Bioresorbable polyelectrolytes for smuggling drugs into cells

Abstract

There is ample evidence that biodegradable polyelectrolyte nanocapsules are multifunctional vehicles which can smuggle drugs into cells, and release them upon endogenous activation. A large number of endogenous stimuli have already been tested in vitro, and in vivo research is escalating. Thus, the interest in the design of intelligent polyelectrolyte multilayer (PEM) drug delivery systems is clear. The need of the hour is a systematic translation of PEM-based drug delivery systems from the lab to clinical studies. Reviews on multifarious stimuli that can trigger the release of drugs from such systems already exist. This review summarizes the available literature, with emphasis on the recent progress in PEM-based drug delivery systems that are receptive in the presence of endogenous stimuli, including enzymes, glucose, glutathione, pH, and temperature, and addresses different active and passive drug targeting strategies. Insights into the current knowledge on the diversified endogenous approaches and methodological challenges may bring inspiration to resolve issues that currently bottleneck the successful implementation of polyelectrolytes into the catalog of third-generation drug delivery systems.

Keywords::

• active and passive targeting
• bioencapsulation
• drug delivery
• endogenous stimuli
• layer-by-layer method

Introduction

Over the last decade, the layer-by-layer (LbL) method (Decher and Hong 1991, Decher 1997) has been employed to assemble charged polymers, termed polyelectrolytes, to form polyelectrolyte multilayers (PEMs). This programmed deposition method has opened up exciting new avenues to accomplish multiple key milestones in drug delivery, including, but not limited to, functionalized bioresorbable PEM capsules, with innovative modes of action for treatment against fatal diseases like HIV (De Haes et al. 2010) and cancer (Cho et al. 2014, Kim et al. 2007, Johnston et al. 2012). The LbL method is based on the sequential electrostatic adsorption of charged polyelectrolytes on top of oppositely charged porous or non-porous templates like melamine–formaldehyde (Sukhorukov et al. 2004, Khopade and Caruso 2004), silica nanoparticles (Itoh et al. 2004), gold nanoparticles (AuNPs) (Schneider and Decher 2004, Pereira et al. 2014), calcium carbonate (Ca CO₃) (Yashchenok et al. 2013), carbon and nickel nanotubes (Mayya et al. 2001, Artyukhin et al. 2004), toluene (Tettey et al. 2010), cyclohexane (Khapli et al. 2009), liquid crystal emulsions (Sivakumar et al. 2008), microbes like bacteria (Neu et al. 2001), viruses (Fischlechner et al. 2005), biomolecule-based cells like erythrocytes (Donath et al. 2002, Georgieva et al. 2002, Kreft et al. 2006), and liposomes (Cuomo et al. 2012), followed by the dissolution of the template. For the hierarchical assembly, natural macromolecules like DNA, RNA, poly(L-lysine) (PLL), poly(L-glutamic acid) (PGA), chitosan, cellulose, alginates, hyaluronan (Becker et al. 2010, Saurer et al. 2010, Recksiedler et al. 2006, Lee et al. 2007, Yan et al. 2011a, Antunes et al. 2011, Szarpak et al. 2010), and chemically modified or synthetic polymers like dextran sulfate (DS), polystyrene sulfonic acid (PSS), poly(allylamine hydrochloride) (PAH), poly(acrylic acid) (PAA), and poly(N-isopropylacrylamide) (PNIPAm) (Shu et al. 2010a, Gao et al. 2004, Jian et al. 2013, Wang and Caruso 2006, Cavalieri et al. 2009) are usually employed to form thin films and hollow capsules. Any therapeutically valuable cargo (chemical compound or drug, macromolecules or nanoparticles) can be incorporated into these capsules through “pre-loading” (adsorption of cargo onto template prior to LbL assembly or entrapment) and “post-loading” (encapsulation of drugs into pre-formed capsules) by initiating changes in the permeability of the shell wall via suitable stimuli.

Graphical abstract.

The foremost challenge in designing a smart drug delivery system is to smuggle the drug to the preferred site in the body, and then to activate the release, while protecting it from biological degradation. The method of wrapping drugs within the polyelectrolyte shells offers a wide range of inimitable properties to preserve the therapeutic potency of the molecules and to tune the drug release at the desired site in a controlled style. In contrast to non-biodegradable nanoparticles, polyelectrolytes assembled by the LbL method can be fabricated from bioresorbable materials with a low toxicity profile. The most remarkable advantage of PEMs assembled by LbL in drug delivery are their modularity, and the ability to respond to various environmental stimuli. For example, very recently, Morton et al. demonstrated LbL as a modular approach to develop targeted drug carriers via adsorption of ligand-functionalized, aqueous polyelectrolytes for tissue-specific targeting (Morton et al. 2014). Currently, incorporation or in situ synthesis of metal nanoparticles, especially gold, silver, and magnetic nanoparticles in PEMs, has emerged as a promising platform in the development of light-, ultrasound-, and magnetic-responsive systems (Volodkin et al. 2012, Radziuk et al. 2007, Angelatos et al. 2005, Muñoz Javier et al. 2008, Anandhakumar et al. 2012). A simple protocol for the in situ synthesis, as well as for the incorporation of biologically synthesized silver nanoparticles (AgNPs) in PAH/DS thin films that showed rupture and deformation in the presence of laser light has been demonstrated (Anandhakumar and Raichur 2013, Sripriya et al. 2013). The AgNPs act as an energy absorption center, locally heating up the LbL film and rupturing it under the effect of laser treatment. This novel method has a significant potential in remotely activated drug delivery. Similarly, other physicochemical stimuli like electrical stimulation (Wood et al. 2008), ionic strength (Antipov et al. 2003, Gao et al. 2004), solvents (Lvov et al. 2001), and mechanical deformation (Fernandes et al. 2010) can increase the permeability of the LbL films and would help in opening and closing the PEM capsules for encapsulation and release of drugs. Despite these triggers, biological properties or the presence of biomolecules specific to a disease or organ can be used to activate the release of drugs. In addition, dual or multi stimuli-responsive PEMs that can respond to more than one stimulus have also been reported (Nukolova et al. 2011, Tokuda et al. 2013, Shu et al. 2010b, Chang et al. 2011, Lo et al. 2005, Liang et al. 2011, Costa and Mano 2014).

The process of activating the drug release from PEMs by means of various stimuli can be rationally classified into: endogenous (enzymes, glucose, glutathione, receptors, pH, and temperature) and exogenous (light, ultrasonic, electric, magnetic, ionic, and mechanical) stimuli. The increase in pH and temperature is often associated with several disease conditions offering the ability for self- regulated control over drug release, thereby avoiding the need for external trigger. Hence, the thermal- and pH-responsive PEMs have been grouped under endogenous stimuli. So far, many reports on polyelectrolyte drug delivery systems have been published, in addition to the overwhelming reviews on multifarious stimuli-responsive PEMs that already exist (del Mercato et al. 2014, Deshmukh et al. 2013, Ariga et al. 2011, De Geest et al. 2007, de Villiers et al. 2011, Delcea et al. 2011, Antipina and Sukhorukov, 2011, Yan et al. 2011b, Wohl and Engbersen, 2012, Vergaro et al. 2011, Fleige et al. 2012, De Koker et al. 2012, Pavlov et al. 2013). However, a detailed review on different active and passive targeting strategies is lacking in the literature. Surface functionalization of PEM capsules with ligands offer methods to design drug delivery systems specific to diseased cells. Bioresponsive PEMs are attractive for intracellular drug delivery and have unique potential for in vivo applications, since they do not need any external triggering. The present review gives an overview of the diverse endogenous approaches available to activate the drug discharge from PEMs assembled by the LbL method, and elucidates the recent advances and challenges in this field.

Enzyme-mediated drug release

The incorporation of specific enzymatic substrates as structural components in drug delivery systems would allow a change of material property upon an enzymatic reaction. The over expressed, disease-associated enzymes or the enzyme that is found at higher concentrations at the target site can be utilized to trigger the drug release inside the cell. The wall properties of the PEM capsules can be either disintegrated in the presence of various enzymes like protease, hyaluronidase, chitosanase, and DNAse, or readily altered by varying the enzyme concentration. In the following section, the different kinds of enzyme-degradable polyelectrolyte systems for drug delivery applications are reviewed.

Proteases

Proteases or proteinases include a diverse group of enzymes like trypsin, α-chymotrypsin, pronase, etc. that could cleave or degrade the other proteins by hydrolyzing the protein–peptide bonds. The multifarious roles of proteases in diseases such as cancer, diabetes, neurological disorders, rheumatoid arthritis, cardiovascular conditions, and Alzheimer's disease make the disease-associated proteases an attractive target for drug delivery (Gilmore 2012). Usually, the protease-sensitive drug delivery systems are composed of a protease-sensitive substrate that releases the drug due to the cleavage at the peptide cleavage site in the presence of proteases.

De Geest et al. (2006) were the first to report the intracellular degradation of poly-L-arginine and DS in the presence of the enzyme pronase, spontaneously in VERO-1 cells, after lipid-raft-mediated uptake. Another example for pronase degrading polyelectrolytes was well-designed by Borodina and colleagues for the sustained release of encapsulated DNA. Pronase was captured by micron-sized calcium carbonate particles that were subsequently embedded into poly(L-arginine) and poly(L-aspartic acid) shells. Upon EDTA treatment to extract the calcium carbonate, pronase was released into the capsule's interior and started to digest the surrounding polyelectrolyte shell. They report that the lifetimes of such self-disintegrating capsules could be successfully adjusted to seconds, hours, or days by varying the amount of encapsulated pronase (Borodina et al. 2007).

Apart from pronase, trypsin enzymes are distinctly significant since they can cleave peptides at the carboxyl side of amino acids lysine and arginine. Lee et al. successfully fabricated densely packed trypsin-responsive multilayers, by layering oppositely charged PLL and small interfering RNA (siRNA) on the surface of AuNPs. The multilayers thus prepared, termed as sRAuNPs, whose outer surface layer is PLL, could deliver siRNA into the tumor cells and silence the target gene effectively without toxicity (Lee et al. 2011). Similarly, Krishna et al. demonstrated the synthesis of capsules made of FDA-approved arginine-rich cationic protein, protamine (PRM), and heparin (HEP) onto a CaCO₃ template doped with PSS. The trypsin treatment in about 0.04 mM phosphate-buffered saline (PBS) resulted in the immediate rupturing and sudden release of loaded dextran-tetramethylrhodamine isothiocyanate (TRITC) (Radhakrishnan and Raichur 2012). Figure 1 shows the scanning electron microscopy (SEM) images of loaded microcapsules before and after trypsin treatment. The dual enzyme-responsive hollow microcapsule with arginine and chondroitin sulfate (CS) that disintegrates and releases the drugs upon exposure to either trypsin or in the presence of an endo-b-N-acetylhexosaminidase or hyaluronidase, has also been designed (Radhakrishnan et al. 2013).

Figure 1. CLSM images of TRITC-dextran loaded (PRM/HEP)₂ microcapsules before (A) and after (B) 7 weeks incubation in PBS (pH 7.4) and TRITC-dextran loaded (PRM/HEP)₂ microcapsules after incubation in a 0.04 mM solution of trypsin in PBS (pH 7.4) (C) (the scale bar = 5μm). SEM images of (PRM/HEP) ₂ microcapsules before (D) and after (E) incubation with 0.04mM solution of trypsin in PBS (pH 7.4). Reproduced with permission from reference (Radhakrishnan and Raichur, 2012) Copyright Royal Society of Chemistry 2012.

Ochs and co-workers demonstrated a few modular approaches using PGA modified with alkyne moieties for the design of tailored drug delivery vehicles (Ochs et al. 2008, 2010a, 2010b). In another study, multilayer films of chitosan and pyrene-labeled poly (2-acrylamido-2- methylpropanesulfonic acid) (APy) were destroyed after the incubation in pepsin, due to enzymatic degradation of chitosan and desorption of APy (Wang et al. 2007). One more protease-assisted triggering phenomenon was put forth by Victor et al. by using poly(lactic acid) (PLA) nanoparticles that were enzymatically degraded by a-chymotrypsin (Orozco et al. 2010). Likewise, a range of protease-responsive PEM drug delivery systems have been developed (Rivera-Gil et al. 2009, Ren et al. 2005, Marchenko et al. 2012), while in contrast, protease-protective polyelectrolyte systems have been designed by Woitiski et al. (2009) and Pechenkin et al. (2013) for the protection of encapsulated proteins from proteolytic activity.

In addition to these studies, an intriguing approach was described by Marina and co-workers. They fabricated a thin polypeptide film at a pH of 7.4 with alternating layers of two charged polypeptides, PLL/PGA, on a surface mimicking nonwoven. Complete or near complete degradation of the PLL/PGA nanofilm was obtained when the multilayer was exposed to a bacterial protease V8 glutamyl endopeptidase originating from Staphylococcus aureus, while no degradation of the nanofilm was seen in the presence of human neutrophil elastase (Craig et al. 2012). The results from this study open avenues for products where the wound infection is the trigger for release of bioactive substances from a wound dressing. Thus, PEMs can be potentially tailored to design smart protease-triggered drug delivery systems that could facilitate the release of drugs for the diseases associated with deviant protease enzyme activity.

Hyaluronidase

Hyaluronic acid (HA) is an anionic mucopolysaccharide, occurring naturally in all living organisms. In mammals, the enzymatic degradation of HA results from the action of three types of enzymes—hyaluronidase (hyase), b-d-glucuronidase, and β-N-acetyl-hexosaminidase. HA is non-immunogenic and found throughout the body in various forms (Necas et al. 2008). It is a major ligand for the CD44 receptor, and therefore can be used to target cancer cells that overexpress CD44 glycoproteins (Auzenne et al. 2007) . The hydrophobic anticancer drug paclitaxel (PTX) was successfully loaded in PEM films designed by pre-complexing PTX with a chemically modified derivative of hyaluronic acid(alkylamino hydrazide) containing hydrophobic nanocavities, and subsequent assembly with either PLL or quaternized chitosan as polycations (Boudou et al. 2012). As tumors are known to secrete high levels of hyaluronidase, the PTX-loaded microcapsules were cytotoxic to MDA-MB231 breast cancer cells due to the enzymatic degradation of the microcapsules in contact with cancer cells, leading to a transfer of the PTX from the hydrophobic nanodomains to the lipid membrane of the cells. Similarly, Vodouchê et al. (2006) designed a surface coating based on PLL/HA multilayers that act as a reservoir for PTX.

Szarpak et al. reported PEM capsules by LbL assembly of HA and PAH or PLL that are responsive to the tissue enzyme hyaluronidase (Szarpak et al. 2008, 2010). Both chemically crosslinked HA/PLL as well as HA/PAH capsules become rapidly (i.e., within 2 h) internalized in endosomatic/lysosomatic vesicles upon incubation with in vitro cultured macrophages. Subsequently, fast intracellular opening of the capsule was observed, without the need for external triggering. Figure 2 shows the confocal microscopy studies of cellular uptake of capsules in which the polyelectrolyte shells remain stable outside the cells but readily break open once internalized by cells, suggesting their potential as carriers for intracellular drug delivery. Likewise, Lee et al. (2007) found that the bovine serum albumin (BSA) is released faster from cross linked HA/PLL as soon as hyaluronidase concentration is equal to or greater than 10 U/mL. These studies clearly indicate the perspective to use the HA-based PEM capsules for intracellular drug delivery.

Figure 2. Confocal microscopy images of (A) cross-linked HA/PAH and (B) cross-linked HA/PLL capsules after 2h coincubation with RAW mouse macrophages. Capsules are stained green fluorescent using HA-Fluorescein isothiocyanate (FITC), while the cellular lysosomes are stained using Lyso Tracker Red. The left pane gives the overlay of the green and red channel, the middle pane is the Differential interference contrast (DIC) channel and the right pane is the overlay of green, red and DIC. Colocalization between the green and red channel is observed as a yellow/orange color. Reproduced with permission from reference (Szarpak et al. 2010) Copyright American Chemical Society 2010.

Chitosanase

Chitosanase is an enzyme that hydrolyzes the liner polysaccharide chitosan. Designing LbL systems with chitosan will enable the drug release through multilayer disassembly in the presence of chitosanase. Serizawa et al. (2002, 2006) were the first to demonstrate the hydrolysis of chitosan and DS by cationic chitosanase. Sustained release of the encapsulated proteins was attained via LbL assembly of chitosan and DS on protein-entrapping mesoporous silica particles (Itoh et al. 2006). With increasing time, the chitosan component was degraded by chitosanase, which was explored by using SEM (Figure 3). Recently, Almalik et al. (2013) proved that high molecular chitosan nanoparticles are more prone to degradation by chitosanase than low molecular weight chitosan. Although chitosanase does not subsist in mammalian cells, this approach can be used to design degradable carriers for a large variety of other living organisms that produce chitosanase.

Figure 3. SEM images of the (Chitosan/DS)₃ hollow capsules before (A) and after (B-D) enzymatic degradation at 37°C respectively for 5 min (B, C), and for 24 h (D). Panel C is a higher magnification of panel B. Reproduced with permission from reference (Itoh et al. 2006) Copyright American Chemical Society 2006.

Nucleases

The negative charge on each nucleotide unit in nucleic acids enables the possibility of assembling LbL systems. Such assembled systems can be disintegrated by means of the enzyme nucleases that slice DNA at specific sites. Johnston et al. demonstrated that films and capsules assembled solely from DNA can be engineered to contain restriction-enzyme-cut sites. The capsules degraded as the oligonucleotide in the capsule wall were cleaved specifically in the presence of EcoRI, and released the encapsulated BSA (Johnston et al. 2009). But the utility of these capsules is limited since the restriction enzyme EcoRI is found only in bacteria. Earlier, Serizawa et al. (2003) reported the time-controlled desorption of an LbL assembly of DNA and poly(diallyl dimethylammonium) chloride (PDDA) on the basis of the enzymatic hydrolysis by deoxyribonuclease I (DNase I), which is condensed on the PDDA surface and activated by metal ions in the media.

Redox-responsive polyelectrolyte drug delivery systems

Glutathione is a tripeptide consisting of glutamate, cysteine, and glycine. The concentration of thiol- containing glutathione (GSH) is 100–1000 times higher in intracellular compartments like cytosol, lysosome, and nucleus when compared to that in the extracellular space. Besides, the level of GSH in tumor cells is about seven times higher than in the normal cells (Meng et al. 2009). These facts can be utilized to design GSH-responsive drug delivery systems that can release the drug selectively inside the intracellular space, in the presence of GSH at a concentration of 2–10 mM. In order to introduce redox responsiveness, polymeric systems are functionalized at the terminus of the polymer chains with thiol groups that could enable the formation of disulfide links. When they reach the intracellular compartment in the cell, the high concentration of GSH will cleave the disulfide bond by a redox reaction to free the thiols, thereby releasing the drugs. For example, self-assembled micelles of amphiphilic block copolymers (ABPs) consisting of a pendant disulfide-labeled methacrylate polymer block (PHMssEt) and a hydrophilic poly(ethylene oxide) (PEO) block were investigated (Figure 4). In response to GSH as a cellular trigger, the cleavage of pendant disulfide linkages in hydrophobic PHMssEt blocks of micellar cores caused the destabilization of self-assembled micelles due to the change in the hydrophobic/hydrophilic balance, and led to enhanced intracellular release of the anticancer drug doxorubicin (DOX) (Khorsand et al. 2013).

Figure 4. Illustration of PEO-b-PHMssEt as effective intracellular drug delivery nanocarriers exhibiting enhanced release of DOX in response to GSH in cancer cells. Reproduced with permission from reference (Khorsand et al. 2013) Copyright American Chemical Society 2013.

Similar redox-responsive polymeric LbL systems that contain disulfide crosslinks have been designed by many groups (Li and Haynie 2004, Zelikin et al. 2010, Haynie et al. 2004, Kim et al. 2010, Sun et al. 2013, Zheng et al. 2013, Li et al. 2013). Caruso and colleagues reported comprehensively on redox- responsive systems (Zelikin et al. 2005, 2006, 2007, 2008, De Rose et al. 2008, Sivakumar et al. 2009, Chong et al. 2009a, 2009b, Sexton et al. 2009, Yan et al. 2010, Ng et al. 2011, Becker et al. 2009). They described capsules comprised of a synthetic polymer, poly(methacrylic acid) (PMA), stabilized via biodegradable disulfide linkages (Zelikin et al. 2008). The capsules are obtained via the sequential deposition of thiolated PMA (PMASH) and poly(vinylpyrrolidone) (PVPON) by means of hydrogen bonding onto the template silica particles. The capsules degrade in the presence of intracellular concentrations of a natural reducing agent, GSH. These capsules have been used as carriers for globular proteins (Zelikin et al. 2006), single- and double-stranded DNA of varied lengths (Zelikin et al. 2005, 2007, Ng et al. 2011), drug-loaded oil droplets (Sivakumar et al. 2009), oligopeptides (Chong et al. 2009b), antigens (De Rose et al. 2008), and DOX, to target colorectal cancer cells (CRC) (Yan et al. 2010). Apart from GSH, polycationic and polyanionic poly(ferrocenyl silane)s featuring redox-active ferrocene (Ma et al. 2006, Song et al. 2013a, 2013b) and sulfhydryl-containing reducing agent 1, 4-dithiothreitol (DTT) (Blacklock et al. 2007) have also been used for the preparation of redox-controlled drug delivery systems. The results from these reports advocate the proficiency of the redox-responsive polyelectrolyte drug delivery systems for intracellular drug release, particularly for the rapid release of anticancer drugs.

Temperature-responsive polyelectrolyte systems

Certain polyelectrolytes like PNIPAm respond to the change in their environmental temperature. They can be engaged in drug delivery systems to discharge the drugs at physiological temperature or specifically to the diseased cell associated with the increased temperature. Polymers which become insoluble upon heating have a lower critical solution temperature (LCST), and those that become soluble upon heating have an upper critical solution temperature (UCST) (Schmaljohann 2006). If the polymer solution is heated above the LCST, the polymer chain will dehydrate, become more hydrophobic, and will collapse. This change in the hydrophilic–hydrophobic balance causes the micelles to disassemble and therefore release the encapsulated drug (Fleige et al. 2012). PNIPAm shows inverse temperature-dependent solubility, meaning it is hydrophilic below its lower critical solution temperature (LCST), and it becomes hydrophobic above this temperature (Wohl and Engbersen 2012). Since the PNIPAm is not toxic and its LCST is around 32°C, close to the body temperature, it is an attractive candidate for drug delivery. PNIPAm has been extensively exploited by various research groups to design thermo-sensitive microcapsules and films for biomedical applications (Wei et al. 2006, Yuanpei et al. 2009, Qin et al. 2006, Hales et al. 2004, Chen et al. 2006, Li et al. 2006, Zhao et al. 2011, Steitz et al. 2002, Glinel et al. 2003, Ye et al. 2005, Prevot et al. 2006, Dähne et al. 2001).

Qin et al. prepared thermosensitive thin films of PNIPAm and PAA by the LbL method. The resulting multilayer films ([PAA/PNIPAm]₁₀) were capable of loading and unloading the dye Rhodamine B when the solutions were exposed to elevated temperature (Qin et al. 2006). An interesting approach to treat hyperthermia was put forth by Li et al. (Yuanpei et al. 2009) involving the use of poly (N, N-diethylacrylamide-co-acrylamide)-block-poly (gamma-benzyl L-glutamate). The antitumor drug PTX was encapsulated and it showed thermo-sensitive controlled behavior. The cytotoxicity of PTX-loaded nanoparticles increased obviously under conditions of hyperthermia. Sukhishvili and colleagues constructed a new type of hydrogen-bonded and non-toxic LbL films for repeated on-demand release of drug from an assembly of tannic acid (TA) with the temperature-responsive block copolymer micelles (BCMs), which were pre-formed by heating solutions of a neutral diblock copolymer, poly(N-vinylpyrrolidone)-b-poly(N-isopropylacrylamide) (PVPON-b-PNIPAM), to a temperature above the LCST of PNIPAm (Figure 5A). The model drug, DOX, was incorporated into these films and was retained within the hydrophobic BCM pores at 37°C, whereas exposure to a lower temperature (20°C) triggered fast DOX release, demonstrating a cytotoxic effect on breast cancer cells with more than 90% killing efficacy (Figure 5B). Further, the films maintained their structural integrity in PBS, and each film could be repeatedly loaded with drug and used for more than 15 times with only ∼7% loss in film thickness and no obvious changes in reloading capacity or release profiles (Zhu et al. 2013). Earlier, the same group reported similar hydrogen-bonded temperature-sensitive systems (Zhu and Sukhishvili 2009, Zhuk et al. 2009, Xu et al. 2010, Tan et al. 2011). These studies suggest that the polymeric micelles display highly sensitive swelling/deswelling responses to temperature variations. Moreover, they are highly proficient in entrapping the drugs, and hence can be devoted to design thermosensitive delivery vehicles for pulsated and controlled drug delivery.

Figure 5. (A) Schematic illustration of the layer-by-layer assembly of PVPON-b-PNIPAM BCMs with TA into stable films at a temperature above PNIPAm's LCST. Right: hydrogen bonding formed between TA and BCMs within the LbL films. Reproduced with permission from reference (Zhu et al. 2013) Copyright Elsevier 2013. (B) (a) Cytotoxicity of [BCM/TA]_40.5 films against MCF-7 cells after multiple temperature cycles between 37°C and 20°C. Films were immersed in culture medium at a constant temperature of 37°C, and then cooled down to 20°C for 30 min at cumulative time of 1, 2 and 3 days. After each 37°C/20°C cycling, the medium was separately collected and replaced with fresh medium at 37°C to start a new release cycle. MCF-7 cells cultured overnight were treated with collected medium samples for 24 h and analyzed by MTT assay to determine cell viability. (b) Immunofluorescence images of MCF-7 cells cultured for 24 h. F-actin was stained red with phalloidin-TRITC and cell nuclei stained blue with DAPI. (c) LIVE/DEAD staining of MCF-7 cultured for 24 h. Red = live; green = dead. Reproduced with permission from reference (Zhu et al. 2013) Copyright Elsevier 2013.

Glucose-sensitive polyelectrolytes

Polyelectrolyte multilayer films and microcapsules have been developed to release insulin in response to change in the glucose level in blood. The presence of sugar molecules can either disrupt the linkage between polymeric multilayers or induce change in the permeability, resulting in the formation of small pores (Sato et al. 2011). Sato et al. reported the concanavalin/glucogen LbL systems that are sensitive to sugars (Sato et al. 2004), and later, drug delivery systems that can be fully disintegrated in the presence of D-glucose, D-mannose, and derivatives were also demonstrated (Sato et al. 2004, 2005, 2009). Novel glucose-sensitive polyelectrolyte capsules bearing phenylboronic acid moieties that are able to respond to a stimulus provided by the human body (increased glucose concentration) for the controlled delivery of insulin were reported by De Geest et al. (2006). In this study, the capsules disassembled when exposed to glucose at a concentration of 2.5 mg/L or above, in 5 min, due to the intermolecular repulsion between the sulfonate groups of PSS and the glucose/phenylboronic acid complex. Since the normal glucose concentration does not exceed 1.5 mg/L in healthy human beings, these capsules will remain intact at normal (healthy) glucose levels. Qi et al. successfully constructed glucose-responsive glucose oxidase/catalase multilayer shells by the LbL method onto insulin particles by glutaraldehyde crosslinking, for the controlled release of hypoglycemic drugs. It was observed that the coupled enzymatic reactions of glucose to gluconic acid in the assembled capsules reduced the pH value and led to the enhancement of the shell's permeability, simultaneously increasing the solubility of insulin (Qi et al. 2009b).

In recent years, star polymers have been utilized in a range of drug delivery applications, and their unique solution properties and high functionality make them attractive building blocks for the preparation of nanostructured thin films (Connal et al. 2008). Chen et al. studied self-assembled positively charged poly [2-(dimethylamino) ethyl methacrylate] (PDMAEMA) star polymers and negatively charged insulin and glucose oxidase (GOD). The multilayer films showed faster release of insulin in response to 3 g/L glucose solution than that without glucose, and were found to be non- toxic to mouse fibroblast cells (Chen et al. 2011a). In addition, the same research group successfully fabricated the glucose-sensitive multilayer film by the LbL method with 21-arm poly[2-(dimethylamino)ethyl methacrylate] (star PDMAEMA) and negatively charged insulin and GOD in the form of {(Star PDMAEMA/Insulin)(4) + (Star PDMAEMA/GOD)(4) + Star PDMAEMA} (Chen et al. 2011b). The multilayer film shows an on–off regulation of insulin release in response to stepwise glucose challenge in vitro. Furthermore, the multilayer film could continuously release enough insulin in vivo after being subcutaneously implanted in streptocozotin-induced diabetic rats, and reduced the blood glucose level for at least two weeks. Similarly, a series of insulin-incorporated LbL films with different arrangement sequences of insulin/GOD/star-PDMAEMA were successfully fabricated, and a facile strategy to prolong the drug release time from LbL films was also demonstrated by this group (Chen et al. 2012). The in vivo experiments indicate that the sustained release time of insulin, as well as the hypoglycemic effect, could be prolonged from 17 to 36 days, just by rearranging the assembly sequence of LbL building blocks. Similarly, many research groups have developed LbL shells and films that deliver insulin in response to sugars (Miura et al. 2006, Levy et al. 2008, Ding et al. 2009a, 2009b, Qi et al. 2009a). In addition to these studies, glucose-responsive systems have also been designed to target cancer cells, since the tumor cells accumulate glucose faster than the normal cells. Manna and Patil (2010) fabricated multilayer thin films of poly(vinyl alcohol), borate, and chitosan on colloidal particles (MF particles) to deliver the anticancer drug DOX. After 10 h, 80% of encapsulated DOX was released in the presence of 25 mM glucose solution, at the physiological pH.

pH-responsive polyelectrolyte drug delivery systems

The pH gradient varies inside the human body, especially in a diseased state like cancer in which the pH of the tumor microenvironment is between 6.0 and 7.0. The pH value will drop further, from the extracellular microenvironment of a tumor to intracellular organelles, such as endosomes (pH = 5.5) and lysosomes (pH < 5.5) (Yang et al. 2014). Therefore, these abnormal changes in pH can help in designing pH-sensitive controlled drug delivery systems for cancer treatment. By tuning the pH values, drugs can be either encapsulated or released from PEM capsules via decomposition or enhanced permeability. PEM capsules with weak polyelectrolytes undergo protonation and deprotonation when the pH is altered, relative to the pKa value of the polyelectrolyte in the outer shell. Addition of H + ions or protonation leads to increased permeability because of stronger repulsion, resulting in the swelling of the capsule, whereas deprotonation lowers the shell permeability and results in shrinkage of the polyelectrolyte capsules. Many studies on PEM-based pH-responsive drug delivery systems have been reported earlier, for the delivery of proteins (Li et al. 2012b, Yoshida et al. 2010, Ye et al. 2006), anticancer drugs (Luo et al. 2012, Shen et al. 2013, Minati et al. 2013), genes (Tseng et al. 2013, Li et al. 2012a), and acid-susceptible drugs (Gujarathi et al. 2012). In addition, reversible capsules like PAH/PMA that could open at pH levels < 4 and close at pH levels > 5 (Anandhakumar et al. 2010) have been designed; and carboxymethyl cellulose (CMC)/PAH that open at pH ≤ 6 and close at pH ≥ 7 (Tripathy and Raichur 2013) for protein encapsulation and delivery have also been designed. A detailed literature on PEM-based pH-responsive LbL films and capsules was documented by Sato et al. (2011). These studies evidently indicate the significance of pH-responsive PEMs in drug delivery.

Drug release to target cells by active targeting

As several classes of drug delivery systems lack specificity for diseased cells, the drugs get diffused to all other parts of the body, causing toxicity to the normal cells. In addition, the drug gets partially degraded, thereby reducing its potency. The development of drug delivery systems in response to antibody binding or receptor–ligand interaction can hinder the side-effects of drugs by targeting the diseased cell, and would help in increasing the drugs’ efficacy. In many pathological states, cells express new surface molecules that are absent in healthy cells, or strongly overexpress some of these molecules when compared with cells in a healthy state (Grazú et al. 2012). These molecules encompass antibodies or their fragments, lectins like asialoglycoproteins, lipoproteins, blood plasma glycoproteins like transferrin, peptides, hormones, polysaccharides, and some low molecular weight ligands such as folate. PEM capsules can be surface-functionalized with these molecules and can be directed to an antigen or receptor overexpressed in the diseased cell.

Antibody-mediated drug delivery

One way to improve the immune response is to target the drug carriers specifically to dendritic cells (DC) that present the antigens on the surface to the T cells. Therefore, LbL capsules have been designed such that they can deliver the antigenic cargo directly to the DCs, to evoke potent immune responses (De Geest et al. 2012a, 2012b, De Temmerman et al. 2012, De Koker et al. 2009, De Rose et al. 2008). In vitro and in vivo studies have shown that ovalbumin (OVA)-loaded capsules elicit presentation of capsule antigen by both the major histocompatibility (MHC) complex pathways (MHC I and MHC II) (Sexton et al. 2009, De Geest et al. 2012a). Masereel et al. developed AuNPs with LbL coating of PAA and PSS, followed by surface coating with immunoglobulin (anti-BSA IgG) linking, through an amide bound. This coupling of AuNPs to antibodies targeting surface proteins specifically expressed in cancer cell would be an attractive challenge for the treatment of cancer (Masereel et al. 2011).

The human A33 antigen is expressed by 95% of human colorectal tumor cells. Binding of humanized A33 monoclonal antibody (HuA33 mAb) to the A33 antigen is known to activate cellular internalization mechanisms, which provides a mechanism for delivering the particles inside the cells. Cortez and colleagues successfully biofunctionalized the core/shell particles and capsules made of PAH/PSS by the LbL method, with a HuA33 mAb that can bind to the A33 antigen present on the colorectal cancer cell (CRC) lines (Cortez et al. 2006). They also investigated the influence of size, surface properties, cell line, and kinetic parameters such as dosage (particle concentration) and incubation time on the specific binding of HuA33 mAb-coated LbL particles and capsules to CRC. In these studies, it was observed that HuA33 mAb binds to the A33 antigen present on almost all CRC, and the particles have demonstrated great promise in clinical trials as an immunotherapeutic agent for cancer therapy (Cortez et al. 2007). Later, the same group analyzed the interaction of alkyne-modified poly (N-vinylpyrrolidone) (PVPON_Alk) capsules surface-coated with HuA33 mAb with human CRC (Figure 6A and B) (Johnston et al. 2012). Recently, the targeting of mouse splenic DCs using polymer capsules functionalized with DC-specific antibodies was investigated. The results from this study emphasize that simply selecting a molecule that an antibody can specifically bind to and be internalized by the target cell is not sufficient for effective internalization of a material. Consideration must be made as to whether the process by which the antigenic cargo will be internalized can accommodate the material attached to the targeting antibody (Mintern et al. 2013).

Figure 6. (A) Deconvolution microscopy images of (a) HuA33 or (b) IgG-functionalized PVPON_Alk capsules incubated with LIM1899 CRC cells at 37°C for 24 h. Capsules are labeled with AF647 (red), cells are labeled with Lava Cell (fake colored green), and the nucleus is labeled with Hoechst 33342 (blue). Reproduced with permission from reference (De Rose et al. 2008) Copyright American Chemical Society 2012. (B) Representative imaging flow cytometry (a-c) and deconvolution microscopy (d) images of HuA33 mAb functionalized PVPON_Alk capsules (red) incubated with LIM2405 + CRC cells (green) at 37°C for 4 h. (a) Surface bound capsules. (b) Cells with 1–2 internalized capsules. (c) Cells with multiple internalized capsules. (d) Deconvolution microscopy image to distinguish surface bound capsules (red) from internalized capsules (yellow). Reproduced with permission from reference (De Rose et al. 2008) Copyright American Chemical Society 2012.

Biotin-responsive LbL systems

Avidin is a protein known to strongly bind biotin and analogs (Wilchek and Bayer 1988). Multilayer thin films have been prepared by the LbL method by using avidin as the polycation, and -poly(vinyl sulfate) (PVS) (Anzai and Nakamura 1999), concanavalin A (Con A) (Hoshi et al. 2002), 2-iminobiotin-labeled poly(ethyleneimine) (ib-PEI) (Inoue et al., 2004), PSS (Endo et al. 2011), PSS, PVS, and DS (Anzai et al. 2000), as polyanions on the surface of quartz slides in which avidin retained its binding activity to biotin and its analogs. In the studies demonstrated by Caruso et al. the monoclonal antibodies, HuA33 mAb and epidermal growth factor receptor monoclonal antibody (EGFR mAb) were biotinylated, complexed with neutravidin, and individually coupled to biotin-containing disulfide- stabilized poly (methacrylic acid) capsules. They demonstrated significantly enhanced cellular binding and internalization of the antibody-functionalized capsules, compared with control human immunoglobulin (IgG) - functionalized capsules, suggesting that these capsules can specifically interact with cells through antibody/antigen recognition (Shimoni et al. 2012).

Folate

The folate receptor has been broadly utilized in targeted drug delivery, especially to target cancer cells, as expression of the folate receptor is selectively upregulated in tumor sites of various organs such as colon, lung, prostate, ovaries, mammary glands, and brain (Morris and Sharma 2011). Folic acid (FA), or folate, specifically binds to the folate receptor, and therefore folate has been conjugated to many polyelectrolytes (Jeong et al. 2005, Kim et al. 2007, Chiu et al. 2011, Nukolova et al. 2011, Zhu et al. 2011, Chiang et al. 2013, Mandal et al. 2011) for active targeted drug delivery to cancer cells. For example, Zhou et al. fabricated PEMs composed of chitosan and alginate assembled by the LbL method on the surface of the BSA- stabilized biocompatible poly(lactide-co-glycolide) nanoparticles. FA-grafted poly(ethylene glycol) (PEG–FA) was covalently bounded to the PEMs via carbodiimide chemistry. The combination of the chitosan/alginate LbL coating with carbodiimide chemistry provides a simple but effective way of tailoring nanoparticle surfaces in a sequential way, to first reduce unspecific interactions and then to attach other molecules to achieve specific targeting properties (Zhou et al. 2010). Hou et al. (2011) used FA as a tumor-targeting agent and designed FA− and methoxypoly(ethylene glycol) (mPEG)-conjugated chitosan nanoparticles for targeted and prolonged drug delivery systems for tumor cell-selective targeting treatments. Biodegradable polymeric micelles, self-assembled from a di-block copolymer of poly (D, L-lactic-co-glycolic acid) and PEG, were prepared to achieve delivery of DOX targeted to the folate receptor (Yoo and Park 2004). Besides cancer, Agarwal et al. developed a formulation to combat diabetes by LbL coating of PAH/PAA over the liposomes to which FA was appended as targeting ligand, by synthesizing an FA− poly(allylamine) hydrochloride conjugate. The insulin entrapped within the freeze-dried formulation was found stable, exhibited stability in simulated biological fluid, and showed prolonged hypoglycemic response (for up to 18 h) in diabetic rats (Agrawal et al. 2013).

Transferrins

The transferrins (Tf) are a family of iron-binding proteins, typically monomeric glycoproteins with a single polypeptide chain of 670–700 amino acids. Tf have been widely applied as ligands for active targeting of anticancer agents, proteins, and genes to primary proliferating malignant cells overexpressing transferrin receptors (Tfr). Antitransferrin receptor–drug conjugates could deliver drugs and genes across the blood–brain barrier for treatment of a broad spectrum of central nervous system diseases (Li and Qian 2002). Earlier, polymeric drug delivery systems targeting Tfr were designed for delivery of the anticancer drug PTX (Shah et al. 2009, Xu et al. 2005), DOX (Dufes et al. 2004, Sivakumar et al. 2009), and for the systemic delivery of oligonucleotides (ODN) to the brain (Vinogradov et al. 2003). PLGA nanoparticles were loaded with PTX, and the surface was epoxy- activated and modified with Tf. When Tf-modified poly(D,L-lactide-co-glycolide) was administered to Sprague-Dawley rats, the distribution of PTX to tumor tissue was highest when compared to that achieved using poly(D,L-lactide-co-glycolide) nanoparticles without Tf (Shah et al. 2009). In another study, Tf was conjugated to the polyelectrolyte-coated gold nanorods for targeted in vitro delivery to cancer cells. The intracellular uptake of the nanorods predominantly occurred in HeLa cells via the Tf− TfR interaction, and the nanorods were localized in vesicular structures such as endosomes (Ding et al. 2007). Fisher et al. showed that Tf, vascular endothelial growth factor (VEGF), and basic fibroblast growth factor (bFGF), can each be linked to poly N-(2-hydroxypropyl) methacrylamide (HPMA)- coated poly (L-lysine)/DNA complexes. All ligand-targeted complexes demonstrated increased uptake into receptor-positive cells, thus permitting versatile targeting of genes to selected cells (Fisher et al. 2000).

The use of pathogenic microorganisms producing an oncolytic agent may be an approach for apoptotic therapy in cancer treatment. Granicka et al. investigated the targeting efficiency of Bacillus subtilis bacterial cells coated with polyelectrolyte shells modified with Tf to increase affinity toward the target tumor cells. It was observed that the bacterial cells coated with modified shells with incorporated Tf exhibited stronger lethal impact on leukemia cells as compared to bacterial cells with no modification of the shell coating (Granicka et al. 2014).

Peptides

Peptides derived from the sequence of cell surface proteins, such as intercellular adhesion molecule-1 (ICAM-1), LHRH, Bombesin, and LFA-1, have shown potent binding affinity to the target cell surface receptors (Majumdar and Siahaan 2012). Peptides chemically bound to polyelectrolytes and adsorbed or embedded in the PEM films were found to retain their biological activities. PEM films made of PLL and PGA have been functionalized by covalent binding of a synthetic analog of the anti-inflammatory peptide α-melanocyte-stimulating hormone (α-MSH) to PGA, to create biologically active coatings for tracheal prostheses (Schultz et al. 2005). An elegant strategy for bi-functionalization of PEM films was developed by Meyer et al. These authors showed that it is possible to functionalize the PEM film, both for cell transfection and for activation via a peptide signaling pathway. They prepared films containing plasmid DNA pre-complexed with poly(ethyleneimine) (PEI) and a peptide molecule NBPMSH. This peptide, grafted to PGA, was used as a signal molecule for melanoma cells B16–F1, and for its ability to enhance gene delivery in a receptor-independent manner (Meyer et al. 2008).

Integrins are among the most widely studied receptors for tumor-targeted drug delivery, since they play a key role in cancer metastasis and tumorigenesis. There are over 25 known integrin receptors, and most of these recognize the small tripeptide turn sequence arginine–glycine–aspartic acid (RGD), which is a part of extracellular proteins such as fibronectin and vitronectin. Toublan et al. used the electrostatic adhesion approach to adhere RGD-containing peptides to the surface of serum albumin microspheres to target tumor cells. These microspheres are selectively taken up by HT29 human colon cancer cells in vitro (Toublan et al. 2006). The RGD peptide has been conjugated with chitosan using a thiolation reaction. RGD enhances selective intratumoral delivery by delivering plexin domain-containing protein 1 (PLXDC1)-targeted siRNA into the alphanubeta3 integrin-positive tumor endothelial cells in A2780 tumor-bearing mice. This approach resulted in significant inhibition of tumor growth, compared with controls (Han et al. 2010).

RGD peptides coupled to PEM films augment osteoblast adhesion and can be used to spatially control cell adhesion. Chua et al. functionalized chitosan–HA polyelectrolyte multilayers and immobilized RGD-containing peptide on PEM substrates, to achieve enhanced osteoblast functions while retaining antibacterial efficacy (Chua et al. 2008). Tsai et al. (2009) immobilized dual peptides containing RGD (a cell-binding domain) and LHRRVKI (a heparin-binding domain) onto polystyrene by LbL, and the effects on osteoblast cell culture were investigated. In another study, the cell proliferation in osteoblasts was enhanced in the presence of PEM films of PLL/PGA₅–PLL/PGA-RGD (Picart et al. 2005). Similarly, Gribova et al. investigated PEM films made of PLL and PGA as substrates of tunable stiffness that can be functionalized by an RGD-adhesive peptide, to investigate important events in myogenesis, including adhesion, migration, proliferation, and differentiation. (Gribova et al. 2013).

Asialoglycoprotein

The asialoglycoprotein receptor is expressed on the surface of liver hepatocytes, and specifically recognizes the terminal b-D-galactose and N-acetylgalactosamine residues (Ruiz and Drickamer 1996). Galactose-bound catonized polymers such as galactosylated chitosan enable direct delivery and internalization of DNA into liver through the asialoglycoprotein receptors to which galactose binds (Gao et al. 2005). D-galactose has been effectively incorporated into the multilayers by LbL assembly of galactosylated polyelectrolyte, by Zhang et al. (2006, 2008a, 2008b). They described the construction of hepatic- targeting microcapsules by self-assembly of chemo-enzymatic synthesized poly(vinyl galactose ester-co-methacryloxyethyl trimethylammonium chloride) (PGEDMC) containing galactose branches, which can be specifically recognized by membrane-bound galactose receptors (ASGPR) for the acyclovir (ACV) controlled release system (Zhang et al. 2008b). Earlier, the cationic D-galactose-branched copolymer PGEDMC alternated with PSS to form thin multifilms by LbL on different solid surfaces such as quartz slides, poly(ethylene terephthalate) (PET) films, silicon wafers, and polystyrene microparticles was reported. PGEDMC/PSS planar films and capsules carrying β-galactose as recognition signals have specific recognition abilities with peanut agglutinin (PNA) lectin rather than concanavalin A (Con A) lectin, observed by fluorescence spectroscopy (Zhang et al. 2008a). Another significant study about multilayers of galactosylated chitosan and plasmid DNA was reported by Cai et al. (2008) in which the films had a specific higher transfection rate on hepatoma G2 cells.

Growth factors

The growth factors and their receptors play an important role in cell growth and proliferation, and are overexpressed in a variety of tumors. Muller et al. demonstrated that vascular endothelial growth factor (VEGF) adsorbed on (PAH/PSS)₄ maintained its bioactivity in vitro and stimulated endothelial cell proliferation (Müller et al. 2008). In another study, the polyelectrolyte films were functionalized with brain-derived neurotrophic factor (BDNF) and semaphorin 3A (Sema3A). The viability of motoneurons on polyelectrolyte multilayers was higher compared to that on polyelectrolyte monolayers (Vodouhê et al. 2005). When basic fibroblast growth factor (bFGF) was immobilized onto quartz slides and collagen films by assembly with chondroitin sulfate (CS), it was found that the immobilized growth factors retained maximal bioactivity (Ma et al. 2007). These kinds of polyelectrolyte architectures may find diverse application in tissue engineering.

Passive targeting

Passive targeting is based on the accumulation of drugs in the areas around the tumors with leaky vasculature, by an enhanced permeation and retention (EPR) effect. Schneider and his colleagues employed the smallest (25–100 nm) LbL nanoparticles [AuNPs coated with five primer layers of PAH and PSS and functionalized with N-(2-hydroxypropyl) methacrylamide (HPMA) copolymers] to enhance the accumulation of active drug in the tumor tissue (Schneider et al. 2009). In the studies by Poon et al. (2011b) carboxylated quantum dots (QDs) were sequentially coated with PLL and DS, then terminally capped with an outer layer of either PLL, DXS, or HA. After systemic intravenous injection in mice, the blood concentrations of the nanoparticles decreased in a two-phase manner (Figure 7). All particles were observed to undergo a rapid distribution phase within 1 h of injection. The longer persistence of HA-terminated nanoparticles in the blood stream corroborates their superior stability and biodistribution profile. They also monitored the accumulation of the HA-terminated nanoparticles in subcutaneously induced KB tumors with intravital fluorescence imaging over a period of 72 h. It was observed that 24 h post injection, the HA-terminated particles were detectable in the tumors. Nevertheless, the nanoparticle accumulation and clearance profile in the tumor is typical of EPR-based targeting, which is short-lived, with a maximum accumulation of 24 h post injection, as there are no active mechanisms in place to promote cell uptake or extend their residence time in the tumor interstitials. Previously, trilayer architecture of PLL modified with iminobiotin, followed by the linker protein neutravidin and biotin end-functionalized PEG that enabled the layered nanoparticles to avoid rapid reticuloendothelial system (RES) clearance, allowing their accumulation in tumor interstitials due to EPR effect, was reported (Poon et al. 2011a).

Figure 7. Blood circulation and tumor targeting of optimized LbL nanoparticles. (A) Blood circulation profiles of QD₇₀₅, QD₇₀₅/PLL/[DXS/PLL]₃/HA and D₇₀₅/PLL/[DXS/PLL]₃/DXS. The longer persistence of QD₇₀₅/PLL/[DXS/PLL]₃/HA in the blood stream corroborates their superior stability and biodistribution profile. (B) Enhanced permeation and retention (EPR) based targeting of solid KB tumors induced subcutaneously on both hind flanks using QD₇₀₅/PLL/[DXS/PLL]₃/HA. Image is taken at the 24 h time point. (C) Time dependant accumulation of QD₇₀₅/PLL/[DXS/PLL]₃/HA in KB tumors. Accumulation of the nanoparticles in tumors is transient and typical of EPR dominated targeting. Data is given in mean ± SEM, n = 6. Reproduced with permission from reference (Poon et al. 2011b) Copyright American Chemical society 2011.

Conclusion and outlook

In this review, the recent progress and the various research studies carried out in the field of endogenous stimuli- responsive polyelectrolyte drug delivery systems has been summarized. As put forth by Park, there are at least four modulated delivery systems that have to be developed in the future third generation drug delivery systems. They are the glucose-sensitive transient insulin delivery with on–off switching capability, targeted delivery of anticancer agents or siRNA to tumors, long-term drug delivery ranging from 6 months to 1 year, and in vitro testing methods that can predict in vivo pharmacokinetic profiles (Park 2014). Of these, the most challenging is a mode of delivery which mimics the pulsatile levels of insulin in blood. At the moment, it is not clear which endogenous stimuli could be utilized for such a delivery system. Inspiration might come from exogenous principles. One potential route, which holds promise in modulating blood insulin levels in vivo, is based on the principle of the light-cell-pump (Sommer et al. 2010), which has been recently exploited to force permeable nanovesicles to release their cargo (Carter et al. 2014).

PEMs are particularly attractive for third generation applications because of their responsiveness to a variety of environmental stimuli, and because their properties can be easily tailored to specific needs. Moreover, PEMs hold great promise for designing theranostic platforms. For instance, recently Zhang et al. (2014) prepared multifunctional magnetic PAH/PSS microcapsules that can be used in drug delivery, hyperthermia, and also simultaneously as MRI contrast agents for remote cancer cell therapy and imaging. However, there are a few bottleneck challenges to the successful implementation of PEMs into the catalog of third generation drug delivery systems. One major challenge would be designing PEM-based delivery systems with stability against physiological barriers. Furthermore, detailed knowledge about the internalization behaviour of PEM within the cells, the biocompatibility, and biodistribution in vivo is only partially available. Recent progress regarding these aspects was achieved by Morton et al. in a model study which employed a poly (lactic-co-glycolic acid) drug-containing core as a substrate for LbL deposition. The authors fabricated a range of coated systems to investigate the potential of these polyelectrolytes to stabilize the drug for delivery, as well as to improve the pharmacokinetics of both the drug and carrier (Morton et al. 2013). Figure 8 shows the stability assessments of different nanoparticle architectures using the IVIS system, a highly versatile in vivo imaging technology, for simultaneous tracking of both drug and nanoparticle. In another study, the multiple internalization pathways of PEMs into the mammalian cells were explored by Kastl et al. (2013).

Figure 8. Biodistribution channels for both drug, cardiogreen (CG₈₂₀) and nanoparticle (PLL₇₀₀) with significant differences in the accumulated signal intensities for different formulations of the layered architectures in the liver, relative to the controls, within the first 30 min post-administration on both channels. Reproduced with permission from reference (Morton et al. 2013) Copyright Elsevier 2013.

Thus, there is ample evidence that biodegradable polyelectrolyte nanocapsules are multifunctional vehicles which can smuggle drugs into cells, and release them upon endogenous activation. Therefore, the interest in the design of intelligent PEM drug delivery systems is clear, and the need of the hour is a systematic translation of PEM-based drug delivery systems from the lab to clinical studies. In addition to applications in drug delivery, polyelectrolytes find increasing applications in other different fields like biosensors (Campos et al. 2014), biomaterial coatings (Panayotov et al. 2014), encapsulation of living cells (Stange et al. 1993, Granicka et al. 2009), bioelectrochemistry (Wiemann et al. 2008), photoelectrochemistry (Yuan et al. 2014) etc. and consequently, any progress in LbL-based PEM could lead to significant advancement in these fields as well.

Declaration of interest

The author reports no declaration of interest. The authors alone are responsible for the content and writing of the paper.