Bioelectricity for Drug Delivery: The Promise of Cationic Therapeutics

Abstract

Biological systems overwhelmingly comprise charged entities generating electrical activity that can have significant impact on biological structure and function. This intrinsic bio-electrical activity can also be harnessed for overcoming the tissue matrix and cell membrane barriers, which have been outstanding challenges for targeted drug delivery, by using rationally designed cationic carriers. The weak and reversible long-range electrostatic interactions with fixed negatively charged groups facilitate electro-diffusive transport of cationic therapeutics through full-tissue thickness to effectively reach intra-tissue, cellular, and intracellular target sites. This article presents a perspective on the promise of using rationally designed cationic biomaterials in targeted drug delivery, the underlying charge-based mechanisms, and bio-transport phenomena while addressing outstanding concerns around toxicity and methods to mitigate them. We also discuss electrically charged drugs that are currently being evaluated in clinical trials and identify areas of further development that have the potential to usher in new treatments.

Introduction

Our growing understanding of disease pathogenesis at molecular, genetic, cellular, and tissue levels is paving way for the discovery of an increasing number of therapeutic targets. This provides the opportunity and a critical need to develop systems that can enable effective local delivery of biological cargo to target tissue, cellular, and nuclear sites with specificity.¹ Functional delivery systems are increasingly becoming part of the drug development process rather than being viewed as product enhancements.

Biological systems overwhelmingly comprise electrically charged entities, including tissues consisting of fixed negatively charged proteoglycans (PGs), cells with negatively charged glycocalyx, phospholipid cell and nuclear bilayer membranes, and several types of charged macromolecules, proteins, amino acids, and electrolytes. The fixed electrically charged regions within a tissue or cell generate short-range local electric potentials across boundaries of adjacent tissues (or within tissues) of varying chemical composition or across the cell membrane.² These potential differences, existing over nano-length scales, produce strong local electrical fields, and the resulting transport of mobile charges creates electric currents in the microenvironment. For example, bioelectric resting potential across a cell membrane is about 50 mV.³ Cartilage tissue contains a high density of negatively charged aggrecan, comprising of sulfated glycosaminoglycans (GAGs), that generates a similar magnitude of electric potential drop at the synovial fluid-cartilage interface over nanometer length scales, creating strong inward pointing electric fields of 10⁶–10⁷ V/m, equivalent to that required for dielectric breakdown of air.²

These bio-electrical phenomena have very significant impacts on both biological structure and function. For example, negatively charged PGs in tissues provide hydration, swelling pressures, compressive stiffness, and signal transduction. Aggrecans in cartilage hold water and resist compressive loading due to electrostatic repulsion between negative charges as they come closer. Small PGs such as decorin, biglycan, fibromodulin, and lumican are closely associated with collagen fibrils and are believed to regulate collagen fibril diameter.⁴ At the cell membrane, PGs stabilize ligand-receptor binding, triggering signaling complexes that regulate cell proliferation, migration, matrix adhesion, and endocytosis.⁴ Disruptions to this micro-electrical environment can, therefore, impair active transport of biomolecules (both direction and speed), protein conformational changes, signaling pathways, and downstream transcriptional changes, leading to complex degenerative diseases.⁵

This intrinsic bio-electrical activity can also be harnessed for overcoming the tissue matrix and cell membrane barriers, which have been outstanding challenges for targeted drug delivery, by using rationally designed cationic peptide carriers (CPCs), protein-, lipid-, and polymer-based drug carriers.^6–9 Their weak and reversible long-range electrostatic interactions (that persist over nanometer length scales)^10,11 with fixed negatively charged groups facilitate their electro-diffusive transport through the full-tissue thickness to reach intra-tissue and intracellular target sites. While incorporating positive charges is the underlying design mechanism, carrier specificity can be further tuned based on the net fixed charge density (FCD) of the target tissue, state of disease, where cell targets reside, and the physicochemical properties of the relevant drug.⁶ Recent advances have brought the potential of bioelectricity in targeted drug delivery to the forefront, and increased research and clinical interest in the field as in the case of cell penetrating peptides (CPPs) that interact with negatively charged GAG moieties on the plasma membrane to enhance their uptake across the membrane barrier.

Here, we review the mechanisms underlying charge-based bio-transport and their applications to achieve multilevel targeting in the tissue, cellular, and nuclear domains (Fig. 1). We explore research advances and discuss outstanding questions, including those around toxicity concerns and methods to mitigate them. We also discuss a few charged materials that are being evaluated in clinical trials and identify areas of further development of charged delivery systems that have the potential to usher in treatments for a set of largely unaddressed diseases.

FIG. 1.

Cationic carrier or carrier-drug conjugates can interact with negatively charged fixed intra-tissue and cellular moieties, enabling multilevel drug targeting at tissue, cellular, and intracellular levels. (A) Cationic carriers can rapidly penetrate the tissue in high concentrations owing to electrostatic interactions with the high density of negatively charged matrix GAGs, after which they can successively penetrate (B) the cell and (C) the nucleus to reach their target sites via charge-mediated pathways. GAGs, glycosaminoglycans.

Cationic Biomaterials for Drug Delivery at the Tissue, Cellular, and Intracellular Level

Tissue targeting

There exist a number of dense, negatively charged tissues in the body such as the musculoskeletal tissues, including cartilage, meniscus, and intervertebral disks (IVDs), vitreous humor of the eye and mucosal membrane that remain an outstanding challenge in the field of drug delivery. Their degeneration is associated with several common diseases, such as osteoarthritis, back pain, macular degeneration, inflammatory bowel disease, etc., that affect millions of people worldwide. Several of these tissues are also avascular, rendering systemic methods of drug delivery ineffective in reaching intra-tissue cellular sites. In addition, the complex extracellular matrix of such avascular tissues consists of a dense meshwork of collagen fibrils, PGs with highly negatively charged GAG chains, and cells, which hinders transport of locally injected macromolecules, especially if they are anionic.

Despite the existence of chemical entities and biologics that have been shown to inhibit or reverse disease progression in preclinical studies, many degenerative diseases affecting these tissues remain untreated due to a lack of effective drug delivery methods.¹² Targeted therapies that can penetrate through the full thickness of these tissues to reach their matrix, cellular, and intracellular target sites and provide sustained drug release can enable the clinical translation of these drugs and address several untreated diseases of avascular, negatively charged tissues.

Recent research has shown that the high negative FCD of tissues provides an opportunity, rather than a challenge, for targeted drug delivery by modifying therapeutics to add optimally charged cationic domains, thus utilizing the weak, reversible nature of electrostatic interactions to enhance their intra-tissue transport, uptake, and retention.^2,6,13,14 The concentration of cationic carriers administered locally increases sharply at the interface of negatively charged tissues by a Donnan partitioning factor, K, due to the internal electrical field created by the negative FCD of the tissue (Fig. 2A).² The enhanced interface concentration of cationic carriers (KC versus C in the bulk bath) induces steep concentration gradients across the tissue thickness, resulting in faster intra-tissue transport rates and higher uptake compared with neutral molecules.² In fact, positively charged carriers have shown to be uptaken up to 400 times more than neutral versions of the same drug carriers in cartilage.^6,7 These cationic carriers also demonstrate increased intra-tissue residence time due to the presence of a high density of negatively charged groups that enable binding with the cationic carriers.

FIG. 2.

(A) Transport mechanism of cationic carriers in negatively charged tissues. Concentration of the cationic carrier is enhanced from C at the administration site to KC at the interface of negatively charged tissue due to Donnan partitioning effect, which increases the intra-tissue concentration gradients, resulting in enhanced electro-diffusive transport and high uptakes. Weak-reversible binding interactions enable carriers to penetrate through the full-tissue thickness to effectively reach their cell targets in sufficient doses. Dashed line represents early initial concentration profile of the cationic carrier across the tissue, whereas solid line across the tissue refers to intra-tissue concentration of cationic carriers after equilibrium has been reached. (B) Targeted drug delivery to cartilage by using optimally charged cationic carriers. (i) Confocal images showing depth of penetration of CPCs in cartilage within 24 h. (ii) Charge-dependent equilibrium uptake of CPCs of various charge (between +8 and +20) in cartilage. Figure adapted with permission from Vedadghavami et al.⁶ (iii) Immunohistochemical analysis showing presence of Avidin through the full thickness of rabbit articular cartilage even at 3 weeks post-ACLT, which correlates with the spatial GAG density stained by Safranin-O (red). Figure adapted with permission from Bajpayee et al.²² (C) Targeted drug delivery to ocular tissues by using an electrostatically formed complex of pRFP, PAMAM, and penetratin peptide. Fluorescent images of rat retina 4 h after topical administration of the complexes (blue, green, and red represent cell nuclei, penetratin, and pRFP, respectively). Figure adapted with permission from Liu et al.²⁵ (D) Drug delivery across mucosal barrier. Insulin-encapsulated PLGA nanoparticles coupled with cationic R8 and negatively charged Pho, capable of dissociation through digestion with IAP and resulting in a net positive charge. R8 and Pho were coupled to PLGA by using a DSPE-PEG bridge. Fluorescent images of rat intestinal villi post-oral administration (white arrows show absorption of particles in the interior of intestinal villi). Figure adapted with permission from Wu et al.²⁹ ACLT, anterior cruciate ligament transected; CPCs, cationic peptide carriers; DAPI, 4,6-diamidino-2-phenylindole; DSPE-PEG, 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[amino-(poly{ethylene glycol})]; FAM, 5-carboxyfluorescein; IAP, intestinal alkaline phosphatase; NPs, nanoparticles; PAMAM, poly(amidoamine); Pho, phosphoserine; PLGA, poly(lactic-co-glycolic-acid); pRFP, red fluorescent protein plasmid.

Although the positive charge of solutes facilitates tissue penetration in dense, negatively charged, avascular tissues such as cartilage, it has been observed that the effectiveness of tissue penetration is highly dependent on the magnitude of net charge.^6,15 Mechanistic transport studies using peptides with different net charges across tissues of varying FCD (cartilage, vitreous of eye, mucin, skin) have shown that an optimal charge can be determined for a peptide carrier to take advantage of Donnan partitioning-induced enhanced transport such that the electrostatic interactions are weak enough so the carriers can rapidly penetrate the tissue superficial zones but strong enough to bind within the deep zones for long-term retention.^6,14,16,17

Conversely, carriers with excessive high net positive charge can get trapped at the tissue surface because of the strong binding that will hamper their ability to reversibly bind and reach target sites.^6,16 Recently, the effect of net positive charge on the transport of arginine-rich short-length CPCs of varying net charge (between +8 and +20) was investigated in cartilage.⁶ The depth of penetration of fluorescently labeled CPCs in cartilage was determined by using confocal microscopy. It was noted that CPC +8 rapidly penetrated through full thickness of cartilage due to weak and reversible charge interactions with negatively charged cartilage aggrecans. However, CPC +20 was halted within the cartilage superficial zone because of strong binding resulting from its high net positive charge, impeding its further intra-tissue diffusion (Fig. 2B-i).⁶ In addition, equilibrium uptake of CPCs in cartilage was measured 24 h after submerging cartilage explants in baths of fluorescently labeled CPCs. The highest uptake was demonstrated by CPC +14 while still maintaining full-tissue thickness penetration (Fig. 2B-ii).⁶

Transport properties of cationic peptides of varying net charge (between +7 and +23) were also studied in the vitreous humor of the eye, which comprises a meshwork of negatively charged hyaluronic acid. It was discovered that peptides of up to +15 charge effectively diffused through the vitreous humor while increasing the net charge beyond this threshold to +19 and +23, resulting in their complete immobilization in the vitreous humor.¹⁶ The effect of length of polyarginine in enhancing penetration of nanostructured lipid carriers across the skin was also investigated by using polyarginines of varying length between 8 and 15 residues. Nanoparticles (NPs) modified with 11-arginine residues (R11) showed the highest intra-skin penetration compared with ones modified by the other peptides.¹⁷ Therefore, modulating the net positive charge of a carrier depending on target tissue FCD is an important design criterion for effective tissue targeting, penetration, and uptake.

Cationic carriers are increasingly being utilized for targeted drug delivery to negatively charged tissues. Avidin, a positively charged glycoprotein, due to its optimal size and net charge, demonstrated 200 × higher intra-cartilage uptake and rapid penetration through the full thickness of cartilage after its intra-articular administration into the knee joints of rats and rabbits.^18–21 Avidin, when conjugated with an anti-inflammatory drug, dexamethasone (Dex), was found to be present in the full thickness of knee articular cartilage of anterior cruciate ligament transected (ACLT) rabbits even after 2 weeks following its intra-articular injection (Fig. 2B-iii). A single low dose of this cationic conjugate effectively suppressed ACLT-induced joint inflammation and catabolic activity through sustained Dex release from its intra-cartilage drug depot over 3 weeks; its efficacy was significantly greater than the unmodified Dex.²² Recently, multiarm Avidin nano-constructs with enhanced drug loading content have also been developed.^23–25

Similarly, poly(amidoamine) (PAMAM) dendrimers were used to locally deliver insulin like growth factor 1 (IGF-1) to rat cartilage.⁹ Sixth-generation PAMAM dendrimers were partially PEGylated to minimize NP's cytotoxic response. The amines were covalently conjugated to IGF-1, and the remaining amine groups on PAMAM provided the positive charge to facilitate electrostatic interactions with the cartilage matrix. PAMAM conjugation significantly enhanced joint residence time of IGF-1; a single intra-articular injection of PAMAM–IGF-1 provided 30.4 days of drug at therapeutic concentrations, whereas the free IGF-1 injection only provided 2.9 days of effective therapeutic response. Cartilage targeting using cationic carriers via the systemic route has also been explored. Recently, cationic cysteine-dense peptides (CDPs) with at least three reducible disulfide bonds were shown to accumulate in cartilage and IVDs at a higher concentration compared with blood, muscle, and liver after its systemic administration in rats whereas uncharged neutral CDPs did not, further reinforcing the importance of charge interactions in targeting negatively charged tissues.²⁶ The cysteine-knotted folds preserved the peptide from proteolytic degradation. The CDP was then conjugated to triamcinolone acetonide by using a hydrolysable dimethyl adipic acid for selective targeting to joint tissues on intravenous injection. The conjugate accumulated at 11 × and 33 × higher concentrations in knee cartilage and IVDs, respectively, compared with the free drug and significantly reduced the ankle diameter in a rat arthritic joint within 24 hours after the first dosing.

The first identified CPP, trans-acting transcriptional activator (TAT), isolated from HIV, has been applied to a number of drug delivery systems since its discovery. This cationic CPP has been used to enhance penetration of PEG-poly(lactic-co-glycolic-acid) (PLGA) NPs through the ocular barrier after topical administration.²⁷ The NP was further modified with an integrin α_vβ₃ specific binding peptide to selectively target choroidal neovascularization (CNV). The dual modified NPs demonstrated superior corneal penetration (5.5 × higher than unmodified NPs) and specific targeting toward CNV. A gene delivery system using a combination of 3rd-generation PAMAM and penetratin peptide was developed for reaching posterior eye tissues.²⁸ Complexes for gene delivery were formed by using the anionic red fluorescent protein plasmid (pRFP) gene model with PAMAM and penetratin (P) (pRFP/PAMAM/P) using electrostatic interactions and were administered topically in rat eyes. Free pRFP failed to reach the retina 4 hours post topical administration, whereas pRFP/P complex was able to penetrate into the retina. Further, the ultimate pRFP/PAMAM/P complex demonstrated higher accumulation in the retina compared with pRFP/P (Fig. 2C).

In addition, PLGA NPs capable of reversing their net charge from net neutral to positive have been developed to overcome both mucosal and epithelium barriers to enable oral delivery of insulin.²⁹ Insulin encapsulated PLGA NPs were functionalized with octa-arginine (R8) and dissociable negatively charged phosphoserine (Pho) to form PLGA-R8-Pho NPs. The initial neutral charge of the systems allows rapid penetration through negatively charged mucus barrier. Thereafter, in presence of the epithelium, intestinal alkaline phosphatase cleaves the negatively charged Pho groups, giving the NP an overall positive charged that enhances epithelium uptake of the NPs. This charged-reversal strategy significantly enhanced penetration of PLGA-R8-Pho through the mucosal and epithelium barriers into the bloodstream compared with PLGA NPs (Fig. 2D).

Cell targeting

The negatively charged, GAG-containing glycocalyx and the anionic phospholipid bilayer of cell membranes serve as the first barriers to cellular entry by especially repelling negatively charged, hydrophobic, or large-molecular-weight solutes. This leads to poor intracellular targeting and uptake of therapeutics.³⁰ Cationic CPPs utilize electrostatic interactions to overcome the cell membrane barrier and facilitate delivery of various cargoes to the cytoplasm and nuclear sites via direct penetration, endocytosis, or via transitory modification of structure to form inverted micelles (Fig. 3).³¹ Cationic solutes can interact electrostatically with the negatively charged plasma membrane, temporarily disrupting it and forming transient pores to facilitate direct penetration and cellular uptake.^32,33 Peptides with amphipathic properties (containing both hydrophobic and charged regions) can help temporarily disrupt the bilayer to further promote direct penetration into the cell.³⁴ CPPs, particularly regions with polyarginine residues, exhibit strong binding affinity to negatively charged PGs, especially with heparin sulfates in the glycocalyx, leading to aggregation of these drug carriers in high concentrations, facilitating rapid cellular uptake via receptor-mediated or nonspecific endocytosis.^35–37

FIG. 3.

Mechanisms facilitating cell membrane translocation of cell-penetrating peptides: (A) direct penetration, (B) receptor-mediated or nonspecific endocytosis, and (C) formation of transitory structures such as inverted micelles.

For example, internalization of arginine-rich CPPs such as TAT happens through macropinocytosis, a nonspecific form of fluid-phase endocytosis where large wavy cellular extensions trap extracellular solutes and bring them into the interior of the cell.³⁸ Finally, cationic peptides (or other types of carriers) can also achieve membrane translocation through the formation of transitory structures within the membrane, specifically inverted micelles. Biologically, micelles are single-layer phospholipid membranes, typically oriented with the hydrophilic phospholipid heads on the outside, associating with the polar solvent, whereas the hydrophobic tails are shielded within the structure. Cationic CPPs associate with negatively charged phospholipids through electrostatic interactions and form an intracellular vesicle with the negatively charged hydrophilic phospholipid heads surrounding the positively charged CPP, and the nonpolar hydrophobic lipid tails pointing outward (inverted micelles), allowing the CPP to remain in a hydrophilic environment as they enter the cell.^39,40

Nucleus targeting

The cell nucleus represents a critical biological target in the delivery of a number of therapeutics. Many anti-cancer drugs exert cytotoxic effects by inhibiting the unwrapping and transcription of DNA, thus preventing replication and leading to cell death.^41–43 In addition, gene therapy and genome editing have the potential to address countless genetically linked diseases and conditions.⁴⁴ Therefore, efficient delivery of therapeutics to the nucleus remains a major area of interest. However, unmodified drugs exhibit relatively low transport efficiency into the nucleus, limiting their therapeutic effect.⁴⁵ As such, several peptide-based drug delivery systems have been developed to specifically target the cell nucleus and subnuclear regions.^46–48 Cationic CPPs display a similar ability to translocate across the nucleus envelope, comprising a two phospholipid bilayer membrane, and several sequences have been identified that display affinity for nuclear and subnuclear targets.⁴⁹

For example, portions of the cationic antimicrobial peptide 18 were shown to exhibit high cellular internalization and accumulation within nuclei. Further, when conjugated to the anti-cancer agent doxorubicin, which interferes with DNA transcription, the CAP18 conjugate showed greater cytotoxicity than the control, indicating successful delivery to the nucleus.⁵⁰ In a similar study, a highly basic nuclear factor kappa-light-chain-enhancer of activated B cells peptide, CB5005, displayed effective cellular internalization and accumulation within the nuclei of human glioma cells and tumor spheroids. In addition, in vivo, CB5005 was conjugated to doxorubicin, resulting in reduced tumor volume and initiating apoptosis in a glioma-bearing mouse model.⁵¹ Cationic co-polypeptides combining various ratios of histidine-rich peptide, a known CPP, and nuclear localization sequences have been developed for gene therapy, demonstrating significant cellular uptake, nuclear localization, and successful delivery of small interfering RNA to the nucleus.⁵²

Undesired Biological Outcomes of Cationic Carriers

Despite several advantages that cationic biomaterials offer in multilevel drug targeting to penetrate the tissue and successively reach its tissue, cellular, and nuclear target sites, their long-term biological activity largely remains unknown. Concerns around their side-effects associated with the use of high doses persist and need to be addressed.

At the tissue level, high cationic charge of biomaterials can cause charge shielding, reducing the electrostatic repulsion effect between negatively charged PGs and thereby reducing intra-tissue swelling pressures (Fig. 4A).⁵³ This can lead to decreased water content and some loss of PGs, which can, subsequently, affect organization of collagen fibrils, intra- and intermolecular crosslinks, interaction of PGs with extracellular matrix proteins such as collagen, as well as with cells, leading to altered downstream signaling and changes at sub-cellular levels. These effects are, however, highly dose dependent. For example, a cationic glycoprotein, Avidin, an effective drug carrier for intra-cartilage targeting,^7,23,24 was shown to cause GAG loss only at the very high concentration of 100 μM (which is two to three orders of magnitude higher than that used for drug delivery).^18,22 Even at this high dose, it did not cause any chondrotoxicity or changes in biosynthesis rates of proteins and PGs.¹⁸ Similarly, recent work on cartilage penetrating cationic peptides showed small but measurable changes in tissue Donnan osmotic swelling pressures but only at very high peptide concentrations.⁶ The change in swelling pressure was insignificant and did not cause any GAG loss.^6,18 In addition, no effects on cell viability or biosynthesis were observed.

FIG. 4.

Undesired bioactivity of cationic carriers. (A) High concentrations of cationic carriers can shield the intra-tissue electrostatic repulsion between fixed negatively charged groups. Highly positively charged carriers can displace the sodium ions (Na⁺) to create a higher density of positive charges near the fixed negative groups, thereby resulting in a steeper exponential drop of electric potential, ɸ (x) and reducing the spacing between the negatively charged groups (x′ < x) causing osmotic deswelling of tissues. (B) Undesired nonspecific binding of cationic drug carriers with serum proteins and blood cells, resulting in formation of aggregates and hindering their delivery to desired target sites. These aggregates can also disrupt the phospholipid cell membrane and facilitate poration. (C) Excessive pumping of protons to help ionization of amines after intracellular uptake of highly charged cationic carriers results in an excessive inward flow of chloride ions (Cl⁻) and water molecules to maintain charge electroneutrality in the endosome (proton-sponge theory), resulting in osmotic swelling and rupture of the endosome.

Similarly, cationic CPPs such as TAT and penetratin have shown only minor effects on cell proliferation and membrane integrity at concentrations up to 50 μM.⁵⁴ Transcriptomic studies have also indicated insignificant genetic alterations with TAT in a variety of cell lines.^55,56 High doses of octa-arginine, however, were found to upregulate several immune-related genes in a macrophage cell line.⁵⁷ Increasing amphiphilicity of peptides, on the other hand, can cause membrane disruption and changes in cellular metabolism at much lower doses due to the increased ability of highly amphipathic peptides to penetrate the cell membrane.⁵⁸ This principle has also been confirmed by recent proteomics studies.⁵⁹

When administered in the circulatory system or close to the target organ, heavily charged cationic biomaterials can interact with negatively charged serum proteins such as albumin⁶⁰ and red blood cells, precipitate in large clusters, and adhere to the cell surface. This can destabilize the cell membrane and induce immediate toxicity, including eliciting an immune response.⁶¹ Once out of circulation, if used in high doses, cationic drug carriers can aggregate in cell glycocalyx regions with high concentrations of PGs, causing poration of cellular membranes (Fig. 4B).⁶² In addition, opsonization of cationic liposomes by serum proteins increases their uptake by macrophages of the reticuloendothelial system (mainly in liver and spleen), resulting in rapid clearance of the cationic liposomes from blood circulation post-systemic administration.^63,64 PEGylation has been used for enhancing circulation time of cationic liposomes by providing a hydrophilic shield surrounding the liposome, which prevents opsonization and clearance of the cationic liposomes.^65,66

After cellular uptake and on reaching the lysosomal compartment, the amine group on the material surface could also perturb proton pump activity.⁶⁷ The proton sponge theory suggests that the high H⁺ buffering capacity of titratable amines could interfere in endosomal acidification, which can lead to exaggerated proton pump activity, osmotic swelling, and, ultimately, the rupture of the endosome (Fig. 4C), causing a reduced rate of mitosis and vacuolization of the cytoplasm.⁶⁸ A recent study showed that high doses of positively charged liposomes and nanocarriers synthesized from polymer polyethyleneimine and chitosan competitively interacted with Na⁺/K⁺-ATPase cationic binding sites, resulting in overload of Na⁺/K⁺ in cells and causing cytoplasm swelling. This resulted in cell necrosis and leakage of mitochondrial DNA and other damage-associated molecular patterns, triggering severe inflammatory response.⁶⁹

Modification Methods to Overcome Toxicity Concerns and Improve Target Specificity

In addition to controlling the dose, the net charge and its spatial distribution on a biomaterial must also be optimized to ensure that targeting benefits outweigh toxicity concerns. In the past few years, new modification techniques for commonly used cationic biomaterials and delivery systems have been devised to overcome this problem. For example, PAMAM dendrimers have shown generation, surface charge, and concentration-dependent toxicity.^9,70 To overcome this, PEGylation has been used as a strategy to partially shield cationic charge to minimize cytotoxic effects while not hindering its tissue or cell targeting effectiveness.^9,23,24,71 Other methods utilize the incorporation of acid-labile linkers within large-molecular-weight polymers to allow for rapid degradation into low-molecular-weight structures in acidic endosomes, which has shown to significantly reduce toxicity.⁷² Disulfide bonds can also be incorporated in polymer or peptide designs to enable easy cellular reduction and enhance cellular transfection while minimizing cytotoxicity.⁷³ To avoid cationic carriers from associating with negatively charged serum proteins, spatially distributing the positive charge along the peptide sequence has been investigated. This approach involves using neutral amino acids as spacers without affecting the net charge, or importing a heterocyclic ring such as imidazolium, guanidinium, or pyridinium as a substitution for the linear amine head group in vector design to help spread the positive charge of the headgroup.^61,74

Charge-based binding can be further stabilized by utilizing synergistic effects of short-range hydrogen bonds and hydrophobic interactions that are strongest over angstrom length scales, as demonstrated recently by arginine-rich cartilage-penetrating peptide carriers that showed enhanced retention in arthritic cartilage with significantly reduced negative FCD.^6,75 The specificity of cationic carriers can be further improved by incorporating specific cell receptors or matrix binding motifs.^76–78 Their cellular uptake and delivery of drug can be controlled by using activatable linkers that are responsive to their microenvironment. This approach has mainly been used in tumor-affected tissues where the microenvironment has upregulated proteases, acidic pH, lower membrane potential, and hypoxia, including in a breast cancer model where the cleavage of the linker allowed the activation of the polycationic CPP.^79–81 For example, a 54 amino acid, lysine-rich CPP known as KRP was screened out of phage display for its high affinity to CD105, or endoglin, an accessory receptor for transforming growth factor-β that is upregulated in rapidly proliferating endothelial cells. KRP's hydrophilicity, basicity, large molar mass, and affinity for cancerous endothelial cells allow for selective targeting of the tumor and its acidic microenvironment, and improved cellular uptake due to the enhanced permeation and retention effect. KRP-doxorubicin conjugates were shown to exhibit significantly stronger tumor selectively, cellular internalization, and anti-tumor properties in human osteosarcoma cells compared with a noncancerous control cell line.⁸²

In a second study, a novel, polyarginine CPP named RIPL was rationally designed as a homing peptide for hepsin, an extracellular protease overexpressed in numerous cancer cells. By exploiting the upregulation of proteases, acidic pH, and lower membrane potential of the tumor microenvironment, cationic RIPL-liposomes were able to effectively target and enhance intracellular drug delivery to hepsin-expressing cell lines.⁸³ In addition, RIPL-liposome-docetaxel conjugates displayed significantly greater intracellular drug delivery in hepsin-expressing human ovarian and prostate cancer cells, while inhibiting tumor growth in an in vivo model of ovarian cancer.⁸⁴ These studies indicate that CPP sequences can be fine-tuned to target specific tumor cell lines by harnessing the properties of the tumor microenvironment, displaying their potential for delivering chemotherapeutics with higher specificity.

Clinical Translation

The promising preclinical data highlighting the successful application of cationic carriers in targeted therapies have resulted in their clinical translation for treatment of various diseases (Table 1). Cationic liposomes have gained attention as a promising candidate for targeted drug delivery for cancer treatment due to their preferential binding with angiogenic blood vessels compared with normal blood vessels.⁸⁵ The lack of glycocalyx barrier on the surface of tumor endothelia exposes the negatively charged phospholipid bilayer of the epithelium, facilitating preferential binding and internalization of the cationic liposomes through electrostatic interactions.⁸⁶ In addition, the acidic microenvironment of the tumors⁸⁷ can further ionize cationic liposomes, making them more positively charged and resulting in enhanced selective uptake.

Table 1.

Cationic Carriers in Clinical Trials

Compound	Cationic carrier	Therapeutic cargo	Company	ClinicalTrials.gov Identifier
EndoTAG™-1	Cationic liposome	Paclitaxel	SynCore Biotechnology	NCT00377936
				NCT00448305
				NCT00542048
				NCT01537536
				NCT03002103
				NCT03126435
SGT-53	Cationic liposome	p53	SynerGene Therapeutics	NCT00470613
				NCT02340117
				NCT02340156
				NCT02354547
KAI-9803	TAT	δ-protein kinase C inhibitor peptide	KAI Pharmaceuticals	NCT00093197
				NCT00785954
XG-102	TAT	c-Jun N-terminal kinase inhibitor peptide	Xigen	NCT01570205
				NCT02235272
				NCT02508337
AM-111	TAT	c-Jun N-terminal kinase inhibitor peptide	Auris Medical	NCT00802425
				NCT02561091
				NCT02809118
RT001	TAT derivative	Botulinum toxin type A	Revance Therapeutics	NCT00968942
				NCT00888914
				NCT01064518
				NCT01124552
PsorBan	R7	Cyclosporine A	Cell Gate	N/A
AZX100	PTD-4	AZX100	Capstone Therapeutics	NCT00811577
				NCT00825916
				NCT00892723

PTD, protein transduction domain; TAT, trans-acting transcriptional activator.

EndoTAG™-1 (SynCore Biotechnology) is a cationic liposomal formulation encapsulating the chemotherapeutic paclitaxel, which has advanced to clinical studies for targeted cancer treatment.⁸⁸ By exploiting the negatively charged tumor microenvironment, these positively charged drug carriers bind preferentially to actively dividing endothelial cells and deliver chemotherapeutics that prevent angiogenesis and inhibit tumor growth. In a phase I trial, EndoTAG-1 was generally well tolerated.⁸⁹ Combined administration of EndoTAG-1 and a standard chemotherapy drug increased the survival rate of patients with advanced pancreatic cancer⁸⁶ and advanced triple-negative breast cancer (TNBC)⁹⁰ in separate randomized controlled phase II clinical trials. Efficacy of EndoTAG-1 combined with other chemotherapy agents in patients with visceral metastatic TNBC is currently under investigation in a phase III clinical trial.⁹¹

A similar cationic liposomal drug carrier for cancer therapeutics, SGT-53 (SynerGene Therapeutics), encapsulates wild-type p53 and features surface transferrin receptor antibodies as a targeting moiety. Loss of the tumor suppressor gene p53 is implicated in a vast majority of human cancers; therefore, this gene therapy technique attempts to restore function of p53, which improves tumor sensitivity to radiation and chemotherapeutics.⁹² Transferrin receptor levels are correlated to cell proliferation rate and are greater in actively dividing cancer cells. Cationic SGT-53 actively targets tumors undergoing angiogenesis by harnessing the overexpression of transferrin receptors, basic tumor microenvironment, and decreased membrane potential to achieve intracellular delivery of p53. In a phase Ib trial, combined SGT-53 and docetaxel was well tolerated after systematic administration, with only 2 out of 14 patients experiencing the dose-limiting toxicity of neutropenia.⁹³ Three patients experienced partial response with significant tumor shrinkage, whereas another two patients showed no further disease progression.⁹⁴

Several cationic peptide-based drug delivery systems, mainly consisting of the TAT peptide, have also advanced to clinical studies. KAI-9803 (KAI Pharmaceuticals), a δ-protein kinase C inhibitor peptide conjugated to TAT peptide, was used to restore blood flow after myocardial infarction.^95,96 In vivo studies showed that TAT peptide enabled enhanced, rapid penetration of KAI-9803 in tissues such as heart, kidney, and liver on intravenous injection in rats compared with the cargo peptide and also showed superior internalization in cardiomyocyte cells in vitro.⁹⁷ In a phase I/II trial, KAI-9803 demonstrated a safe and tolerable profile in a dose-escalated study when injected intra-coronarily during angioplasty.⁹⁶ Although the study suggested improvement trends compared with the concurrent placebo, the study was not powered to conclude efficacy of the treatment using biomarkers.⁹⁸ In a follow-up phase II clinical trial, KAI-9803 failed to reduce infarct size in patients with ST elevation myocardial infarction through intravenous injection.⁹⁹ It remains unclear whether change in route of administration, choice of patients, sample size, or ineffectiveness of the treatment contributed to the contradictory results.

XG-102 (Xigen), a TAT coupled to a c-Jun N-terminal kinase inhibitory peptide, was designed with the goal of reducing ocular inflammation. XG-102 has shown to accumulate in various parts of ocular tissues, including iris epithelium, retinal pigment epithelium, and neural retina, upon intravitreal injection.¹⁰⁰ Safety and tolerability of XG-102 was confirmed by using intravenous injection in a dose-escalating phase I clinical trial in healthy individuals¹⁰¹ and upon subconjunctival injection in patients with intraocular inflammation in a phase Ib study.¹⁰² In a phase II randomized, double-blind, parallel-group, controlled, multicenter trial, a single-dose subconjunctival injection of XG-102 was shown to be noninferior to repeated Dex eye drop administration for the treatment of postoperative ocular inflammation within 28 days post-surgery.¹⁰³ The same peptide electrostatically formulated in a hyaluronic acid gel has also been investigated for prevention of acute inner ear hearing loss under AM-111 (Auris Medical), which has received Fast Track Designation by the Food and Drug Administration.¹⁰⁴ The TAT peptide enables accumulation of AM-111 in cochlear hair cells and neurons, whereas the therapeutic cargo prevents the cell apoptosis initiated by extreme noise exposure.¹⁰⁵ Intratympanic administration of AM-111 demonstrated safety and promising improvement in hearing of patients with severe-to-profound acute sensorineural hearing loss in a double-blind, randomized, placebo-controlled phase II study.¹⁰⁶ Further, in a phase III study, significant treatment effect of AM-111 was confirmed in patients with profound idiopathic sudden sensorineural hearing loss.¹⁰⁷

RT001 (Revance Therapeutics, Inc.), a topical gel formulation for the treatment of facial wrinkles, was able to reach phase III clinical trials.¹⁰⁸ RT001 is composed of 150 kDa botulinum toxin type A (BoNTA) and a cationic peptide that is a reverse sequence of basic residues of TAT.¹⁰⁹ The peptide electrostatically binds to BoNTA such that its protein transduction domains (PTDs) are directed outward to attach to cell membranes.¹¹⁰ Phase II clinical trials were promising; however, in phase III clinical trials, RT001 did not achieve its co-primary endpoints, and as a result, Revance Therapeutics has developed an injectable formulation for future clinical trials.¹¹⁰

Cationic peptides other than TAT have also advanced to clinical studies. PsorBan (Cell Gate, Inc.) consisting of cyclosporine A conjugated to a polyarginine carrier was developed for treatment of psoriasis through topical administration.¹¹¹ Phases I and IIa clinical trials of PsorBan demonstrated effective transdermal penetration of the conjugate and potential benefit in treatment of mild-to-moderate psoriasis.¹¹¹ However, future clinical trials were terminated, as the release rate of the cyclosporine drug was not fast enough to compete with clearance despite efficient uptake.¹¹¹ Capstone Therapeutics developed AZX100 for treatment of dermal scars resulting from surgical operations.^112,113 AZX100 comprises phosphorylated peptide analog of heat shock protein 20 coupled to PTD-4 peptide (YARAAARQARA), an alanine-substituted TAT sequence.^78,114 However, AZX100 was later not pursued owing to financial reasons.¹¹⁵ The safety and tolerability of the cationic carriers as demonstrated in these clinical trials emphasizes that using cationic carriers for drug delivery applications is practical if the dosing needed to induce an effective biological response remains below the cytotoxic concentration of the carrier.

Conclusion

The increasing number of cationic carrier-drug conjugates that are reaching the clinical trial stage is a testament to the promise they hold in translating targeted therapies naturally aided by the overwhelming number of negatively charged entities in our bodies. These negative charges play critical structural and functional roles but can also be harnessed for targeted drug delivery. The associated cytotoxicity concerns can be minimized through rational design of carriers. Methods include incorporating just enough charge on the carrier for effective tissue, cellular, or intracellular drug delivery; incorporating structure modifications that include PEGylation, spatial charge distribution, and the use of reducible or acid labile linkers for easy breakdown; and controlling dose. By utilizing the body's internal electric fields and incorporating rational design, bio-electroceuticals with reversible, tunable, and dynamic properties are becoming an exciting new direction of research and clinical work, with the promise of translating targeted therapies for treatment of several unaddressed diseases.

Authors' Contributions

All authors contributed toward the research, drafting, and revising of the article, and figure creation. The final version of the article was reviewed and approved by all authors. This article has been submitted solely to this journal and is not published, in press, or submitted elsewhere. The responsible authors for the integrity of this article are: C.C.Y., A.V., and A.G.B.

Author Disclosure Statement

No competing financial interests exist.

Funding Information

This work was supported by the United States Department of Defense through the Congressionally Directed Medical Research Programs under contract W81XWH-17-1-0085, National Institutes of Health (NIH) R03 EB025903-1, and NIH Trailblazer R21 EB028385-01.