19 July 2018
Publisher Name: Springer, Cham
Print ISBN: 978-3-319-78258-4
Online ISBN: 978-3-319-78259-1
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Non-viral Gene Delivery

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