Polymeric siRNA gene delivery – transfection efficiency versus cytotoxicity

Journal Pre-proof PolymericsiRNA gene delivery– transfection efficiency versus cytotoxicity Anna Kargaard, Joost P.G. Sluijter, Bert Klumperman

PII:

S0168-3659(19)30610-8

DOI:

https://doi.org/10.1016/j.jconrel.2019.10.046

Reference:

COREL 9995

To appear in: Received Date:

14 July 2019

Revised Date:

23 October 2019

Accepted Date:

23 October 2019

Please cite this article as: Kargaard A, Sluijter JPG, Klumperman B, PolymericsiRNA gene delivery– transfection efficiency versus cytotoxicity, Journal of Controlled Release (2019), doi: https://doi.org/10.1016/j.jconrel.2019.10.046

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Polymeric siRNA gene delivery – transfection efficiency versus cytotoxicity Anna Kargaarda,b, Joost P.G. Sluijterb,c, Bert Klumpermana a Stellenbosch University, Department of Chemistry and Polymer Science, Private Bag X1, Matieland 7602, South Africa b University Medical Center Utrecht, Experimental Cardiology Laboratory, Department of Cardiology, Division of Heart and Lungs, P.O. Box 85500, 3508 GA, Utrecht, The Netherlands c Utrecht University

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Graphical abstract

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Abstract

Within the field of gene therapy, there is a considerable need for the development of non-viral vectors that are able to compete with the efficiency obtained by viral vectors,

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while maintaining a good toxicity profile and not inducing an immune response within the body. While there have been many reports of possible polymeric delivery systems, few of these systems have been successful in the clinical setting due to toxicity,

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systemic instability or gene regulation inefficiency, predominantly due to poor endosomal escape and cytoplasmic release. The objective of this review is to provide

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an overview of previously published polymeric non-coding RNA and, to a lesser degree, oligo-DNA delivery systems with emphasis on their positive and negative attributes, in order to provide insight in the numerous hurdles that still limit the success of gene therapy.

Abstract ...........................................................................................................................1 Introduction: Gene delivery..........................................................................................2 Cationic polymeric delivery systems ..........................................................................5

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Poly(2-(N,N-dimethylamino)ethyl methacrylate) (pDMAEMA)........................................ 18 Poly(L-lysine)........................................................................................................................ 23 Other polypeptide delivery systems ................................................................................. 30 Polyester delivery systems ................................................................................................ 32 Phosphonium-containing polymers .................................................................................. 38 Amphiphilic polymers ......................................................................................................... 42

Charge reversing polymers ........................................................................................ 43 Polymers directly bonded to RNA ............................................................................. 53 Nanoparticles ............................................................................................................... 56

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Future outlooks ............................................................................................................ 59 Funding Acknowledgement ....................................................................................... 60 Conflicts of Interest ..................................................................................................... 60

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References .................................................................................................................... 60

Introduction: Gene delivery

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Gene therapy encompasses the delivery of genetic material to cells in order to modulate gene expression, and ultimately to treat diseases. There are two approaches

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for gene therapy, where the first is to deliver a functional copy of a gene that is defective/absent, while the second is to deliver RNA interference (RNAi) that can suppress pathological gene expression. RNAi includes the following nucleic acids:

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short interfering RNA (siRNA), microRNA (miRNA), short hairpin RNA (shRNA) and antisense oligonucleotides (AONs). We will focus specifically on polymeric systems for

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the delivery of RNAi, also known as polyplexes, since these delivery systems have received considerable attention in recent years and may fulfill the high expectations of gene therapy.

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RNAi was first discovered by Fire et al. [1], and the importance of RNAi has become apparent over the past 20 years, with siRNA and miRNA becoming accepted as playing a vital role in gene regulation. Their role in gene regulation has been studied extensively and it has become clear that apart from their role in healthy individuals, these nucleic acids have been linked to playing a decisive role in diseases such as cancer, neurological diseases and heart disease. [2] However, the lack of success in gene therapy, despite a large number of gene therapy-related clinical trials, can often

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be attributed to the absence of safe, reliable and efficient tools for targeted delivery of these RNA molecules to their intracellular locations in diseased cells. This lack of success is also due to the RNA molecules’ size, negative charge, instability and problematic cellular uptake. [3] Although there has been some success in local delivery of naked RNA molecules, especially siRNA, [4-6], complications arise when targeting tissue via systemic administration. Numerous studies have investigated diverse means of protecting RNAs via delivery systems, with the intention of transporting the therapeutics to reach their site of action. Davis et al. reported the firstin-human phase 1 clinical trial in 2010, in which they showed systemic administration of siRNA via a

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PEG-coated nanoparticle delivery system for the treatment of solid cancer. [7] Since then, many delivery systems have been developed, including viral vectors, [8-10] liposomes, [11, 12] dendrimers, [13-16] peptides [17-19] and polymers [20-22]. In general, polymeric delivery systems do not cause an immune response like those seen for viral vectors [23, 24] and, due to the development of controlled polymerization

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techniques, the ease with which polymeric delivery systems can be modified make them a strong alternative as vectors within gene therapy.

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There are a number of factors that are important when designing a polymeric gene delivery system, and these factors are based on the end function of the gene therapy

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vehicle. Figure 1 summarizes the journey of a polyplex system from complexation to gene modulation. It is vital that the vector protects the payload from degradation, which can be caused by the constituents within the extracellular fluid such as serum

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components and proteins, while not causing an immune response. Moreover, a delivery system should not form non-specific aggregates or cause aggregation of blood components, as this can cause capillary occlusions. [25] It is also important that

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the vehicle facilitates cellular uptake, [26] usually through endocytosis, [26, 27] where a major obstacle remains the endosomal/lysosomal release of the polymeric delivery

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system into the cytoplasm. [28, 29] Thereafter, one further hurdle is the release of the payload from the polyplex in order for it to be able to perform its intended purpose of regulating gene expression. [30]

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ro of -p re lP na ur Jo Figure 1. The journey of polycationic gene delivery systems, from complexation, to systemic delivery and extravasation, endocytosis and cytoplasmic release to the final steps of the formation of the RNA-induced silencing complex and mRNA inhibition/degradation.

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Thus far, no delivery vehicles have been able to address all requirements listed above, but large strides have been made over the past decade, with some systems undergoing Phase 3 clinical trials [31-38] and one drug, Onpattro (Alnylam) [39] recently having been approved by the US Food and Drug Administration (FDA) for use in the clinic. The knowledge gained from oligo-DNA delivery vectors, both positive and negative, can be applied to non-coding RNA delivery, although some slight refinement might be needed in order to achieve its new purpose. Therefore, knowledge of oligoDNA delivery can be used as a type of shortcut for the design of delivery vehicles for non-coding RNAs. Here, we will discuss a variety of different systems that have

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been designed for both non-coding RNAs and oligo-DNAs, with an emphasis on their positive and negative attributes.

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Cationic polymeric delivery systems

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Figure 2 a) Polyplex formation via electrostatic interactions; b) schematic of endocytotic cellular uptake of polyplex and subsequent release into the cytoplasm. Based on ref [40].

Cationic polymers have emerged as highly attractive vehicles for gene therapy and

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are, by far, the most deliberated polymeric gene delivery vehicles for both non-coding RNAs and oligo-DNAs. Generally, these polymers are composed of amines that can be protonated, and which are thus capable of interacting with the non-coding RNAs and condensing them to form polyplexes, as can be seen in Figure 2. By condensing the nucleotides, the polycations protect them from degradation by decreasing their interactions with proteins and enzymes. Furthermore, polymeric complexation of nucleotides

decreases

the

chances

of

certain

motifs

in

the

non-coding

RNAs/oligoDNAs from being recognized by Toll-like receptors (TLRs) through the

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nucleotides’ interaction with membrane surfaces, which would activate an undesirable immunogenic response. In addition, the polycations reduce systemic clearance of the nucleotides, enable their systemic transportation, increase their half-life time, and facilitate cellular uptake through increased interaction with cells. [41] However, a pivotal disadvantage of polycation systems is the inherent cytotoxicity that they cause due to their positive charge. [42] Table 1 gives an outline of the major advantages and disadvantages of the key polycations that have been used in gene delivery applications. Table 1 Summary of the pivotal advantages and disadvantages of core polycations applied in gene delivery

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Polyethylenimine (PEI)

Pros Cons - Regarded as gold - Very cytotoxic and as standard for in vitro usage such unsuitable for in - Good RNAi complexation vivo gene delivery - efficiency Causes haemolysis - Positive charge allows for enhanced cell permeation - Elusive mechanism of endosomal escape ('proton sponge' effect) - Aggregation and protein interaction Higher RNAi complexation efficiency than PEI due to sterically unhindered and free nature of the amines

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Polymer system

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Cationic polymer systems

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Poly(2(N,Ndimethylamino)ethyl methacrylate) (pDMAEMA)

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Positive charge allows for enhanced cell permeation

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Can be polymerized by RDRP techniques, increasing control over molecular weight and architecture RDRP allows for the incorportation of 'shielding' copolymers other than PEG

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Inherently cytotoxic, like PEI Transfection efficiency of PEI remains significantly higher Causes haemolysis Unknown mechanism of endosomal escape

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Amphiphilic polymers

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Phosphoniumcontaining polymers

- Poor transfection - efficiency - High cytotoxicity - Low buffering capacity Improved toxicity profile - Cationic nature still infers cytotoxicity compared to ammonium Higher levels of in vitro and in vivo stability in solution Amphiphilic nature adds - Decreased cellular liposome-like uptake compared to characteristics, while still polycations easily modified since they are synthetic polymers Less cytotoxic than polycations Still able to facilitate cellular uptake

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- Instability towards protease causes short circulatory half-life time rendering them unsuitable for in vivo gene delivery

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Poly(L-lysine)

- Aggregation and protein interaction

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RDRP allows for incorportation of comonomers to decrease toxicity and improve transfection efficiency, e.g. butyl methacrylate Improved biocompatibility since its monomeric units are amino acids Polymer is biodegradable - Can be copolymerized with other monomers for decreased cytotoxicity and improved transfection efficiency - Primary amines accessible for further functionalization

The nitrogen to phosphorus charge (N/P) ratio is the ratio of positive charge on the

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polymer due to the amine groups to the negative charge on the nucleic acid as a result of the phosphate groups. The N/P ratio is an important physicochemical factor to consider when forming polyplexes as it has an impact on the size, stability and net charge of the delivery system. Higher N/P ratios have been seen to form smaller micelles and enhance gene expression in vivo. [43, 44] This enhanced gene expression is due to the presence of an excess of cations, which increases intracellular delivery. However, when the charge neutral point is surpassed due to these excess

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cations, the cytotoxicity of the polyplexes increases due to non-specific aggregation with constituents within the extracellular fluid. [45] Fischer et al. [46] conducted an in vitro study that explored the structural effects of the polymer that influenced the polycation’s impact on cell viability and on haemolysis. This study highlighted that the cytotoxicity of a polymer directly relates to the cationic charge density and the molecular weight. Thus, significance is placed on finding the correct N/P ratio for each system, i.e. the balance between transfection efficiency and toxicity profile. One way in which to improve the toxicity profile of polyplexes is to add a hydrophilic surface-coating polymer block, e.g. poly(ethylene glycol) (PEG). This also offers other

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advantages such as an increased circulation time in vivo, decreased aggregation, and more protection against blood components. [47-50] On the other hand, due to shielding of the positive charge, PEGylation has been seen to decrease complexation efficiency [51] and cellular interactions, causing a decrease in cellular uptake. This issue has been addressed by actively targeting cells by adding targeting ligands or antibodies

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on the surface of delivery systems. [28, 48, 52-56] Furthermore, cellpenetrating peptides have been attached to the surfaces of polyplexes due to their ability to assist

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in penetrating cell membranes. [57-61] Similarly, nuclear localization signal peptides have been attached in order to direct the gene carriers into the nucleus. [62, 63]

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A review by Zhao et al. [64] gives a comprehensive overview of multi-targeting peptides for gene delivery, including an insight into both cell-penetrating peptides and nuclear localization signal peptides. PEGylation increases passive targeting of polyplexes, i.e.

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it increases the chance of polyplexes to reach their targeted site due to an increased circulation time. [65] Incorporation of targeting ligands and antibodies allows for active targeting through interaction with receptor proteins on specific cell types, which

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subsequently facilitates cellular uptake. [66] For example, Allen et al. [67] synthesized poly(1-vinyl-imidazole)

(PVIM)

quaternized

with

various

t-Boc-protected

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bromoalkylamines (up to 30% quaternization) and then, after Boc-deprotection, was decorated with folic acid, see Figure 3. Common luciferase transfection assays [68, 69] were used in order to test the transfection efficiency of the polyplexes since it is a very convenient way to get comparative results. In short, a luciferase luminescent reporter gene is transfected into mammalian cells, and the transfection efficiency of a polyplex, containing a corresponding gene, is assessed by measuring the luciferase luminescence. [68] Although these polyplexes wholly underachieved in comparison with polyethylenimine (PEI) and SuperFect Transfection Reagent, [70] which is an 8

activated dendrimer molecule used as control due to its high in vitro transfection efficiency and good reproducibility, in terms of transfection efficiency. Nevertheless, the results underscore the remarkable impact that folic acid has on the cellular uptake of polyplexes, whereby the presence of folic acid decoration on the delivery system increased the luciferase expression up to 250 times compared to the unmodified analogues. Thus, this reiterates the importance of targeting ligands for enhanced

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cellular uptake, such as folic acid, within gene therapy.

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Figure 3 Post-polymerization functionalization of alkyl-functional poly(1-vinylimidazole) through conjugation with NHS-folate in order to achieve receptor-mediated nonviral delivery. [67]

The main advantages and disadvantages of PEGylation are shown in Table 2. The transfection efficiency of PEGylated systems decreases due to diminished intracellular release. It has been repeatedly observed that endosomal escape is of greater significance than cellular uptake when attempting to achieve efficient gene knockdown. [71-74] Some evidence shows that, through the use of pH labile,

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PEGylated polymer systems, it is possible to improve the endosomal escape of the delivery systems. [75-79] In the case of polycations, a controversial theory, the proton sponge effect, suggests that the breaking of stimuli responsive linkers between PEG and the polycation, causes an elevation in osmotic pressure due to the number of products. Consequently, this increases the pressure within endosomes, which in turn causes the swelling and rupture of the endosome. Therefore, the use of these pHlabile linkers can facilitate gene regulation through delivery of the vector to the cytoplasm. [74, 80-82] Table 2 Advantages and disadvantages of PEGylation

PEGylation Pros - Decreases toxicity - Decreases aggregation

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Cons - Decreases cellular uptake efficiency - Decreases endosomal escape

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- Improves stability and circulation time - Decreases the overall transfection and decreases non-specific interactions, efficiency, particularly in vivo thus facilitating passive targeting

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Reversible deactivation radical polymerization (RDRP) techniques, such as atom transfer radical polymerization (ATRP) and reversible addition-fragmentation chain

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transfer (RAFT)-mediated polymerizations, allow for the synthesis of polymers with predetermined molecular weights and architectures. These techniques have enabled the elucidation of the effects of polymer molecular weight and architecture on polyplex

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transfection efficiency and toxicity. The results of numerous studies have shown that, by increasing the molecular weight of the polycation, or by increasing the molecular weight of the polycation block within a block copolymer, the oligonucleotide

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condensation capabilities and the transfection efficiency of the polyplex are also increased. However, these increases also cause a rise in the inherent cytotoxicity.

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[8387] This increase in cytotoxicity has been linked to two causes. Firstly, the increased molecular weight causes an increase in the polyplex size, which thus increases its propensity to bind to negatively charged peptides. Secondly, the polyplex interacts more with the cell membrane, causing destabilization and, consequently, cell death. Hence, it is imperative to optimise molecular weight when developing polymers for gene delivery applications.

Similarly, the polymer backbone architecture, e.g. homopolymers, block copolymers, statistical or random copolymers, graft copolymers, and star-shaped polymers, imparts 10

important properties that influence gene delivery. Due to a lower charge density, random copolymers impart lower cytotoxicity than block copolymers. [88] Furthermore, polyplexes produced from grafted and branched copolymers display better toxicity profiles, as well as higher transfection efficacy, than those from linear homopolymer analogues. [83, 89, 90] These grafted copolymers display higher electrostatic interactions with the nucleic acids and cell membranes, subsequently improving membrane disruption and increasing uptake. Ziebarth and Wang [91] used coarse-grained molecular dynamics simulations to investigate the effect of polymer architecture, as well as molecular weight of blocks,

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within copolymers on the polyplex structure. These theoretical models revealed that polyplexes that consist of longer, linear block copolymers were smaller and possess more well-defined core-corona structures compared to copolymers containing lower molecular weight blocks. This is due to the improved ability to condense the RNA/DNA molecules for the higher molecular weight blocks. Linear copolymers were reported to

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be homogeneously dispersed with RNA/DNA molecules whereas polyplexes, prepared from star-shaped polymers, form layered structures. Furthermore,

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experimental investigations have shown that polyplexes from star-shaped polymers display higher transfection efficacy. This is due to the enhanced condensation of RNAs

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by star-shaped polymers compared to their equivalent linear or randomly branched polymers. [92-97]

Georgiou et al. established that, beyond the importance of polymer’s architecture, the

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positioning of monomers within a polymer is of equal significance, as the architecture impacts the transfection ability of the resulting polyplex. [98, 99] Synatschke et al. [95] suggested that ideal polyplexes are those prepared from branched polymers with

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intermediate molecular weight, as this would help researchers to discover the optimal balance between transfection efficiency and cytotoxicity.

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A number of different polycations have been investigated for use in gene delivery systems. A few of these systems will be discussed in more detail in the subsequent sections. The most well researched polycation is PEI, see Figure 4. [100, 101]

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Polyethylenimine (PEI)

Figure 4 Structural representations of linear and branched polyethylenimine (PEI).

PEI has been named the “gold standard” of non-viral based delivery systems and, due

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to this, there have been a tremendous number of studies that vary PEI’s architectures, add modifications or conjugations, PEGylation and target different cell/tissue types. PEI’s success as a gene delivery vehicle can be rationalized by its ability to complex with non-coding RNAs and oligo-DNAs due to its amines that can be protonated to

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bear a strong positive charge. Key advantages and disadvantages of PEI have been described in Table 1. The most important point is its positive charge which also causes

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the inherent cytotoxic nature of PEI. [45, 102] The levels of its cytotoxicity have been strongly correlated to the size and structure of PEI. [103] Thus, the charge ratio of PEI

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polyplexes is very important, with increased N/P ratios resulting in smaller particles with more positive charge, and hence allow for enhanced permeation into cells. [104] Of course, this leads to increased cytotoxicity. In an in vivo study conducted by Chollet

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et al., [105] a clear enhancement in the luciferase activity in the lung was noted with an increase in the concentration of linear PEI/DNA polyplexes. However, an adverse reaction of this increase in concentration was necrosis in the liver and fatality. Fischer

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et al. [46] discovered that the nature of cell death caused by linear PEI and polyplexes was indicative of a nectrotic type of cell death due to the fact that cell apoptosis was

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not detected. Comparatively, a study by Kafil et al. [106] reported that branched PEI is capable of eliciting apoptosis in target cells. This study noted that branched PEI (25 kDa) caused higher levels of toxicity than linear PEI (25 kDa). The increased cytotoxicity was thought to be related to the higher z-potential of branched PEI compared to linear PEI. Interestingly, a subsequent study has contradicted previous studies and relayed evindence that linear PEI demonstrates higher transfection efficiency than branched PEI. [107]

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It has become a well-known phenomenon that, as the molecular weight of PEI increases, at a constant N/P ratio, so too the cytotoxicity rises. However, the transfection efficiency of low molecular weight PEI is also notably lower. Therefore, studies have been conducted in which chains of low molecular weight PEI are connected via a reducible link, in order to obtain biodegradable high molecular weight PEI. [103, 108] This has proven to be effective in obtaining superior RNAi activity with lower cytotoxicity. Gosselin et al. [108] incorporated a disulfide bond between two low molecular weight PEI molecules, using dithiobis(succinimidylpropionate) and dimethyl‚3,3′-dithiobispropionimidate, see Figure 5. These polymeric systems showed

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slightly lower transfection efficiencies compared to high molecular weight PEI; however, the cytotoxicity was also somewhat reduced via the inclusion of the disulfide bridge. [108] In comparison to this, Breunig et al. investigated the viability and efficacy of disulfide crosslinked, low molecular weight, linear PEI in numerous cell lines. They obtained a decrease in cytotoxicity (>90% viability across all cell lines), while improving

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the transfection efficacy. Nonetheless, the maximum efficacy of these reducible PEI polyplexes required higher N/P ratios relative to the non-reducible PEI polyplexes.

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Thus, one can deduce that, although the use of reducible bonds between low molecular weight PEI can decrease cytotoxicity to some extent, the associated

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transfection efficiency may be compromised when compared to non-reducible PEI of

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the same ultimate molecular weight. [109, 110]

Figure 5 Disulfide bonds incorporated between two low molecular weight PEI molecules via dithiobis(succinimidylpropionate) (DSP) and dimethyl‚3,3′-dithiobispropionimidate bonds (DTBP). [108]

As previously mentioned, there can be no true RNA delivery without endosomal escape, and hence no gene modulation can occur, as the delivered RNA molecule

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must be able to enter the cytosol in order to interact with the targeted RNA. Understanding the mechanism of endosomal escape is pertinent in order to improve actual delivery systems. However, this mechanism is still rather elusive in the case of PEI. A number of theories abound regarding endosomal escape. The most accepted, and most widely debated, theory is the “proton sponge” effect, vide supra. [111] This theory is based on the large buffering capacity of polycations, including PEI. It is believed that the endosome will eventually rupture due to an increase in osmotic pressure, caused by the uptake of chloride ions that are passively transported into the endosome due to the charge gradient. Although popular, this theory has also been

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greatly disputed. A study by Benjaminsen et al. [112] reported evidence that suggested that no decrease in pH occurs within the lysosome, thus shedding doubt on the accuracy of the proton sponge effect. Yet, no alternative explanation currently exists to explain how PEI polyplexes causes gene regulation, only that they do. In elaboration of the abovementioned, PEGylation has become common practice in order to address

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problems with toxicity, stability and circulation time of PEI polyplexes. [113-115] Due to the disadvantages of PEGylation, see Table 2, studies have been conducted in

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order to investigate the effect of PEG chain length, in the hope of discovering the optimum architecture for the simultaneous protection and maintenance of cellular

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uptake and transfection efficiency. [113, 114, 116] However, due to proof that hydrophilic polymers on the external surface of a delivery device decrease cellular uptake and thus transfection efficiency, [117, 118] many PEGylated PEI polyplexes

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have included targeting ligands or antibodies on the surface, in order to promote active cell targeting via enhanced uptake. [119-121] As such, folic acid has become a popular targeting ligand in gene therapy due to its selectivity towards cancerous cells that

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overexpress folate receptors, thus increasing its internalization efficiency. [120, 122] Acid-labile linkers have also been incorporated in PEGylated PEI polyplexes in order

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to improve the endosomal escape, and consequently the transfection efficacy. [75, 118, 123] Knorr et al. [75] synthesized novel PEG-acetal-PEI polyplexes through a maleimide moiety on the PEG chain-end and mercaptan-functional PEI, see Figure 6. The resultant PEG-acetal-PEI was seen to break down after 3 minutes at pH 5.5 – the pH corresponding to that within the endosomal compartment. A 10-fold increase in transfection efficiency was obtained, which was notable in comparison to PEGylated PEI with a pH stable linkage.

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Figure 6 Synthesis of novel, pH sensitive PEG-acetal-PEI polymers via Michael addition between a maleimide moiety on the PEG chain-end and a mercaptan-functional PEI for application in gene delivery. [75]

Other examples have incorporated disulfide links between the hydrophilic polymer and

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PEI, for example Carlisle et al. [124] connected PEI and poly(hydroxypropyl methacrylate) (pHPMA) via a reducible disulfide bond. They hypothesized that the increase obtained in terms of transfection efficiency was caused by the reducing

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environment in the late-endosome/lysosome, which thus caused the breakage of the disulfide linker. However, in more recent studies, the use of disulfide links has been

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rendered less effective than pH-labile bonds. This will be discussed further in the section entitled Nanoparticles.

Despite all these investigations, the transfection efficiency of PEI still remains

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inadequate, especially when the toxicity profile is kept in check. Thus, although research on PEI polyplexes remains active, there has been a flurry of new research on gene delivery, as many no longer consider PEI to be the ultimate vector for these

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delivery systems. Instead, “smart” polymeric systems have become leaders in this field. Their “smart” nature can be attributed to a range of different factors, including

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biodegradability, responsiveness to stimuli such as pH, temperature, reducing environment, etc. However, all of these systems have one thing in common – they aim to enhance biodistribution, biocompatibility, cellular uptake, endosomal escape and cytoplasmic release – and ultimately their main function is the delivery of the RNA species in order to control gene expression. Accumulatively, this is an attempt to improve what PEI could not achieve. Few of these “smart” polymeric systems, if any, are able to fulfil all the listed objectives. Many of them have utilized RDRP techniques,

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e.g. RAFT-mediated polymerization and ATRP, to obtain more sophisticated polymers and copolymers with better controlled molecular architecture. A major advantage of RDRP techniques is that they enable the synthesis of polymers in a sequential manner and allow for facile means of incorporating functional modifications, such as targeting ligands, fluorescent markers or other imaging modalities, or even the RNA molecule itself. A more recent example by Li et al. [125] used PEI in order to design a mixed micelle composed of comb-like and grafting copolymers, synthesized in part by ATRP and ring-opening polymerizations. Their comb-like copolymer consisted of poly(lactide-

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co3(S)-methyl-morpholine-2,5-dione)-poly(poly(ethylene glycol) monomethacrylate) (PLMD-PPEGMA) linked to an endothelial cell-specific targeting ligand, Cys-Arg-GluAsp-Val-Trp (CREDVW). They co-assembled these polymers with poly(lactide-co3(S)methyl-morpholine-2,5-dione)-g-polyethylenimine (PLMD-g-PEI), an amphiphilic graft copolymer. Compared to PEI/pDNA, this mixed micelle system demonstrated

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increased gene transfection efficiency, measured by monitoring the ZNF580 protein expression via western blot analysis. The CREDVW-modified micelles doubled the

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cellular internalization efficiency compared to scrambled CREVDW-modified micelles. Furthermore, wound healing assays demonstrated the cell migration ability of the

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delivery system by visualizing the migration process as a function of time on an inverted microscope. However, the study failed to show evidence of significant endosomal escape in the fluorescence micrographs, where a large portion of the

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fluorescent signal of the DNA therapeutic overlapped with the lysotracker signal.

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Figure 7 Chemical structure of comb-like PLMD-PPEGMA-CREDVW copolymers and grafting amphiphilic copolymer, PLMD-g-PEI, which were assembled into mixed micelles for gene delivery. [125]

Wang et al. [126] produced hyaluronic acid-PEI carbon dots without the use of any additives (acid, base or oxidant) by microwave-assisted pyrolysis. Hyaluronic acid is the carbon source for the carbon dots, while PEI acts as the surface passivation agent. In this case, hyaluronic acid acts in a similar shielding manner to PEG due to its biocompatibility and negative charge. Moreover, hyaluronic acid has high affinity towards cell-specific surface markers, e.g. cluster determinant 44 (CD44) receptors

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that are overexpressed on a range of tumour cells. Six different carbon dots were produced by varying the microwave irradiation time from 5 to 50 minutes. By DLS measurements, it was noted that an increase in irradiation time led to a decrease in hydrodynamic size, attributed to an increased consumption of PEI and hyaluronic acid

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residues as the reaction time progressed. FTIR spectroscopy, NMR spectroscopy and elemental analysis indicated that all six carbon dots were very similar in composition

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and structure. However, increased irradiation time did seem to have some negative effect on the resulting photoluminescence, postulated to be caused by an increase in hyaluronic acid dehydration leading to a larger carbon core with more PEI passivated

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onto the surface. A gel retardation assay revealed that all carbon dots were capable of complexing one pDNA per carbon dot (1:1 ratio, w/w), significantly lower than most

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polyplex systems. The complexed carbon dots at different carbon dot/pDNA (w/w) ratios had a particle size ranging from 240 to 360 nm. With an increase in the w/w ratio, the z-potential also increased from –5 mV to +44 mV. The study revealed that at

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a w/w ratio greater than 6, the z-potential was higher than +20 mV, which is advantageous for efficient cellular uptake. Cytotoxicity was evaluated by transfection

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of the carbon dots at different w/w ratios in HeLa cells for 24 hours. Although the cytotoxicity increased slightly as the w/w ratio increased, at a w/w of 14, 80% viability was still achieved. In comparison, a control experiment of PEI/pDNA (w/w 14) only showed 20% viability of cells. A luciferase assay in HeLa cells was used to establish the transfection efficiency of various w/w ratios of the carbon dots, and compared to 25 kDa PEI (N/P 10). It was established that a w/w ratio of 6 is required for adequate transfection efficiency, and the efficiency increased with an increase in the w/w ratio. Furthermore, at these w/w ratios, the carbon dots managed to show a 3-4 times greater

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transfection efficiency than PEI. Wang et al.

also determined that a microwave

irradiation time of 10 min produced carbon dots with the highest transfection efficiency. Serum tests revealed that hyaluronic acid improved the serum tolerance of carbon dots 50 times compared to unmodified PEI polyplexes. Cellular uptake studies with and without the presence of serum was undertaken using Cy5-labelled pDNA. Transfected cells were analyzed by flow cytometry 4 days after transfection. Up to 10% serum, no decrease in cellular uptake was observed, with approximately 99.7% cells positive for Cy5-labelled pDNA. At a serum content of 50%, 77.8% of cells were still positive for Cy5-labelled pDNA. These results are remarkably higher than those of PEI,

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where the positive cells decreased to 55.5% in the presence of 10% serum. Cellular imaging of transfected cells support that carbon dots are effectively being transported into cells. It would have been interesting to see whether the carbon dots managed to escape the endosome/lysosome, however this was not reported. That said, these hyaluronic acid-PEI carbon dots represent an exciting pathway towards a potential

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gene delivery vector.

Poly(2-(N,N-dimethylamino)ethyl methacrylate) (pDMAEMA)

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The anionic nature of RNA remains a simple means by which to form polyplexes. Poly(2-(N,N-dimethylamino)ethyl methacrylate (pDMAEMA), Figure 8, is a wellstudied polymer for these sorts of delivery systems [127]. The key advantages and

na

disadvantages of pDMAEMA applied as a gene delivery vector can be found in Table 1. Like PEI, its complexation is also based on the protonation of amines, thus causing endosomal escape to be based on the hypothesis of the debated proton sponge effect.

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Owing to the sterically unhindered and free nature of the amines of pDMAEMA, they have a higher capacity of absorbing protons, and thus possess a higher proton sponge

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buffering potential compared to PEI. [110, 128] In a comparative study, pDMAEMA with tertiary amines was compared to those with quaternized amines. It was found that, although the complexation with nucleic acids is higher for quaternized pDMAEMA, the toxicity also increases, while the transfection efficiency decreases. The authors hypothesized that this decrease in transfection efficiency is caused by the inability of quaternized amines to be further protonated, which then decreases the buffering capacity of the polymer. [83, 129]

18

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Figure 8 Chemical structure of pDMAEMA

Via RDRP techniques, many investigations have been conducted on the effect of molecular weight on complexation efficiency, toxicity, cellular uptake and transfection efficiency of pDMAEMA. Like PEI, pDMAEMA is inherently cytotoxic due to its cationic

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nature, and this toxicity increases with an increase in molecular weight. [111, 130] It has been observed that transfection efficiency rises with an increase in molecular

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weight, and that the only limiting factor is toxicity. In addition, the effect of polymer architecture on the polyplex properties has been studied and, as such, linear,

lP

starshaped, [86, 94, 95] graft copolymers, [83, 89, 130] statistical copolymers [79, 131133] and block copolymers of DMAEMA [84, 134] have all been investigated. Studies of pDMAEMA grafted onto poly(2-hydroxyethyl methacrylate), [89]

na

hydroxypropyl cellulose, [83] dextran [129], and chitosan [90] have highlighted that these grafted or branched polymers have better toxicity profiles compared to their high molecular weight, linear pDMAEMA counterparts. Moreover, their transfection

ur

efficiency is also higher, even at lower cation incorporation. [83] This same phenomenon was also seen for hyperbranched pDMAEMA of the same molecular

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weight as linear pDMAEMA. [135] Synatschke et al. [95] investigated the effect of branched pDMAEMA on transfection efficiency and cytotoxicity by synthesizing linear, 3-arm and 5-arm pDMAEMA of different molecular weights. They found that a minimum threshold molecular weight is required in order to obtain cellular uptake, the value of which differs according to the polymer architecture. Moreover, in line with other studies, they reported that linear pDMAEMA is more cytotoxic than star-shaped polymers. However, they also specified that they noticed an upper limit of the dependency of N/P ratio on cellular uptake. This is more likely due to a stronger

19

complexation of plasmid DNA to the polycation at higher N/P ratios, which as described by Vader et al. [136], causes a shielding effect of the green fluorescent protein encoded in the plasmid. This in turn causes an experimental inaccuracy of the decrease in fluorescent signal within the cell obtained via flow cytometry. Vader et al. [136] described an experimental method to negate this shielding effect by chemically disrupting the polyplex and simultaneously causing cell lysis, followed by measuring the fluorescent signal in the lysis supernatant. In this study, they showed that the cellular uptake of polyplexes increases proportionally with the N/P ratio. PEG has been used extensively as a protective block for pDMAEMA in a similar

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manner to its use with PEI. However, as controlled polymerization techniques lend themselves to easy incorporation of comonomers, there have also been investigations into the use of other hydrophilic protective blocks. For example, Song et al. [134] synthesized graft copolymers of poly(N-vinylpyrrolidone) (PVP) and pDMAEMA (PVPg-pDMAEMA), as well as further chain extensions with methyl methacrylate

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(PVP-g(pDMAEMA-b-PMMA)) via ATRP, see Figure 9. PVP-g-pDMAEMA is able to form coils in aqueous medium, while PVP-g-(pDMAEMA-b-PMMA) forms micelles with

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PMMA in the hydrophobic core. The incorporation of PMMA allowed for better condensation of pDNA, a higher buffering capacity and subsequently a higher

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transfection efficiency. As such, both of these systems display an increased

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ur

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transfection efficiency when compared to PEI (25k) at low N/P ratios.

Figure 9 Schematic of the synthetic route via ATRP to obtain PVP-g-(pDMAEMAbPMMA). [134]

In order to tackle toxicity, and inspired by reducible PEI, You et al. [137] synthesized reducible high molecular weight pDMAEMA by coupling low molecular weight pDMAEMA via disulfide bonds. The use of RAFT-mediated polymerization makes synthesis of well-defined reducible pDMAEMA easier than that of reducible PEI. You

20

et al. synthesized a bisfunctional RAFT agent and, post-polymerization, removed the trithiocarbonate-end groups in order to obtain α,ω-dithiol-pDMAEMA. Thereafter, this telechelic pDMAEMA was oxidized to form the high molecular weight product. They obtained lower cytotoxicity than the non-reducible analogue, and equivalent transfection efficiencies. However, transfection using PEI as the vector remains significantly higher than that using pDMAEMA. In order to further reduce toxicity, studies have worked towards developing smarter pDMAEMA polymers that possess reduced charge, by incorporation of neutral comonomers. Methyl methacrylate (MMA), [134, 138] butyl methacrylate (BMA) [131,

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139, 140] and butyl acrylate (nBA) [80] have all been copolymerized with DMAEMA in order to decrease the cytotoxicity and improve the transfection efficiency. The copolymers exhibit a good ability to complex the nucleotides, nonetheless, slightly higher N/P ratios are commonly necessary to achieve this complexation compared to homopolymers of DMAEMA. While incorporation of nBA improves stability and lowers

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cytotoxicity of the polyplexes, it reduces their pH-dependent disruptive behavior. Thus, this leads to a decrease in RNAi activity, which may be due to endosomal entrapment.

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[131] Conversely, copolymerization of DMAEMA and BMA produces pH tunable polymer micelles, with a pH directly related to the fraction of BMA incorporated in the

lP

polymer chain. In a study conducted by Nelson et al., various percentages of BMA were incorporated into PEGylated copolymers of DMAEMA and BMA. It was found that polymers with a BMA content of ³ 50 mol% form micelles at pH 7.4. The pH at

na

which the micelles disassociate decreases with increasing fraction of BMA. At 70% BMA, the micelles are stable – even at pH 4.0. This suggests that they do not gather

ur

enough positive charge to destabilize the increased hydrophobic interactions brought about by the higher BMA content. [131] The investigation revealed that the cellular uptake of the polyplexes increases as the percentage of BMA decreases. However,

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polyplexes based on 50% BMA copolymers exhibit the highest luciferase silencing relative to all other copolymers. This indicates that the polyplexes, based on the 50% BMA, show better endosomal release compared to other BMA compositions, as a result of their pH dependent profile. Stayton and coworkers [133] incorporated propylacrylic acid (PAA) into this system, forming poly(DMAEMA-b-[DMAEMA-coPAAco-BMA]) via RAFT-mediated polymerization, Figure 10. This system shows improved serum stability and, most importantly, the endosomal escape of the polyplex is

21

enhanced due to the tuning of the pH-dependent membrane disruptive behavior of the polyplex. They later incorporated a folate targeting ligand through a folatecontaining RAFT agent. This system was reported to further facilitate specific cellular uptake

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through folate-receptor interaction. [132]

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Figure 10 Chemical structure of poly(DMAEMA-b-[DMAEMA-s-PAA-s-BMA]) designed by Stayton and coworkers. [133]

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To improve endosomal escape, pH-labile linkers have been incorporated between the hydrophilic surface polymer and pDMAEMA. Lin et al. [141] incorporated a pH labile ortho-ester as linker between pDMAEMA and PEG. Unfortunately, they found that the

na

ortho ester linker is not the ideal bond type for a gene therapy system as it is stable around pH 6.5. Thus, this means that the deshielding effect of PEG was incomplete. Trützschler et al. [142] conducted a study to compare the gene transfection efficiency

ur

of methacrylate polymers bearing primary, secondary and tertiary amine pendant groups, namely polymers based on (2-aminoethyl)-methacrylate (AEMA), N-methyl(2-

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aminoethyl)-methacrylate (MAEMA), and DMAEMA. The lowest transfection efficiency was observed for polymers containing more DMAEMA, while the polymers containing higher mol% of primary amino groups showed the highest transfection efficiency. The study revealed that with transfection efficiency in mind, the amino functionality was more important than polyplex size, buffering capacity and rate of cellular uptake of the polyplexes. The buffering capacity of the polymers had little to no impact on the endosomal release, which supports the idea that methacrylate polymers escape the

22

endosome by the disruptive formation of pores rather than via the “proton sponge” effect. Poly(L-lysine) Poly(L-lysine) was one of the first cationic polymers to be used for the complexation of nucleotides due to its biodegradable nature. [143, 144] Nevertheless, it was judged to be a second rate choice compared to other investigated polymers, such as PEI or pDMAEMA, due to its poor transfection efficiency, caused by a lack of buffering capability, short circulatory half-life time and high cytotoxicity. [145] In fact, it has been

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revealed to possess an equivalent toxicity to branched PEI. [46] The key advantages and disadvantages of poly(L-lysine) have been summarized with the other important polycations in Table 1.

It has been possible to incorporate other polymer blocks with poly(L-lysine) in order to improve transfection efficiency and/or decrease toxicity. [146, 147] One such example

-p

was investigated by Patil et al., [148] who synthesized triblock poly((amido amine)b(ethylene glycol)-b-(L-lysine)) nanocarriers, as seen in Figure 11. Poly(L-lysine)

re

provides the polycationic nature required for the condensation of RNAs, as well as encourages cellular uptake. Conversely, the poly(amido amine) enhances endosomal

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escape and cytoplasmic delivery of RNAs due to the presence of its tertiary amine groups, i.e. it functions as proton sponge. Gene knockdown is substantially decreased when poly(amido amine) is not present in the nanocarrier, thus confirming its essential

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ur

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role in making the delivery system effective.

23

Figure 11 Nanocarrier designed by Patil et al. composed of PAMAM-PEG-PLL. [148]

Another example in which a system was built around poly( L-lysine) was the use of PEG-b-poly(L-lysine)-b-(L-leucine) for gene delivery. [149] It was observed that, only by tuning the poly(L-lysine) and poly(L-leucine) ratio, was it possible to obtain better transfection efficiency than with poly(L-lysine) homopolymer. A similar study incorporated poly(aspartamide) derivatives into PEGylated poly( L-lysine) in a welldefined A-B-C triblock copolymer. [147] The poly(aspartamide) was introduced with the intention of enhancing in vivo micelle stability, which was previously seen to be lacking in polyplexes that spontaneously assemble due to electrostatic interactions.

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The presence of both hydrophilic nucleotides and hydrophobic polycations is believed to cause some destabilization in vivo. [147, 150, 151] The abovementioned stabilized polyanion micelles by Kim et al. [147] were compared to equivalent random block copolymers as hydrophobic control micelles, as well as non-PEGylated diblock

-p

copolymer micelles. They reported that the non–PEGylated micelles are more readily taken up by cells, while the “randomly” hydrophobic control micelles are taken up the

re

least. However, the cellular uptake did not tally with efficient gene silencing, as the performance of the non-PEGylated micelles were the least successful. This was

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determined to be due to the fact that these micelles are unable to release the siRNA in the cells, whereas the triblock micelles release their payload more effectively. The triblock micelles also outperformed the “randomly” hydrophobic control micelles, which

na

was correlated to their balance between serum stability and RNA release. [147] Interestingly, a more recent investigation by Zhang et al. [152] discovered a relationship between the length of the hydrophobic block and the gene transfection

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efficiency. They synthesized poly(L-lysine)50-block-poly(L-leucine)n, with n 10, 15 and 25. The larger the hydrophobic block, the higher the transfection efficiency. The

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poly(Lleucine) block was capable of negating the cytotoxicity usually associated with poly(Llysine) above DP 30, while the block copolymers still showed transfection efficiencies comparable to that of PEI.

Kataoka and coworkers have extensively explored PEG-b-PLL modified with thiol and thiolane functionalities for use as polyion complex (PIC) micelles. Their polymer functionality started by containing thiol groups which crosslink and stabilize the core through reducible disulfide links, see example in Figure 12. [153-156] The disulfide crosslinking stabilizes the structure of the micelles in physiological conditions, while 24

also allowing for cytoplasmic release of RNAi due to the reductive environment within the cell. These responsive systems are able to achieve a 100-fold increase in siRNA activity compared to the unmodified PEG-b-poly(L-lysine). Although the blood circulation of the PIC micelles was improved, the RNAi activity was low. This was improved by decorating the surface of the PIC micelles with cRGD peptides, resulting in better cellular uptake, better subcellular distribution, blood stability and targeted tumour accumulation. In 2011, the group published that the large majority of 2iminothilane modifications undergo intramolecular ring-formation which results in Nsubstituted 2-iminothiolanes due to the instability of the amidines. [158] The result is

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a decrease in the polymer charge density at physiological conditions, thus requiring a higher N/P ratio for optimal PIC micelle formation. However, the results also indicate that the modified polymer has higher bloodstream and buffer stability when it contains a fraction of N-substituted 2-iminothiolanes. This underscores the importance of nonionic and noncovalent interactions in the stability of PIC micelles, as well as the

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importance of a proper balance between micellar stability and RNAi activity. The group has further functionalized these polymers with different targeting ligands/antibodies,

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ur

na

lP

re

including cRGD, [159] anti-TF Fab’ [160, 161]

Figure 12 Chemical structure of stimuli responsive core shell-type polyion complex (PIC) micelles containing PEG-b-poly(L-lysine) with iminothilane modification which enables a cross-linked core via disulfide bonds. [157]

25

Oishi et al. [162] conjugated siRNA and lactocylated PEG via a β-thiopropionate acidlabile linkage, and subsequently complexed it with poly( L-lysine), in order to introduce pH-sensitivity into their novel PIC micelles, as seen in Figure 13. Since the βthiopropionate bond is stable at physiological pH, but readily cleaves at the pH of the intracellular endosomal compartment (pH = 5.5), the smart complexes are able to release siRNA from the lactocylated PEG. This allowed them to achieve enhanced gene silencing in hepatoma cells at extremely low siRNA concentrations compared to Lac-PEI-siRNA complexes. In fact, they obtained an almost 100 times enhancement in RNAi activity for their pH sensitive PIC micelles when compared to the Lac-

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PEGsiRNA conjugate. Very notably, these PIC conjugates retained their RNAi activity even when incubated with 50% serum for 30 minutes prior to transfection. Oishi et al. [162] investigated the importance of the pH-labile linker to RNAi activity by inhibiting endosomal acidification. This was achieved by adding nigericin to the culture medium. As a result, they saw a significant decrease in the RNAi activity, while no change was

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seen in the RNAi activity for the Lac-PEG-siRNA conjugates. This suggests that the cleavage of the acid-labile linker indeed occurs within the intracellular endosomal

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compartment, and that it is probable that the free PEG increases the osmotic pressure within the endosomal compartment, which causes swelling and disruption of the

na

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compartment. Ultimately, this allows for release of the siRNA into the cytoplasm.

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Figure 13 siRNA conjugated to lactocylated PEG via a β-thiopropionate acid-labile linkage, which was subsequently complexed with poly(L-lysine), in order to introduce pH-sensitivity in the polyion complex micelles. [162]

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An interesting trend was observed when comparing pHPMA-oligolysine polymers synthesized via RAFT-mediated polymerization and via conventional free radical polymerization, Figure 14. [163] The transfection efficiency was seen to be similar in both systems, however, the IC50 values (the polymer concentration at which there is 50% cell survival) for the conventional free radical polymers were 10-fold lower than those of the RAFT-mediated polymers, indicating that the conventional free radical polymers cause far greater cytotoxicity.

26

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et

al.

[164]

developed

star-shaped

poly(L-lysine)

polymers

via

na

Byrne

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Figure 14 A schematic illustration of the synthesis of HPMA-oligolysine polymers containing a reducible or non-reducible linker between the polymer backbone and the RNA-complexing, oligolysine side-chains.

Ncarboxyanhydride ring-opening polymerization, which is known to produce polypeptides of varying architecture with good control over molecular weight. Figure

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15 shows the chemical structure of the individual arms of the stars. They investigated the gene delivery potential (both pDNA and siRNA) of a range of well-defined

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starshaped L-lysine polymers with varying number of peptide arms, including linear, 8– , 16–, 32– and 64–arm polymers. They reported that star-shaped poly(L-lysine) was much more efficient at condensing pDNA and siRNA than linear poly(L-lysine). In fact, linear poly(L-lysine) was ineffective at complexing siRNA even at N/P ratios up to 300, while the star-shaped analogues efficiently complexed siRNA at N/P ratios below 5. They presented evidence that the complexation efficiency of their poly( L-lysine) systems was not only dependent on the number of lysine residues present, but also on the architecture of the system, whereby the size of the dendritic core and density

27

of the polypeptide arms played an important role. They noted that when the density of poly(L-lysine)-arms was too high, the interaction between siRNA and cation was limited and the siRNA was located at the periphery of the polymer structure. In comparison, when the polymer architecture contained fewer, higher molecular weight arms, the siRNA was more capable of infiltrating the polymer structure, thus being exposed to more L-lysine residues. The gene modulation efficiency of the 64–arm, low molecular weight poly(L-lysine) was compared to that of the linear analogue at a range of N/P ratios. The star-shaped polyplex was reported to enhance luciferase expression significantly more than the linear polyplex at all N/P ratios tested. However, at N/P

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ratios above 50, the star analogues were reported to display significant cytotoxicity.

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Figure 15 A schematic of the chemical structure of one arm of the 8, 16, 32 or 64 star poly(L-lysine) gene delivery vectors. [164]

Recently, Walsh et al. [97] described the use of structurally identical 16–, 32– and 64– arm star poly(L-lysine) for delivery of pDNA to mesenchymal stem cells (MSCs), see Figure 15. The cellular uptake of the 64–arm analogue was greater at a lower N/P ratio compared to the fewer armed stars. The 64–arm star polymer also displayed the greatest transfection efficiency at N/P 5 compared to the stars with fewer arms. The

28

64–arm star with a DP = 5 per arm was identified as the best transfection vector in its class, with a 1000-fold increase in transfection efficiency compared to its linear analogue. The cellular uptake of these stars was identified to proceed via a rapid, clatherin-independent mechanism. In their latest work, Walsh et al. [165] furthered this research through the functionalization of collagen with these pDNA-star(32/64) poly(Llysine) whereby they demonstrated prolonged transgene expression in MSCs. Again, the 64-arm analogue performed better than both the 32–arm and linear poly(Llysine) equivalents. Furthermore, they reported that the 64–arm analogue showed double the transgene expression compared to PEI. Their in vivo subcutaneous implant

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model results not only showed nanomedicine retention within the implanted scaffold, but also high transfection efficiency into the host cells after 7 days.

Ren et al. [166] modified graphene oxide with poly(L-lysine) and RGDS tetrapeptide targeting ligand for the application of VEGF-siRNA delivery to tumour cells. On the surface of the sheets, graphene oxide contains hydroxyl and epoxide functional groups

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while the edges contain carbonyl and carboxyl functional groups. These functionalities not only ensure biocompatibility but also allow the graphene oxide to be modified, as

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in the example provided by Ren et al. They established that a graphene oxide to poly(Llysine) (w/w) ratio of 5 produced highly dispersed conjugates as a stable

lP

suspension. Agarose gel retardation studies revealed that at an N/P ratio of 10, the VEGF-siRNA was efficiently complexed to the modified graphene oxide. Cellular uptake was confirmed by laser scanning confocal microscopy. Cytotoxicity studies

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revealed that the presence of the RGDS moiety significantly decreased the toxicity of the system, and viability remained above 85% at modified graphene oxide concentrations of 150 µg·mL-1. Gene transfection efficiency was measured by

ur

evaluating the effect on the VEGF mRNA post transfection of HeLa cells. The results from the qPCR revealed a 41% down-regulation of VEGF mRNA after transfection with

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the poly(L-lysine) and RGDS modified graphene oxide complexes (N/P 10). Furthermore, the VEGF protein expression was reduced by 52% compared to the blank control. In vivo studies were performed whereby the inhibition of tumour growth was investigated in mice. Mice injected with poly(L-lysine) and RGDS modified graphene oxide/VEGF-siRNA complexes showed around 50% tumour inhibition while mice injected with naked VEGF-siRNA only showed less than 5% tumour inhibition. The results of these modified graphene oxide complexes compared well to mice injected with DOX (circa 56% tumour inhibition). Furthermore, biodistribution studies 29

tracking the presence of the RGDS related peak [M + K]+ was performed on the heart, liver, kidney, spleen, brain and tumour tissue. The RGDS related peak was only detected in the tumour tissue, indicating no accumulation of the modified graphene complexes in any other organs, thus the RGDS-modified complexes were successfully delivered to the tumour tissue. Further in vivo studies are necessary, however, these modified graphene oxide conjugates represent an exciting new conjugate delivery system which explores previously infererior systems, e.g. poly(L-lysine) with new modifications that enhance biocompatibility, biodistribution and transfection efficiency.

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Other polypeptide delivery systems Shuai and coworkers [167] developed a diblock copolymer of PEG and diethylenetriamine grafted poly(L-aspartic acid) (PEG-PAsp(DETA)) as a delivery vehicle for a miRNA inhibitor of miR-21, as seen in Figure 16. It was tested both in

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vitro and in vivo as a biodegradable delivery system. They found that the copolymers were capable of complexing the miRNA-inhibitor at a minimum N/P ratio of 4, but the optimum N/P ratio was 8, based on their monodisperse particle size (80 nm), spherical

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morphology and optimal z-potential. CCK-8 cytotoxicity assays were performed on both free polymer as well as polyplexes. It was noted that lower cytotoxicity was

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observed for the polyplexes compared to the free polymer, which was attributed to the positive charge of the free polymers. At an N/P ratio below 15, cell viability after 48

na

hours was approximately 80%, while at high N/P ratios, cell viability decreased to 55% (N/P 80). Cell uptake studies revealed that the polyplexes were taken up by PC-3 prostate cancer cells and confocal laser-scanning microscopy confirmed that the

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polyplexes were distributed within the cytoplasm and synapses of the prostate cancer cells. miRNA-21 expression in PC-3 cells was evaluated by RT-PCR and the PDCD4

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protein expression was determined via western blot. Gene and protein expression results compared well with those of Lipofectamine 2000. The in vivo anti-tumour efficacy was evaluated with the PC-3 xenograft model on BALB/c nude mice. After 30 days, the average tumour volume of the group treated with mPEG-

PAsp(DETA)/miRNA inhibitor was significantly smaller than that of the negative control group and PBS control group. RT-PCR results of the harvested organs revealed decreased miRNA-21 expression while PDCD4 expression was significantly increased for the mPEG-PAsp(DETA)/miRNA inhibitor group compared to the control groups.

30

Although the toxicity profile of this system is not optimal, the gene expression results

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reveal an exciting potential new alternative to PEI systems.

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Figure 16 Schematic of the formation of mPEG-PAsp(DETA) from mPEG-PBLA-NH2 as described by Shuai and coworkers. [167]

Chen et al. [168] grafted from PEI (25 kDa) via ring-opening polymerization of L-serine

lP

N-carboxyanhydride to produce a delivery system for the combination treatment with siRNA and pDNA. They then compared the PEI-L-serine to PEI (25 kDa) as delivery vectors. Agarose gel electrophoresis showed that PEI-L-serine/siRNA and PEI(w/w) ratios required for complexation were 0.4 and 0.5 for pDNA and

na

Lserine/pDNA

siRNA respectively, while PEI required 0.3 for complexation to both genes. This slight difference was credited to the higher cationic charge density of PEI, which was

ur

corroborated by measuring z-potentials. Cytotoxicity tests using MTT asssays revealed that the incorporation of

L-serine

increased cell viability compared to

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unmodified PEI, and moreover, cell viability increased proportionally to the degree of substitution with L-serine. However, as the substitution increased, the charge density decreased along with the transfection efficiency, thus calling for a higher w/w ratio for efficient complexation and transfection. Confocal laser scanning microscopy and flow cytometry were used to evaluate the transfection efficacy of labelled siRNA (FAMsiRNA) and pDNA (Cy5-DNA), cotransfected into HeLa cells using PEI-L-serine and PEI. The results confirmed that FAM-siRNA and Cy5-DNA were effectively cointernalized, with a higher fluorescence signal being detected for cells transfected

31

with the L-serine modified delivery system. A cell apoptosis assay for tumour cells was used to evaluate the gene expression efficiency of the combination treatment using PEI-Lserine. The highest apoptotic rate was obtained with the combination treatment using L-serine modified PEI (50.9%), while PEI (25 kDa) obtained 44.3% with the combination treatment. An interesting result in this study revealed that the use of combination therapy increased the apoptotic rate from 15.9% and 44.6% apopotosis with a single treatment of L-serine modified PEI, complexed individually with pDNA and siRNA, respectively. These results show that combination therapy is a possible new avenue to be explored for improvement of gene therapy.

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Polyester delivery systems

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Figure 17 Chemical structure of poly(lactic acid-co-glycolic acid) (PLGA)

Poly(lactic acid-co-glycolic acid) (PLGA), see Figure 17 is a biodegradable polyester

lP

synthesized through copolymerization of lactic acid and glycolic acid. It has found many applications within the medical field due to its biocompatibility and FDA approval as well as to the useful characteristic that its biodegradation time (depending on

na

molecular weight and copolymer ratio) can be tailored. PLGA is able to enhance transfection efficiency because it facilitates the continuous intracellular release of its payload. Within drug delivery, this continuous intracellular release has the ability to

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vary the period of drug release from mere days to weeks. [169] Results indicated that the endosomal escape of PLGA nanoparticles is not due to the proton sponge

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mechanism, but rather due to a selective charge reversal, from anionic to cationic, on the surface of the nanoparticles caused by a change in the pH. The nanoparticle surface is negatively charged in the physiological and alkaline pH while it becomes positively charged when it enters the acidic endosome/lysosome. The positive surface charge cause the nanoparticles to adhere to the endo-lysosomal membranes. [170] The nanoparticles were reported to not adhere to membranes of early endosomes, thus underscoring the need for a decrease in pH in order to facilitate endo-lysosomal

32

escape. That said, the drawback of using PLGA for gene therapy is that it cannot effectively condense nucleotides. Arora et al. [171] attempted to overcome this disadvantage by incorporating PEI into a PLGA delivery system. However, the zpotential of the nanoparticles is not as high as expected, even with the incorporation of PEI. Although a slow, sustained release of the miRNA in vitro was reported, a very significant increase in the time of release in the presence of serum was also noted. This is most likely the result of inefficient complexation of miRNAs. Since then, similar studies have been performed with poly(L-lysine). [172] Nevertheless, the hydrolytic degradability and biocompatibility of polyesters make

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them highly attractive polymers in biological applications – and certainly in gene delivery. [173-175] Although polyesters are structurally very diverse, there are synthetic challenges inherent in generating new monomers, such as linear step-growth polymerization necessitating monomers containing less than 1% impurity. Regardless

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of this obstacle, Nelson et al. [176] recently published their investigation into water soluble polyesters functionalized with imidazolium groups as possible vehicles for gene delivery. Via catalyst and solvent free melt polycondensation reactions, they

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obtained a collection of cationic polyesters of imidazolium diol, neopentylglycol and adipic acid with tunable hydrophobicity and thermal transitions, see Figure 18. The

lP

synthesized polyesters included copolymers with 10-75 mol% ionic content, imidazolium homopolyester and poly(neopentylene adipate) (neutral), which allowed them to investigate the role of charge content and structure on transfection efficiency.

na

The polyplexes achieve DNA complexation at very low N/P ratios (N/P of 4 completely complexes DNA) and show insignificant toxicity in HeLa cells. However, the

ur

transfection efficiency of all the polyester systems is significantly lower than than that of PEI and Superfect controls. Although these results are seemingly disappointing, very few studies have investigated these types of polycations, compared to the

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tremendous number of studies that focus on polycations based on nitrogen ions, e.g. PEI. As these polyesters are able to show transfection at low N/P ratios, by adjusting the cationic charge and structure, it is plausible that there is still some hope that these polyesters may have potential applications in gene delivery.

33

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Figure 18 Catalyst and solvent free melt polycondensation reaction to obtain water soluble polyesters functionalized with imidazolium groups as possible vehicles for gene delivery. [176]

Kozielski et al. [177] optimized the production of a bio-reducible, linear poly(β-amino ester) by incorporating disulfide bonds in the polymeric backbone. These polymers are

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capable of condensing siRNA into nanoparticles, which subsequently release their payload into the cytoplasm of the cells within minutes. This is achieved due to the

re

presence of glutathione in the cytosol. In comparison to their non-reducible counterparts, they are able to obtain efficient green fluorescent protein (GFP)

lP

knockdown with decreased cytotoxicity in human brain cancer cells. GFP expression assays can be used in a similar way to luciferase assays, whereby the gene knockdown is evaluated by GFP levels measured within the cell.

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Unfortunately, the abovementioned bio-reducible nanoparticles display some instability. Thus, in their later work, Kozielski et al. [178] endeavored to address these issues by balancing bioreducibility of the polymer backbone with hydrophobicity from

ur

the poly(β-amino ester). A balance was found between cytotoxicity (lowered by a higher level of bioreducibility and increased with a rise in hydrophobicity), and

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enhanced delivery and stability of the nanoparticles associated with hydrophobicity. This was achieved while still maintaining cytoplasmic targeting imparted by the bioreducible nature of the nanoparticles. They were able to obtain nanoparticles with enhanced properties and gene knockdown. The nanoparticles also display cell specific knockdown, whereby human primary glioblastomas show higher gene knockdown (97 ± 4%) than human primary non-cancer brain cells (27 ± 9%).

Dual-responsive nanoparticles were developed by Wang et al. [179] and used for combination therapy of pDNA and doxorubicin, see Figure 19. The polymer system 34

contained disulfides incorporated into poly(β-amino esters) (ssPBAE) modified with diethylenetriamine and grafted onto oxidized pullulan (oxPL), a natural ligand for the asialoglycoprotein receptor called ASGPR, which is overexpressed by hepatoma cells. The polymer system (ssPBAE-oxPL) was designed to respond to stimuli from both the tumour extracellular pH and the redox state. Doxorubicin was conjugated via an acidcleavable hydrazone bond while pDNA was condensed via ionic interactions with the amino-groups supplied by the diethylenetriamine at N/P ratios above 4. The disulfide linkers were added to decrease the cytotoxicity of PBAE cationic polymers,

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verified by CCK8 assays.

ur

Figure 19 The polymer system (ssPBAE-oxPL) designed by Wang et al. to be both pHresponsive and reducible. [179]

Cellular uptake studies showed nanoparticles were taken up more rapidly by hepatoma

Jo

(HepG2) cells than lung cancer (A549) cells. This was attributed to the ASGPR overexpression in HepG2 cells causing the nanoparticles to be taken up via receptormediated endocytosis by these cells. While for A549 cells, the nanoparticles were thought to be taken up by adsorption-mediated endocytosis, which explains the reduced rate of cellular uptake in A549 cells. These results indicated the possibility that the nanoparticles could partially target hepatomas.

35

They noted that in vitro effects of the nanoparticles on HepG2 cells were significantly higher for combination treatment compared to delivery of pDNA alone, where an 80% cell survival was seen after 48 hours at 500 µg·mL-1 dose. The combination treatment nanoparticles had an IC50 value of 2.1 µg·mL-1 after 48 hours of incubation. Similar particles, but lacking the pH responsive bond, induced a far higher survival rate of the HepG2 cells, with 75% survival after incubation for 48 hours at 24 µg·mL-1 dose. In vivo biodistribution studies for the disulfide-modified nanoparticles revealed that the large majority of nanoparticles ended up in the tumour of the mice bearing a HepG2 tumour after 24 hours, although some of the nanoparticles were detected in the kidney

ro of

and liver. This study reiterates the potential that combination treatment has towards higher therapeutic effects of delivery systems.

Lellouche et al. [180] designed maghemite-incorporated, PEI-capped poly(D,Llactideco-glycolide)-poly(ethylene glycol) (PLGA-b-PEG) nanoparticles, see Figure 20.

-p

The maghemite allowed for the potential of the nanoparticles to be tracked by Magnetic Resonance Imaging (MRI). The influences of molecular weight of the PLGA-b-PEG

re

and the w/w ratio of PLGA-b-PEG to PEI were investigated. It was noted that the nanoparticle size decreased from 70 nm to 50 nm with incorporation of PEI on the

lP

surface compared to naked PLGA-b-PEG nanoparticles. Furthermore, z-potential increased from -50 mV to +40 mV with addition of PEI at pH 6.5-7. Molecular weight of the PLGA-b-PEG seemed insignificant to the nanoparticle size and only a very slight

na

change in z-potential was observed by variation in molecular weight. At high w/w ratios (15:1), highly polydisperse nanoparticles were obtained, attributed to inhomogeneous PEI coating. A 2:1 ratio was selected for further analysis. With the incorporation of

ur

maghemite in the core of the nanoparticles, the nanoparticle diameter increased to 150 nm while the z-potential decreased to -25 mV. SiRNA complexation with the

Jo

maghemite-incorporated, PEI-capped PLGA-b-PEG nanoparticles was observed at N/P 2 via agarose gel electrophoresis. Gene silencing abilities of the complexed nanoparticles were studied using a dual-luciferase reporter assay (Promega), which is based on Firefly [181, 182] and Renilla, [183, 184] after transfection in human U2OS cancer cells. A range of N/P ratios from 1-10 was used for the study. An N/P ratio of 6 resulted in a Firefly knockdown of 94 ±1% with no significant decrease in toxicity observed in the Renilla level. While increasing the N/P to 10 enhanced the Firefly

36

silencing, it also induced some toxicity as seen in the Renilla level (86±3%). The study found that silencing was dependent on the amount of PEI present, and with a decrease in PEI concentration (i.e. at N/P ratios below 4), the nanoparticles were unable to facilitate endosomal escape into the cytoplasm. Further cytotoxicity studies investigating mitochondrial activity via MTT assays confirmed that an N/P ratio of 6 caused no significant changes in cell proliferation or viability, while higher N/P ratios caused some toxicity, which was in accordance with the Renilla assay. Lellouche et al. further investigated the silencing efficiency of SLK-1 siRNA complexed nanoparticles (N/P 6) by transfecting SK-OV-3 human ovarian cancer cells and

ro of

measuring the mRNA levels. A significant knockdown (77 ± 5%) of PLK-1 was observed. It would have been interesting to see how these nanoparticles compare in vitro to unmodified PEI, however, these comparative analyses were not reported in this study. Preliminary in vivo tests revealed that at 1 mg·kg-1 siRNA dose, no significant toxicity or changes to the liver, kidney or the tested hematological

Jo

ur

na

lP

re

concentrations, were observed in the mice.

-p

parameters, including red and white blood cell counts as well as haemoglobin

Figure 20 Design of capped PLGA-b-PEG nanoparticles MRI-trackable due to the maghemite core. Reprinted with permission from Lellouche et al. (2017) RSC Adv. 7, 26912-26920. Published by The Royal Society of Chemistry.

An in vivo study was conducted by Lin et al. [185] investigating the use of biodegradable charged, polyester-based nanoparticles to deliver mutated KRastargeting siRNA in a pancreatic xenograft mouse model. Nanoparticles complexed 37

with fluorescently-labelled siRNA (TRAMA-siRNA) (N/P 8) were injected into mice with pancreatic tumours, peritumourally. The fluorescence signal from TRAMA-siRNA was monitored from 2 to 72 hours post-injection. The fluorescence signal could be observed in the tumour site immediately after injection until 72 hours later, albeit with a marked decrease in intensity. Gene expression was evaluated by real-time PCR, whereby the level of K-Ras mRNA was measured from tumour tissue homogenates. A decrease in mRNA level was observed when the mice were injected with nanoparticles complexed with K-Ras siRNA, while the control containing a scrambled siRNA showed no silencing effect. Western blot analysis showed that K-Ras protein

ro of

expression correlated with the above results. Thus, suggesting that these nanoparticles were capable of functioning as effective gene therapeutic vectors. Repeat peritumoural injections over 20 days revealed that the nanoparticles were able to significantly inhibit tumour growth. By day 7, the tumour volume for mice injected with nanoparticle complexed with K-Ras siRNA compared to those of the control was

-p

visibly reduced (P < 0.01). This pattern of growth inhibition continued throughout the 20 days of the study. Histomorphology of the tumours via Hematoxylin and Eosin

re

(H&E) staining revealed that tumour tissue infiltrated the muscle tissue of the control and nanoparticle-scrambled siRNA groups while no infiltration was observed for the K-

lP

Ras nanoplex group. In vivo toxicity studies revealed that no significant toxicity was caused by the use of the nanoparticles as a therapeutic at 10 times the normal dose. The monitored physical behavior and the weight of the mice revealed no abnormalities.

na

Further, routine blood tests showed no significant toxicity in the kidney or liver. Histological analysis of major organs after the mice were sacrificed showed no indication of toxicity at nanoparticle doses as high as 60 mg/kg. H&E of major organ

ur

sections also confirmed the absence of inflammatory reactions and necrosis.

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Phosphonium-containing polymers There

has

been

some

interest

in

polyplexes

that

are

prepared

from

phosphoniumcontaining polymers. [186-188] The advantages and disadvantages of these systems have been summarized in Table 1. Clément and coworkers were the first to investigate lipids containing phosphonium and arsenium, in comparison to ammonium, for use in gene therapy, with respect to the effect that the cationic core plays on structure and properties/function. [189, 190] They discovered that by replacing the ammonium core with phosphonium or arsenium, they were able to 38

achieve an improved toxicity profile, as well as higher levels of stability in solution – both in vitro and in vivo. Hemp et al. [188]used conventional free radical polymerization in order to synthesize styrene-based ammonium and phosphonium containing polymers. They varied the length of the alkyl substituent attached to the cation in order to elucidate the impact of structure

on

transfection

efficiency.

They

synthesized

poly(triethyl-

(4vinylbenzyl)ammonium chloride (PTEA), poly(tributyl-(4-vinylbenzyl)ammonium chloride (PTBA), poly(triethyl-(4-vinylbenzyl)phosphonium chloride (PTEP) and

na

lP

re

-p

ro of

poly(tributyl-(4-vinylbenzyl)phosphonium chloride (PTBP), Figure 21.

ur

Figure 21 Scheme of conventional free radical polymerization of PTEA, PTBA, PTEP and PTBP. [188]

The transfection efficiency of the four polymers was compared to those of Superfect and PEI, and it was seen that PTBP displays better transfection efficiency than

Jo

Superfect. Moreover, they observed that the transfection efficiency of PTBA is equivalent to that of Superfect. In comparison, the ethyl moieties in PTEA and PTEB seem to affect the transfection efficiency of the polyplexes, as both of these polymers performed inferior to Superfect. As cellular uptake of the ethyl containing polyplexes was seen via fluorescent markers on those polyplexes, it was hypothesized that the poor transfection efficiencies are due to intracellular mechanisms. The consideration was that perhaps the longer alkyl chain in the butyl derivatives aid in membrane

39

disruption and endosomal escape, as has been reported in the past. [133, 191] Further hypothesizing led them to consider that the shorter alkyl chain may have caused stronger electrostatic interactions with DNA, thus decreasing the release of the DNA and inhibiting the end function, as was reported by Song et al. [192] This study stresses that it is not only the type of cation used, but also the immediate structural environment around the cation that impacts transfection efficiency of a polyplex delivery system. However, these polyplex systems are completely ineffective in serum. Therefore, Hemp et al. [187] subsequently developed a phosphonium-containing diblock copolymer based on their previous work, containing a 4-

of

either

oligo(ethylene

glycol)9

methyl

ether

ro of

vinylbenzyltributylphosphonium chloride (TBP) block and a colloidal protecting block methacrylate

(OEGMA)

or

2(methacryloxy)ethyl phosphorylcholine (MPC), all of which demonstrate steric shielding of nanoparticles that led to a resistance towards proteins and prolonged circulation times. [193, 194] Using RAFT-mediated polymerization, as seen in Figure

-p

22, they synthesized their diblock copolymer from a macroCTA containing the stabilizing polymer, which then allowed them to target pTBP with three different chain

re

lengths (DP 25, 50, 75). This enabled them to determine the optimal AB block ratio for gene transfection for this system. In order to establish the effect of the protecting block,

lP

they also synthesized a homopolymer of TBP via RAFT-mediated polymerization, with a molecular weight equal to the intermediate diblock copolymers. As expected, all MPC87TBPy and OEG52TBPy polymers show an increased resistance towards serum

na

proteins, as well as an increased stability at physiological salt concentrations. The diblock copolymers exhibit cell specific uptake in HepaRG cells, with poor cellular uptake, and therefore low transfection, in COS-7 and HeLa cells. The transfection

ur

efficiency in HepaRG cells is equivalent to Jet-PEI, however, the phoshonium diblock

Jo

copolymers cause lower cytotoxicity than Jet-PEI polyplexes.

40

ro of -p

lP

re

Figure 22 Scheme of chain extension RAFT polymerization of TBT on RAFT polymerized macroinitiators of pOEG and pMPC to obtain phosphonium-containing block copolymers for application in DNA delivery. [187]

A study by Ornelas-Megiatto et al. [186] was published at around the same time as the abovementioned research by Hemp et al. [187], and it showed similar results. In this

glycol

na

study Ornelas-Megiatto et al. synthesized poly(acrylic acid) modified with triethylene monochlorohydrin

via

a

hydrolyzable

ester

linkage.

A

second

postpolymerization modification via a nucleophilic substitution of the chloride produced

tris(3-

ur

five different polycations containing triethylphosphinium, tri(tert-butyl)phosphinium,

triphenylphosphonium

and

triethylammonium,

Jo

hydroxy-propyl)phosphonium,

respectively. After eliminating the triphenylphosphonium–containing polymer due to low aqueous solubility, the tri(tert-butyl)phosphonium–containing polymer because of high cytotoxicity and the tris(3-hydroxy-propyl)phosphonium–containing polymer for poor transfection efficiency (even though it was less cytotoxic than the triethylammonium–containing triethylphosphonium–containing

polymer), polymer

it

was

exhibits

ascertained

lower

that

cytotoxicity than

the the

triethylammonium–containing polymer. The phosphonium-based polymers also

41

display a higher transfection efficiency. A notable result from this study was the superior transfection efficiency for the triethylphosphonium–containing polymer in the presence of serum. Thus, this indicates that the polymer is more stable under physiological conditions than its ammonium analog. Amphiphilic polymers A few studies have investigated the use of amphiphilic polymers as vectors for gene delivery. [195, 196] The advantages and disadvantages of amphiphilic polymers can be seen in Table 1. Due to their amphiphilic nature, these polymers incorporate some

ro of

liposome-like characteristics, however, as they are still synthetic polymers, they are also easily modified. Their dual-hydrophilic-hydrophobic nature suggests that they

na

lP

re

-p

would be less cytotoxic than polycations, while still facilitating cellular uptake. [197]

ur

Figure 23 A schematic representation of the pH-response hydrolysis causing a transition from hydrophobic to hydrophilic form, inducing a breakdown of the selfassembled structure. [74]

Jo

Recently, Du et al. [74] developed pH responsive, amphiphilic micelles from poly(ethylene

glycol)-b-poly[(2,4,6-trimethoxybenzylidene-1,1,1-tris(hydroxymethyl)]

ethane methacrylate-b-poly(dimethylamino glycidyl methacrylate), Figure 23. These are

the

first

2,4,6-trimethoxybenzylidene-1,1,1-tris(hydroxymethyl)

ethane

methacrylate (TTMA)-based siRNA delivery systems to be developed. Hydrolysis of the acetal group in acidic pH causes a hydrophobic to hydrophilic transition of the pTTMA, Figure 23Since the hydrophobicity induces self-assembly, this transition causes a breakdown of these self-assembled structures. Furthermore, there is 42

evidence showing that this transition promotes endosomal escape. [198, 199] Consequently, this improves gene knockdown efficiency of the delivery system (at N/P ratios ≥ 20) compared to Lipofectamine 2000, a common lipid-based, in vitro transfection reagent. Importantly, Du et al. [74] reported no significant cytotoxicity caused by the micelles. By selectively inhibiting the mechanisms of clathrin- and caveolae-mediated endocytosis, as well as micropinocytosis, they showed that micelles enter the cell via both clathrin and caveolae-mediated endocytic pathways. These results indicated that the micelles exhibit improved half-life times (30 min) in the body compared to naked siRNA (<5 min). They also investigated the biodistribution of

ro of

the amphiphilic micelles in mice and saw mainly accumulation in the liver, spleen and lungs.

Charge reversing polymers

-

Good RNAi complexation efficiency Enhanced cell permeation Reasonable in vitro gene delivery efficiency

na

-

Cons

re

Pros

lP

Polymer system

-p

Table 3 A comparison of the advantages and disadvantages of cationic polymers versus various charge-reversing polymers

-

-

-

ur

Cationic polymers

Jo

-

43

Very cytotoxic and as such unsuitable for in vivo gene delivery - Causes haemolysis Blood aggregation, protein interaction and immune response - Poor biodistribution and rapid renal clearance Self-assembly of polyplex relient on electrostatic interaction causing structural instability Poor cytoplasmic release of RNAi, causing low gene transfection efficiency Fall short at effective systemic gene delivery PEGylation reduces cellular uptake

-

Blood aggregation, protein interactions and immune response Causes haemolysis

- Poor biodistribution and high renal clearance Dependent on the pH sensititivity of the breakable covalent bond, e.g. amide bond The pH sensitivity occurs at a pH range, therefore complete control over cleavage is not possible

- Core polyplex is still cationic, meaning that the mechanism of endosomal release remains unclear

re

-p

Covalent chargereversal polymers

Decreased cytotoxicity compared to cationic counterparts Improved stability in physiologically relevent environment leading to extended circulation time, less aggregation and decreased renal clearance Improved cellular uptake compared to PEGylated cationic polyplexes Improved endosomal escape facilitated by charge reversal

-

ro of

Decationizing polymers

- RNAi complexation efficiency equivalent to cationic polymers Decationization allows for release of RNAi after cellular internalization - Decreased cytotoxicity compared to cationic counterparts

lP

The charge reversal occurs more rapidly than covalent pHresponsive bonds, e.g. amide bond,

na

Improved cellular uptake compared to PEGylated cationic polyplexes

Jo

ur

Non-covalent, sheddable chargeDecreased haemolytic reversal polymers effect compared to PEI, although not nil effect

The core polymer still consists of the original polycationic systems Core polyplex is still cationic, meaning that the mechanism of endosomal release remains unclear

Improved stability in physiologically relevent environment leading to extended circulation time, less aggregation and decreased renal clearance

The problems associated with the use of polycations are not limited to a systematic issue, i.e. the issues associated with systemic cytotoxicity, blood aggregation and haemolysis, as well as biodistribution and renal clearance, but also expand to the

44

cellular level. Although, PEGylation addresses the systemic toxicity, the cellular toxicity remains a major limiting factor. This is due to disturbances it presents to the cell membrane, [46, 102, 200] including interferences with physiological polyanions, such as enzymes, cell receptors, non-targeted RNA or DNA, etc., [201] and its possible implication in cancer formation through the activation of oncogenes, which also induces apoptosis. [202] Therefore, research has been redirected towards neutral or charge-reversal RNAi delivery systems. A brief summary of the pros and cons of these systems, in comparison to polycationic systems, can be seen in Table 3. One such system, developed by Hennink and co-workers, [203] employs a

ro of

decationized polyplex system for the delivery of siRNA, Figure 24. RAFT-mediated polymerization was employed to obtain poly(2-hydroxypropyl methacrylamideN,N′dimethylaminoethanol-b-N-[2-(2-pyridyldithio)]ethyl methacrylamide)-b-PEG, with folic acid targeting functionality from a (folic acid-PEG5000)2-(4,4′-azobis(4-cyanovaleric acid)) macroinitiator. Complexation with siRNA was attained by employing electrostatic

-p

interactions between the cationic polymer with the negatively charged nucleic acid, in the same manner as polycationic systems mentioned above. However, post-

re

complexation, crosslinking of the polyplexes was obtained via disulfide bonds causing physical entrapment of the RNAs. The polycations were then decationized via

lP

cleavage of the carbonate linking group through hydrolysis, releasing the positively charged side chains from the poly(2-hydroxypropyl methacrylamide) backbone. After cellular internalization of the decationized, reducibly cross-linked polyplexes, the

na

disulfide bonds were reduced in the cytosol, due to the presence of glutathione, allowing for the payload to be deposited. These polyplexes impart lower cytotoxicity

Jo

ur

compared to their cationic counterparts.

45

ro of -p re

na

lP

Figure 24 Schematic illustration of the complexation of poly(2-hydroxypropyl methacrylamide-N,N′-dimethylaminoethanol-b-N-[2-(2-pyridyldithio)]ethyl methacrylamide)-b-PEG via electrostatic interaction between RNA and the 2hydroxypropyl methacrylamide-N,N′-dimethylaminoethanol moiety, and crosslinking via disulfide formation. Subsequent de-complexation is also shown caused by hydrolysis-induced de-cationization. [203]

Monteiro and coworkers [204] developed poly(2-(N,N-dimethylamino)ethyl acrylate) (pDMAEA) which strongly resembles pDMAEMA, however, unlike pDMAEMA,

ur

pDMAEA is able to undergo self-catalyzed hydrolysis in aqueous medium into poly(acrylic acid) and 2-(N,N-dimethylamino)ethanol. This hydrolysis occurs over 10

Jo

hours, but is independent of pH and other external cues. These de-cationizing polymers behave similarly to pDMAEMA, with the marked difference of being able to release their payload within the cytoplasm due to their degradative behavior. [133, 204] Werfel et al. [205] compared PEGylated copolymers of DMAEA and BMA to their DMAEMA analogue and noted an enhanced delivery of the degrading species to the cytosol. It is important to note that hydrophilic protection, e.g. PEGylation, remains an important component for polyplexes regardless of the charged nature of the polymer backbone. [206, 207] 46

In 2004, Prata et al. [208] synthesized charge reversing, functional amphiphiles which transition from cationic to anionic. The concept relies on amphiphiles that are able to bind and release DNA, while also destabilizing bilayers. They designed the amphiphile with three distinct parts: a cationic ammonium headgroup, capable of binding to DNA; lipophilic acyl chains as the hydrophobic section which forms a bilayer; and lastly, terminal benzyl esters which undergo enzymatic hydrolysis upon entering the endosome. In order to assess the necessity of each of the three structural units, 4 polymers

were

synthesized

(a-d),

see

Figure

25,

while

1,2-dioleoxyloxy-

3(trimethylammonio)-propane (DOTAP) (e) was used as a control. DNA complexation

ro of

was observed for 1,3 and DOTAP, while the anionic unit on 2 and the lack of hydrophobic chain on 4 cause unfavourable binding for these two species. Compound

Jo

ur

na

lP

re

-p

1 and DOTAP have similar binding constants with DNA and form 1:1 complexes.

Figure 25 The chemical structures of the five amphiphiles (a-e) tested as potential polymeric gene delivery systems by Prata et al. [208]

47

After incubation of the three DNA complexes in esterase at pH 7.4, the complex formed with 1 dissociated, releasing the DNA therapeutic, while the other two complexes remained intact. The dissociation of DNA with 1 was attributed to hydrolysis of the benzyl ester groups which disturbs the electrophilic complexation. Transfection experiments with β-galactosidase on Chinese hamster ovarian cells showed that the gene complexed with 1 was the most effective of the delivery systems described herein. Further preliminary in vivo studies showed that 1 also facilitates the transportation of gene therapeutics into human embryonic kidney cells, as well as erythroleukemic cells. This research initiated a diversion away from usual cationic

ro of

polymeric gene delivery systems towards the possibility of charge-reversal polymers. It is known that calcium phosphate (CaP) precipitate is able to encapsulate gene therapeutics, and through its incorporation in nanoparticles, it facilitates endosomal escape due to a breakdown of the endosome in a calcium ion-dependent manner. [209] CaP precipitate on its own delivers very low transfection efficiency, suffers from

-p

variation in complex size, and is difficult to implement in vivo. [210] However, it remains an encouraging avenue for gene delivery when incorportated into non-viral vectors.

re

Using their previously designed and effective charge-reversal polymers, [211] Kataoka and coworkers have explored the combination of CaP precipitate combined with their copolymers

of

PEG

and

poly(N″-(N′-((N-cis-aconityl)-2-aminoethyl)-

lP

block

2aminoethyl)aspartamide) (PAsp(DET-ACO)). [212, 213] The cis-aconitic amide linkage of PAsp(DET-ACO) is relatively stable at physiological pH (7.4) but degrades

na

at lysosomal pH (5.5), which converts the polymer from anionic to cationic, see Figure 26. These polymers do not have the inherent cytotoxicity which is associated with cationic polymers, e.g. PEI, however, due to their pH sensitivity and subsequent

ur

charge-reversal, they are still capable of disrupting the endosomal/lysosomal membrane. Included in the nanoparticles designed by Kataoke et al. [212, 213] were

Jo

CaP precipitates, which further facilitate endosomal escape of siRNA.

The initial study focused on the in vivo of CaP/PAsp(DET-Aco)/siRNA nanoparticles for the treatment of subcutaneous pancreatic tumour models in mice, by targeting the pro-angiogenic molecule vascular endothelial growth factor (VEGF). [212] Systemic administration of CaP/PAsp(DET-ACO)/siVEGF nanoparticles resulted in significant inhibition of tumour growth. By day 9, the tumour volumes in mice treated with siVEGF nanoparticles were 66% of the average volume of those in control mice. In a subsequent study, a spontaneous tumour mouse model was used in order to better 48

mimic the tumoural microenvironment in the clinic. These results confirmed the accumulation of nanoparticles in the tumours and reiterated their gene silencing

na

lP

re

-p

ro of

efficiency.

Figure 26 Schematic of the charge-reversal of PAsp(DET-Aco) whereby the amide

ur

bond breaks when the pH decreases [212]

Jo

Kataoka and coworkers further developed these charge-reversal polymers to incorporate a 2-propionic-3-methylmaleic amide as the anionic protective group, as seen in Figure 27. [214] This amide is more sensitive to the decrease in pH, and therefore more responsive to its position in the gene delivery pathway. The second generation PEG-PAsp(DET-PMM) were shown to be capable of forming micelles equivalent in size and stability to PEG-PAsp(DET-ACO). Stability studies of both first and second generation nanoparticles were performed in medium containing 10% fetal bovine serum at pH 7.4 and showed that the nanoparticles were stable for 24 hours. 49

Cellular internalization studies confirmed that there was no difference in their cellular internalization profiles. Most importantly, the second generation charge-reversal polymer significiantly improved the gene silencing efficiency compared to the previous system. These results underscore the importance of the rapid charge-reversal of the polymeric system at pH 5.5, causing enhanced disruption of the endosomal

re

-p

ro of

membrane.

lP

Figure 27 First and second generation, PEG-PAsp(DET-ACO) and PEG-PAsp(DET-PMM) respectively, charge-reversal polymers for gene delivery, developed by Kataoke and coworkers. [214]

na

Based on the premise that charge-reversal polymers are fully dependent on the reactivity of the amide bond, Chen and coworkers have developed charge reversal zwitterionic copolymers. [215, 216] The charge-reversal is caused by electrostatic

ur

shielding interactions between a zwitterionic copolymer of PEI, poly( L-lysine) and poly(L-glutamic acid) (PELG) and the polyplex, consisting of PEI and the DNA

Jo

therapeutic. At physiological pH, the zwitterionic PELG is negatively charged while the polyplex is positively charged. However, in the acidic environment of a tumour, PELG is positively charged, thus it no longer interacts with the polyplex, freeing the polyplex. This enables the unshielded polyplex to interact with negatively charged tumour cells, leading to improved cellular uptake. [215] DLS results revealed that the PELGshielded PEI polyplexes were 200 nm at physiological pH, while non-unimodal, larger sizes were observed when the pH was decreased to 6.8, attributed to aggregation. These sizes are larger than optimal for gene delivery application.

50

The study showed a 10-times improvement in the transfection efficiency of the PELGshielded PEI polyplexes at pH 6.8 compared to at pH 7.4. Haemolytic experiments revealed that PELG-shielded PEI polyplexes caused significantly lower haemolysis than unshielded PEI polyplexes. It would have been interesting if a comparison of the PELG-shielded PEI polyplexes had been performed against PEGylated PEI, since unPEGylated PEI is notorious for causing haemolysis. Their in vivo mouse study showed that the PELG-shielded PEI polyplexes were able to decrease tumour growth rates better than the controls. [215] Thus, pH-resposive shielding aids in improving gene delivery efficacy of cationic polyplexes.

ro of

In a further study by Chen and coworkers, [216] a similar shielding polymer was prepared whereby the copolymer zwitterion replaced PEI with oligoethylenimine (OEI), thus obtaining copolymers of OEI, poly(L-aspartate) and poly(L-lysine) (OEAL). The results observed in this study were equivalent to those described in their earlier work, vide supra. [215] For both these systems, systemic biodistribution and specific

-p

targeting is the next obstacle that needs to be addressed.

Yang et al. [217] developed a pH-responsive, anionic (mPEG45-b-PAEP75-CyaDMMA,

re

PPC-DA), Figure 28, block copolymer which was used as a shielding polymer for branched ssPEI/siRNA polyplexes. In studies by Chen and coworkers, the shielding

lP

polymer and the polyplexes bind via electrostatic interactions. The conjugation of 2,3dimethylmaleic anhydride (DMMA) to the shield polymer introduces the required pHresponsive property. DMMA is anionic at physiological pH but responds to the acidic

na

environment of tumours. The polyplex is made up bioreducible PEI, which they synthesized by linking branched PEI via dithiobis(succinimidyl propionate). The introduction of the shielding polymer was shown not to interfere with the formation of

ur

the polyplex, although it increased the polyplex size from 80 nm to 91 nm. As expected, the zeta potential also changed from positive (unshielded) to negative (shielded).

Jo

Incubation of shielded and unshielded polyplexes in bovine serum albumin revealed that the change in size of the shielded nanoparticles was significantly less than for unshielded nanoparticles, thus revealing that shielded polyplexes would be less likely to form nospecific interactions in vivo. That said, the incubation period was only 120 min and the shielded particles doubled in size during this time. Since the nanoparticles would need to be stable for at least 24 hours for systemic administration, it would be interesting to see results of extended incubation periods.

51

ro of

Figure 28 Schematic of the charge-reversal of 2,3-dimethylmaleic anhydride, whereby the amide bond breaks when the pH decreases. The shift from anionic to cationic causes the electrostatic shielding at pH 7.4 (blood) and subsequent deshielding of the polyplex at pH 6.6 (tumourous environment). [217]

Cellular internalization at varying pH revealed that the deshielding process facilitates cellular uptake of the polyplexes. Furthermore, gene silencing efficiency was observed

-p

to be improved by detachment of the shielding polymer. In vivo accumulation of fluorescently labelled siRNA delivered by the shielded nanoparticles was measured

re

after systemic administration in a mouse bearing an MDA-MB-231 xenograft. The shielded nanoparticles delivered larger amounts of the siRNA compared to the unshielded polyplexes, resulting in higher accumulation in the tumour tissue. The

lP

antitumour effect of shielded nanoparticles was also investigated via systemic administration of the shielded nanoparticles on mice bearing MDA-MB-231 xenografts.

na

It was established that the shielded nanoparticles significantly inhibited tumour growth compared to the control, unshielded polyplex. Since the charge reversal of these abovementioned shielding polymers by Yang et al.

ur

[217] rely on the same amide pH-response as those described by Kataoke and coworkers, [212-214] the concern regarding the rate of response raised by Chen and

Jo

coworkers [215, 216] would still be an issue. The study by Yang et al. [217] showed that this process takes approximately 30 minutes on the bench.

Charge-reversal polymers present an exciting route for gene delivery. Unlike PEGylation, they address the issues of systemic stability of nanoparticles while not compromising on cellular internalization. That said, obstacles associated with endosomal release and cytoplasmic release of RNA remain equivalent to their polycationic counterparts. In comparison, decationizing polymers, e.g. pDMAEA, address issues of cytoplasmic release, while issues associated with polycations

52

regarding cytotoxicity, nonspecific interactions and systemic instability remain prevalent.

Polymers directly bonded to RNA Table 4 Brief summary of main advantages and disadvantages of directly conjugating RNA to polymers for gene delivery applications

Cons

- Improved RNAi stability

- Modification of RNAi (usually 5'-end) necessary which can impact gene activity and stimulate premature degradation

- N/P is 1 causing low polymer toxicity effects

- Polymer-nucleotide purification can be challenging

ro of

Pros

-p

- Reversible linkers between polymer - Guide-passenger delivery system has and RNAi can be used in order to release no method for release of passenger the nucleotide via a trigger strand

lP

re

- Guide-passenger delivery system allows for use of unmodified passenger - Possible immune response nucleotides There have been some attempts towards the conjugation of neutral polymers directly to nucleotides via labile linkers, e.g. disulfide [218] or pH-labile [123] bonds. This

na

means that the nucleotide is modified, usually on the 5’-end, for this functionalization to be possible. The main pros and cons of the direct conjugation of polymers to RNA(i) has been summarized in Table 4.

ur

Maynard and coworkers [218] synthesized RAFT-polymerized poly(PEG acrylate), a polymer similar to PEG, using a pyridyl disulfide (activated disulfide) RAFT agent.

Jo

Poly(PEG acrylate) has been shown to facilitate passive targeting through increased circulation time. Post-polymerization, this polymer was conjugated to a thiol-modified siRNA. They obtained efficient binding, as well as reversibility of the bond. In a subsequent publication, [219] they combined the polymer-siRNA conjugates with a fusogenic peptide KALA, which is able to facilitate endosplasmic escape, in order to assess the in vitro efficiency of the system. They reported that transfection efficiency was comparable, but slightly lower, than that of Lipofectamine complexed siRNA. Low

53

and high molecular weight poly(PEG acrylate), 6000 and 17 400 g·mol-1, respectively, exhibited little impact on transfection efficiency, while the presence of both KALA and poly(PEG acrylate) components was of utmost importance, with transfection efficiency decreasing significantly in the absence of either of the components. Recently, Lin and Maynard [220] described, for the first time, a study in which the siRNA acted as a macroinitiator for the polymerization, although this strategy had previously been used for DNA. [221] As seen in Figure 29, an ATRP initiator was attached to the siRNA via a disulfide bond, thus introducing a reducible link between the polymer and siRNA for later cleavage within the cytoplasm. Poly(ethylene glycol)

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methyl ether methacrylate (PEGMA) and di(ethylene glycol) methyl ether methacrylate (DEGMA) were grafted from the siRNA via AGET ATRP. These monomers were chosen based on having previously exhibited a nuclease stabilizing effect on siRNA. [219] They reported some low initiator efficiency, causing some siRNA to remain unreacted, however, they obtained good control over molar mass and its distribution.

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This technique represents a possible future alternative towards a polymeric delivery

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device. However, no in vitro or in vivo results have been published to date.

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Figure 29 Schematic of the addition of the ATRP initiator to the siRNA in order to obtain a siRNA-macroinitiator in order to obtain a ‘block copolymer’ via AGET ATRP of PEGMA and DEGMA [220]

Perrier and coworkers [222] used surfactant-free emulsion polymerization to synthesize 2-(2-pyridyldithio)ethylamine modified, carboxyl-α-end HOOC-

poly(acrylamide)-stabilized polystyrene latex particles. The disulfide ligand was then conjugated to a miRNA duplex via a reducible disulfide bond, see Figure 30. Thereafter, they successfully characterized the conjugation as well as showed the release of the conjugated miRNA via incubation in the presence of glutathione. No in vitro results were shown, and therefore further investigation would be required in order 54

to evaluate the true potential of these nanoparticles for gene delivery, especially

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focusing on possible immune response, aggregation and serum stability.

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Figure 30 A cartoon representing the latex nanoparticles synthesized by Perrier and coworkers and subsequently conjugated to a miRNA duplex via a disulfide bond. [222]

The direct chemical modification of nucleotides can cause some difficulties. Side

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reactions may occur on the nucleotide and stimulation of premature degradation of the nucleotide may take place. Furthermore, purification of a polymer-nucleotide conjugate

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can be challenging. To combat this, Averick et al. [223] reported the use of three covalently bonded, biocompatible polymers (PEG-methacrylate-pOEOMA; pOEOMAco-MEO2MA and pOEOMA-co-DMAEMA) to a passenger siRNA via click chemistry,

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in order to obtain a polymer-escort siRNA. The passenger siRNA was complementary to the guide siRNA (the biologically active therapeutic), enabling the guide strand to be annealed to the polymer-escort siRNA. This system allows polymers to provide

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nuclease resistance and facilitate cellular uptake, while the guide siRNA remained unmodified. A limiting factor in this system remains, as the guide strand needs to

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dissociate from the passenger strand in order for it to form the RISC complex, and in so doing allowing gene expression to be regulated. No direct evidence of guide strand release from the duplex is included in the study, however, in vitro experiments resulted in efficient luciferase knockdown for all three polymer escort systems, indicating successful release of the guide strand from the duplex. The investigators indicate that improvements for enhanced guide-strand release could be possible by introducing chemical modifications, e.g. 5’-phosphate, to the guide strand.

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Nanoparticles Table 5 Summary of important advantages and disadvantages of nanoparticles for the delivery of gene therapeutics

Cons - Cytotoxicity due to polymers/crosslinkers still a concern

- Less affected by physiological environment, behaviour dependent only on the polymer structure

- Low biodistribution and renal clearance

- Crosslinkers can be environmentally responsive, e.g. dithiol/acid-labile, allowing for cytoplasmic release

- Possible immune response

- RDPR techniques allow for control over nanoparticle size and morphology, as well as modifications such as introduction of targeting ligands

Cytoplasmic release dependent on crosslinker responsivity RNAi loading efficiency

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Pros - Particle superstructure independent of payload improving structural stability

As polyplex formation is dependent on electrostatic interactions, the superstructure of

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a polyplex is thus reliant on its anionic payload, which, in this case, is the RNA/DNA molecule. Because siRNA and miRNAs are too small to guarantee stability resulting from charge, the polyplexes have an inherent flaw in their make-up. This also means

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that the z-potential of the polyplex is dependent on the siRNA and, subsequently, the aggregation behavior of the system varies depending on its payload. As mentioned

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before, polyplex aggregation is of grave concern. As the physicochemical properties of the superstructure are dependent on the RNA payload, competitive interactions with other polyanions in the physiological environment are also at the mercy of the payload.

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In a dynamic system, for example within the physiological environment, the superstructure of the polyplex may be compromised or weakened due to this

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dependence on the payload. For this reason, some researchers have shifted their attention towards nanoparticles and nanogels [224] with a superstructure independent of their payloads. The key advantages and disadvantages of such systems are underscored in Table 5. Nanoparticles and nanogels possess stable size and morphology, as well as stability that is independent of their payload. The first siRNA clinical trial using nanoparticles as gene delivery tools was published in 2010. [7] The nanoparticles were formulated from PEG that had been conjugated to adamantine, making it able to form inclusion complexes with cyclodextrin. The nanoparticles were

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functionalized with transferrin protein targeting ligands in order to facilitate specific cellular uptake, due to the overexpression of transferrin protein receptors on cancer cells. Thereafter, the particles were administered intravenously and a gene knockdown effect was seen in tumour cells. Siegwart et al. [225] developed a library of cationic core shell nanoparticles, with various hydrophilic surface polymers, to protrude from the nanoparticles and protect the shell. They used crosslinkers with secondary and tertiary amine functionalities in order for these nanoparticles to be able to complex with nucleotides, while still ensuring that the particle formation was not based on polymer-nucleotide interactions.

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Cholesterol was covalently incorporated in the nanoparticles in order to target delivery to the liver. In vivo delivery of nanoparticle-siRNA complexes in mice was observed with accumulation of the nanoparticles in the liver, kidneys and lungs. The incorporation of cholesterol on the nanoparticles successfully enhanced delivery to

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liver hepatocytes.

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Figure 31 Synthetic strategy towards pH-responsive crosslinked nanoparticles for siRNA delivery. Reprinted with permission from Leber et al. (2017) SiRNA-mediated in vivo gene knockdown by acid-degradable cationic nanohydrogel particles. J. Control. Release 248, 10-23

Zentel and coworkers [226] developed cationic, nanoparticle hydrogel superstructures from amphiphilic ester precursor polymers, which aggregate in polar aprotic solvents such as DMSO. They followed Siegwart’s example [225] and cross-linked the hydrophobic inner core with amine functionalized cross-linkers, allowing them to obtain a covalently linked, cationic core for siRNA interaction. Due to a low knock-down efficiency of these nanohydrogel particles, they then adapted these crosslinker

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molecules to contain disulfides or ketal (pH labile) linkers [227] in order to obtain second generation, degradable nanoparticles, Figure 31. This, in turn, produced stimuli responsive nanohydrogel particles that can be destabilized under specific conditions, such as the reductive environment in the cytosol or acidic conditions in the endosome. The particle size and stability is independent of their payload. After in vitro analysis of these nanohydrogel particles, it was ascertained that the disulfide containing cross-linker required relatively high N/P ratios compared to the ketalcontaining cross-linker. These results were in agreement with their previously published results. [228] This higher N/P ratio requirement was due to the need for

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elevated concentrations of glutathione to be present in order for the disulfide bond to break. Such levels of glutathione are predominantly present within the cytoplasm. [229] Thus, the nanoparticle hydrogels had to be transported into the cytoplasm in order to disassemble. Predominantly, this does not seem to successfully occur. [72, 230] Therefore, they continued their in vivo work with the ketal-containing cross-linker and

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saw a significant improvement in knock-down compared to their previously unresponsive nanohydrogel particles. In fact, the in vitro knockdown, caused by these

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nanoparticles being capable of effectively delivering their payload, was comparable to that of Lipofectamine. The nanoparticles were tested on different cell lines for

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cytotoxicity, and no significant toxicity was detected in cells after transfection with concentrations of up to 400 nM.

Subsequently, Zentel and coworkers conducted consecutive injections in order to

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establish more information of the clearance capabilities of the degradable nanoparticles compared to the non-degradable, spermine crosslinked analogues. [227] They discovered that biodegradable ketal-modified particles exhibited superior

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clearance properties. The biodegradable property of these nanoparticles, as well as their high in vitro and in vivo transfection efficiency, make these nanohydrogel particles

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stand out as true contenders in the race for gene therapeutic delivery systems. Of course, one area of concern is the fact that, although the particles themselves are degradable due to the cleavage of the crosslinker, their polymer components are not. This means that, while the nanoparticles degrade, the cationic polymers are still present within the cells post-degradation. These polycations may destabilize the endosomal membrane allowing for endosomal release, however, they are still able to interact with the RNA therapeutics and decrease the knockdown efficiency of the

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therapeutics. Furthermore, the presence of these polymers poses the same cytotoxic threat as the original polyplex systems. Klinker et al. [231] recently published a study describing the formation of bioresponsive polypept(o)ide nanohydrogels synthesized from polypeptoid polysarcosine and Salkylsulfonyl-protected

cysteine

N-carboxyanhydrides,

see

Figure

32.

The

Salkylsulfonyl cysteine promotes micelle self-assembly and provides a chemoselective reactivity towards thiols, allowing for reductive disulfide crosslinking. They demonstrated that these nanohydrogels are able to complex with cholesterol-modified siRNA, while still maintaining their size and morphology and exhibiting an almost

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neutral z-potential. This provides an interesting new platform for gene therapy. However, they did not show any results describing the transfection efficiency of the nanohydrogels. It is plausible that they will suffer from the same challenges as Zentel’s disulfide nanogels described above. Thus, further investigation is necessary in order

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to truly assess the future potential of these nanohydrogels in the field of gene therapy.

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Figure 32 Structure of polypeptoid polysarcosine and S-alkylsulfonyl-protected cysteine N-carboxyanhydrides synthesized for self-assembly and subsequent corecrosslinking into bio-responsive polypept(o)ide nanohydrogels. [231]

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Future outlooks

Extensive work has been conducted towards the development of non-viral, polymeric delivery systems for gene therapy. Nonetheless, despite all efforts, the challenges of the past decade remain prevalent. Although cationic polyplexes allow for efficient uptake, they lack biocompatibility and efficiency in knockdown effects, ultimately limiting their possible use in vivo. The issues of endosomal escape and cytosolic release are still very real and are, therefore, the limiting factor within any delivery system. Thus, they should remain the focal point when designing gene delivery

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systems in the future, while also keeping cytotoxicity in mind. It is probable that future research will see a marked shift towards vectors in the form of neutral or amphiphilic nanoparticles/nanogels,

possibly

accompanied

by

further

development

of

chargereversal or sheddable shielding polymers. Previously published research has also convincingly indicated that the topology of the polymer systems can play a very significant role not only on transfection efficiency but also on cytotoxicity. Therefore, the impact that RDRP will surely have on the future of polymeric gene delivery systems cannot be emphasized enough. Once the issues of endosomal and cystosolic release have been properly addressed,

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the next, and no doubt the greatest hurdle will be in vivo targeted delivery. Current in vivo results still see accumulation of delivery systems in the organs of elimination for renal clearance, rendering the delivery insufficient for targeting specific organs. It would be very exciting if it were possible to move from enhanced cellular uptake to

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true targeted delivery to specific organs.

Funding Acknowledgement

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BK and AK acknowledge funding from the South African Research Chairs Initiative of the Department of Science and Technology (DST) and National Research Foundation

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(NRF) of South Africa (Grant No 46855). This work was supported by the Project EVICARE(#725229) of the European Research Council (ERC) to JS.

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Conflicts of Interest

The authours declare no conflict of interest.

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