1 Introduction

Nanotechnology and nanoscience are widely seen as having a tremendous promise to revolutionize the scientific landscape in terms of research and applications. Among various approaches, nanocarriers as a tool of nanotechnology offer great advantages for biomedical applications, viz., drug delivery, sensing, and imaging (Singh and Lillard 2009). Concurrently, the applications of nanotechnology in target-specific delivery of drug and gene have opened up new areas of research in nanomedicine.

Nanotechnology is the design of objects with dimensions conveniently described in units of nanometers (10−9 m) (Thassu et al. 2007). The dimensions of nanoparticles (NPs) lie between 10 and 1000 nm. Objects in the nanometer size range often exhibit properties that are not found in bulk materials of the same composition. However, particles >200 nm are not widely used, and nanomedicine holds interest in particles <200 nm as physicochemical properties of the particles in this range make them attractive for commercial and medical development (Singh and Lillard 2009). Nanocarriers for drug and gene delivery can be classified as lipid-based (solid lipid NPs, liposomes), polymer-based (polymeric micelles, dendrimers or polymeric NPs), and inorganic NPs (metallic NPs, silica NPs (SiNPs), carbon nanotubes (CNTs)) (Estanqueiro et al. 2015). The selection of material for development of nanocarrier is highly dependent on encapsulated payload/bioactives, therapeutic or diagnostic goal, and route of administration (Jabr-Milane et al. 2008).

Nowadays nanosystems focus primarily on the development of target-specific and slow but controlled drug release systems. A major milestone has been achieved in drug delivery systems with the development of technologies that can mask the nanocarriers from the immune system. Specific decoration of nanocarriers has significantly increased the half-life of the drug. The introduction of synthetic lipid derivatives of polyethylene glycol (PEG) confers “stealth” capability on nanocarrier system, due to the hydrophilicity of the PEG chains. This PEG stealth avoids opsonization and reduces fast blood clearance by immune recognition, hence helping in passive accumulation in the tumors via an enhanced permeation and retention (EPR) effect (Gref et al. 1995, 2000; Mosqueira et al. 1999). The first generation of nanosystems was nanometric liposomes. They are able to encapsulate both hydrophilic and hydrophobic drugs (Patel et al. 2015). The next generation were polymeric NPs precipitated with drug molecules and surface functionalized that prevented the immune recognition and nowadays even having targeting moieties in form of antibodies or folic acid. They have potential to enhance therapeutic benefit while reducing side effects as compared to free drug. Furthermore, inorganic NPs are being widely explored because of their chemical, physical and structural characteristics. This chapter considers current status and possible future directions of conventional and engineered pharmaceutical nanocarriers used for drug and gene delivery.

2 Lipid-Based Nanoparticles

Among the group of several nanoformulations, lipid nanoformulations have received much attention due to its biocompatible and biodegradable nature. They have huge potential for delivery of drugs and genes in several diseases. In the field of nanomedicine, they offer interesting benefits, viz., enhancing drug efficacy and providing controlled and convenient drug release. The performance of lipid nanoformulations greatly depends upon the composition and structure of formulations. Lipid-based nanocarriers may be further categorized as liposomes, solid lipid nanoparticles (SLNs), and nanostructured lipid carriers (NLCs) (Lim et al. 2012) (Fig. 6.1). Several lipid-based formulations have already entered in the market, and others are in preclinical and clinical phase (Table 6.1).

Fig. 6.1
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Structure of lipid-based nanocarriers

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Table 6.1 Clinical status of liposomal formulations
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2.1 Liposomes

Almost from the time of their first report in the 1960s by Bangham and coworkers, liposomes have been extensively used by researchers as prospective carriers for various bioactive molecules. Liposomes are self-closed spherical phospholipid vesicular system having size ranges of 20–200 nm. These lipid vesicles are derived by the dispersion of phospholipids in an aqueous media. Various techniques pursued to disperse the lipids into the aqueous media include thin lipid film hydration, extrusion, sonication, etc. The basic component of liposome is formed by phospholipids or any similar amphipathic lipids (having both hydrophilic head and hydrophobic tail) (Onyuksel et al. 2006; Rawat et al. 2008). The phospholipid molecules used are either naturally occurring and/or derived from synthetic sources such as phosphatidylcholine, phosphatidylethanolamine, phosphatidylserine, phosphatidylglycerol, polyethylene glycol, cholesterol, distearoyl-sn-glycero-3-phosphoglycerol, and dipalmitoyl-sn-glycero-3-phosphothioethanol (Arias 2013).

Liposomes have been broadly classified on the basis of their size, number of bilayers, composition, and method of preparation employed, viz., multilamellar vesicles (MLV), large unilamellar vesicles (LUV), small unilamellar vesicles (SUV), conventional liposomes (CL), pH-sensitive liposomes, cationic liposomes, long circulating liposomes (LCL), reverse phase evaporation vesicles (REV), French press vesicles (FPV), ether injection vesicles (EIV), etc. (Samad et al. 2007; Pattni et al. 2015; Matougui et al. 2016).

Liposomes possess superior capability of delivering drugs to target site and thus overcomes the adverse effect and resistance of drug due to reduced non-targeted circulation of the drug. Zhang et al. (2016a) prepared oxaliplatin and irinotecan hydrochloride co-loaded liposomes for the treatment of colorectal cancer. In vitro cytotoxicity results demonstrated that co-loaded liposomes showed higher cytotoxicity as compared to single loaded liposomes in both CT-26 and HCT-116 cells. Also, co-loaded liposomes exhibited greater antitumor therapeutic activity in CT-26-bearing BALB/c mice. In vivo safety evaluation showed that liposomes had less toxicity as compared to free drug solution (Zhang et al. 2016a). Recently, a liposome-based theranostic system was developed for delivery of docetaxel and quantum dots (QD) simultaneously. In this study, liposomes were conjugated with arginine-glycine-aspartic acid (RGD) to target them effectively in brain for theranostic applications. Brain distribution study confirmed higher accumulation of targeted liposomes in comparison to Docel™ and free QD. Results of brain histopathology were also in concordance to above results (Sonali et al. 2016). More recently, the effectiveness and safety profile of trastuzumab and liposome-loaded doxorubicin and paclitaxel combination was evaluated for early or locally advanced breast cancer therapy, clinically. The study was carried out on 30 breast cancer patients, and treatment was found to be well tolerated with very few adverse effects suggesting good therapeutic option for treatment of cancer patients (Uriarte-Pinto et al. 2016). Owing to the several beneficial properties, several liposome-based formulations has been approved by the FDA for various anticancer agents such as daunorubicin (DaunoXome®, Doxil®/Caelyx®), cytarabine (DepoCyt®), vincristine (Marqibo®), etc. (Lian and Ho 2001).

Liposome-mediated gene transfer has also been found as a promising approach not only in the treatment of diseases with genetic disorders but also in the development of strategies for treatment of several fatal diseases, viz., cancer, degenerative disorders, and AIDS. For instance, Landen et al. (2005) developed 1, 2-dioleoylsn-glycero-3-phosphatidylcholine (DOPC)-encapsulated siRNA liposomes using tertiary-butanol and nonionic detergent Tween 20. Developed system was found to be effective in reducing EphA2 expression 48 h after administration of a single dose in an orthotopic model of ovarian carcinoma suggesting its applicability (Landen et al. 2005). Bubble liposomes (containing perfluoropropane gas) possess specific advantage of high gene loading with improved intracellular penetration. Sugii and coworkers formulated the plasmid DNA-loaded bubble liposomes. The developed system was specifically targeted to neovessels via conjugating cyclic RGD peptides on their surface and exhibited high gene transfection efficiency to human umbilical vein endothelial cells (HUVECs) (Sugii et al. 2016). Similarly, folic acid modified liposomes were also loaded with plasmid DNA and binding affinity of gene was modified by agarose gel electrophoresis assays. Gene transfection efficiency was studied against NCI-H460 cells. The results showed improvement in gene transfection efficiency after attaching the targeting moieties (Cui et al. 2016).

2.2 Solid lipid Nanoparticles

SLNs were first developed in 1991 and are used as alternative carrier systems to traditional colloidal carriers, such as emulsions, liposomes, and polymeric NPs (Muller et al. 2002). SLNs consist of 0.1–30% (w/w) lipid, dispersed in an aqueous solution of 0.5–5% (w/w) of surfactant and are solid at both room and body temperature (Rawat et al. 2008). These particles are prepared by using number of solid lipid materials such as mono-, di-, and triglycerides; fatty acids; waxes; and steroids. The lipid materials employed should possess favorable properties such as biocompatibility, biodegradability, and low toxicity. Various surfactants employed for steric stabilization includes phospholipids, poloxamers, and polysorbates. Most commonly, SLNs are fabricated using hot or cold homogenization technique. The average diameter of the SLNs ranges from approximately 40–1000 nm (Mehnert and Mader 2001).

SLNs have gained popularity as these combine advantages of various “soft” drug carriers such as emulsions, liposomes, and polymeric NPs and at the same time avoid or minimize some of their drawbacks. They have progressively been explored for the encapsulation of labile hydrophilic and hydrophobic drugs protecting them from degradation in the body and for sustained release. Some of the drug classes that are being investigated with this system includes antibacterial, antiparasitic, antioxidant, anticancer, antiviral, antiandrogenic, antihypertensive, anti-inflammatory, antipsychotic agents, vitamins, and various bioactive compounds having multi-potential (polyphenols, flavonoids, carotenoids). Numerous studies have shown that bioavailability of drugs loaded in SLNs can be extensively improved and also provide site specificity. For example, Bhandari and Kaur prepared isoniazid-loaded SLNs to improve its oral bioavailability. The study showed 6 and 4 times higher relative bioavailability of the drug in plasma and brain, respectively, when compared to the free drug solution at the same dose in rats. The high plasma drug concentration reported was attributed to the lipidic transport of SLNs through lymphatics and thus bypassing first pass metabolism of isoniazid (Bhandari and Kaur 2013). Pandita et al. successfully developed resveratrol loaded stearic acid-based SLNs using a mixture of surfactants (lecithin/poloxamer 188) as stabilizers. Pharmacokinetic results demonstrated that oral bioavailability of resveratrol was 8.035-fold high in male Wistar rats when compared to its pure suspension (Pandita et al. 2014). Also, in a study, paclitaxel loaded SLNs when administered orally in male Swiss albino mice exhibited tenfold high bioavailability as compared to paclitaxel solution, and the toxicity studies confirmed the relatively safe nature of the SLNs carrier systems with or without drug (Pandita et al. 2011). In another study, curcumin-loaded SLNs prepared with liquid lipid Sefsol-218® showed higher bioavailability and prolonged inhibitory activity in cancer cells. After the i.v. administration to rat, SLNs were capable of enhancing the bioavailability of curcumin to 1.25-fold compared to free drug (Sun et al. 2013).

Although, exhilarating results have been obtained with SLNs as mentioned above, various biological barriers, viz., rapid clearance, serum instability, and nonspecific uptake by the mononuclear phagocytic system, are associated with conventional SLNs. PEGylation was introduced for surface functionalization of SLNs to provide a hydrophilic layer resulting in increased circulation time via overcoming of opsonization process (Uner and Yener 2007). In a study, Madan and colleagues prepared noscapine-loaded SLNs and noscapine-loaded PEG conjugated SLNs. After i.v. administration to mice, the plasma half-life was significantly enhanced up to ~11-fold and ~5-fold in case of noscapine-PEG-SLNs and noscapine-SLN, respectively, in comparison to free drug suggesting applicability of surface modification (Madan et al. 2013).

Since early 2000, SLNs have been investigated for delivery of genetic material in several disease treatments (Olbrich et al. 2001). Radaic et al. formulated SLNs for gene delivery using factorial design approach wherein different formulation parameters were varied w.r.t. DNA load, colloidal stability, in vitro cytotoxicity, and transfection efficiency in prostate cancer cells. Results suggested that the concentration of lipids and surfactant employed affected the size, stability, and transfection efficiency (Radaic et al. 2015). Yu et al. prepared paclitaxel-loaded cationic SLNs and fabricated its complex with human MCL1-specific siRNA. They showed that co-delivery of paclitaxel and MCL1-specific siRNA using cationic SLNs enhanced anticancer efficacy both in vitro and in vivo when compared to each agent alone (Yu et al. 2012). In spite of several beneficial properties as drug-carrier, SLNs have lots of drawbacks such as limited drug loading and leakage of drug during storage. Further, NLCs were developed as next generation of lipid drug carrier systems at the end of 1990s to overcome the drawbacks of SLNs.

2.3 Nanostructured Lipid Carriers

NLCs may overcome the limitations associated with SLNs such as low drug payloads, leakage of drug during storage, and stability (Weber et al. 2014). They are mainly produced by mixing of different lipids, i.e., solid lipids (glyceryl behenate, glyceryl palmitostearate, glyceryl monostearate/monostearin, cetyl palmitate, and stearic acid) with liquid a lipid (caprylic and capric fatty acids with a minor level of caproic, lauric, myristic, corn oil, coconut or palm kernel oils, etc.) which remains solid at room temperature (Naseri et al. 2015). NLCs show higher entrapment efficiency because solid lipids provide higher space for hold of drug and higher solubility of drugs in liquid lipids in comparison to solid lipids (Poonia et al. 2016). NLCs have been investigated for parenteral, oral, pulmonary, and topical delivery of drugs and have also presented good potential in gene delivery.

The attractive properties of NLCs on skin offer increase in drug penetration mainly through occlusion effect. They transfer drugs into deeper layers of the skin with reservoir action and provide sustained release (Muller et al. 2007). Sweety and coworkers formulated NLCs of an anti-acne drug, azelaic acid, using solvent diffusion-solvent evaporation method for enhancing its dermal retention and overcoming its adverse effects. Results of skin retention study showed high retention of azelaic acid from NLCs gel (63.96 ± 4.45%) followed by plain azelaic acid gel (15.12 ± 3.2%) and drug solution (4.78 ± 1.1%) in rat skin (Sweety et al. 2015). NLCs are also employed for the treatment of the damaged or inflamed skin due to nonirritant and nontoxic properties of NLCs. Valdecoxib-loaded NLC carbopol gel demonstrated no skin irritation while the marketed gel of valdecoxib caused slight irritation after 48 h (Joshi and Patravale 2006). NLC-based gel also showed prolonged release activity as compared to market gel. Further, chitosan-coated NLCs showed 7.7-fold higher flurbiprofen residence time on the cornea when compared to uncoated NLCs. Also, transcorneal penetrations were increased up to 2.4-fold compared to uncoated NLCs (Luo et al. 2011). In a short period of time, NLC technology has come out as a big boon for topical delivery and two NLC-based products are available in the market (Cutanova Nanorepair Q10 and FloraGLO®).

Another widely employed route of administration for NLCs is the oral route. Due to the increased drug loading capacity of NLCs, most of these studies have focused on the ability of NLCs to improve the oral bioavailability of poorly water-soluble drugs. For example, lovastatin-loaded NLCs made from mixtures of precirol and squalene were able to promote the oral absorption and led to increased bioavailability of lovastatin. More than 70% of lovastatin was entrapped in the NLCs, which was significantly higher compared to the SLNs. Results of in vivo studies on rats showed that NLCs produced a significant improvement in the bioavailability compared to the free solution (Chen et al. 2010). In another study, saquinavir, a P-gp substrate, loaded NLCs were prepared and transport mechanisms across Caco-2 cells were studied. A 3.5-fold increase in drug permeability was reported in case of NLCs than the drug suspension (Beloqui et al. 2013). In another study, a ~2.5- and ~3-fold higher bioavailability was achieved with silymarin-loaded NLCs than the marketed formulation (Legalon®) and free drug (Zhai and Zhai 2014). Yang et al. (2013) prepared hyaluronic acid-coated NLCs using electrostatic attraction for targeted delivery of paclitaxel at tumors’ site and investigated the in vitro cytotoxicity and in vivo antitumor efficacy against three CD44-overexpressing cell lines. The result of hyaluronic-coated NLCs showed better antitumor efficacy in B16-bearing Kunming mice when compared to Taxol®. Further, hyaluronic-coated NLCs showed higher accumulation in tumor and enhanced the circulation time of paclitaxel in blood.

NLCs as gene carriers can easily penetrate through the biological membranes effectively because lipids are the main components of cell membranes boosting the uptake of genetic compounds (Pathak et al. 2009). In a study, Zhang and coworkers prepared a novel system by modifying NLCs using cetylated polyethylenimine (PNLC). PNLCs showed high transfection efficiency in human lung adenocarcinoma cell line SPC-A1 and Chinese hamster ovary (CHO) cells as compared to free drug. The transfection efficiency of the optimized PNLC was similar to marketed formulation (Lipofectamine™2000) (Zhang et al. 2008). More recently, Han and coworkers (2016) developed a modified plasmid-containing enhanced green fluorescence protein (pEGFP)-loaded NLCs with transferrin (Tf) as an excellent active targeting ligand for improving the A549 cell targeting ability of the carriers. Tf-NLC/pEGFP showed higher gene expression efficiency in cancer cells both in vitro and in vivo than unmodified NLC/pEGFP suggesting enhanced activity after loading into targeted NLCs.

3 Polymer-Based Nanoparticles

A long chain that is made up with a repeated subunit by joining together larger macromolecules or micromolecules is called a polymer. Mainly two types of polymers have been used for the development of polymer-based nanocarriers, i.e., natural and synthetic. Some examples of natural and synthetic polymers include albumin, dextran, hyaluronate, chitosan, and poly(lactic-co-glycolic acid) (PLGA), poly(lactic acid) (PLA), poly(cyanoacrylate) (PCA), poly(N-(2-hydroxypropyl) methacrylamide) (PHPMA), PEG, polyethyleneimine (PEI), etc., respectively (Haag and Kratz 2006). On the basis of method of preparation, various types of polymeric-based nanoformulations have been obtained such as polymeric NPs, nanospheres or nanocapsules, micelles, polymer drug conjugates, etc. (Fig. 6.2). The average size ranges of such nanocarriers are from 10 to 1000 nm (Peer et al. 2007).

Fig. 6.2
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Structure of polymer-based nanocarriers

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3.1 Polymeric Micelles

Polymeric micelles are spherical, colloidal, and nanoscopic core/shell structures and usually have narrow size distributions with diameters ranging from 10 to 100 nm (Jhaveri and Torchilin 2014). They are formed via self-assembly of amphiphilic copolymer chain in aqueous milieu. They present a core and shell architecture; the inner core is hydrophobic part which serves as a microenvironment for poorly water-soluble drugs, whereas the outer hydrophilic corona protects drug from aqueous environment and stabilizes interface between the core and the external medium (Kedar et al. 2010). This self-assembly monomer provides the main driving force behind the micellization process. Polyion complex micelles are subclass of polymeric micelles and are obtained from electrostatic interactions between oppositely charged copolymer (one is charged segment and other is neutral polymer chain) and drug followed by self-assembly of charge-neutralized blocks (Harada and Kataoka 1995; Kataoka et al. 1996). Both polymeric micelles and polyion complex micelles can serve as reservoir for drugs, which may be loaded by different methods, i.e., chemically, physically, or electrostatically, depending on the chemistry of drug and core forming block.

Many drugs present low aqueous solubility and poor accumulation at their target site which lower their efficacy and can promote systemic adverse effects. In addition, macromolecular drugs such as peptides, DNA, and RNA suffer from premature degradation upon administration, low bioavailability, and inefficient cellular entry, compromising their therapeutic outcome. The versatile attributes of polymeric micelles make these systems an attractive nanocarrier for the delivery of these therapeutics. Due to the nonionic hydrophilic corona, polymeric micelles can prevent the adsorption of opsonins, thereby limiting the rapid uptake by the mononuclear phagocyte system (MPS) and prolonging the circulation half-life of the encapsulated drug (Jones and Leroux 1999). Hamaguchi et al. (2005) showed that the incorporation of paclitaxel in PEG-b-P(Asp) polymeric micelle resulted in about 90- and 25-fold increase in plasma concentration and tumor area, respectively, compared to the free drug. This remarkable increase in the AUC, in spite of using equivalent doses, can be ascribed to the greater stability conferred by the micelles which permits long circulation and minimizes the drug leakage (four to six times longer elimination half-life vs. the free drug) and thus allow the passage and accumulation of the drug in the tumor. Polymeric micelles can be functionalized with targeting ligands in order to achieve active targeting. Perche et al. (2012) demonstrated superior cellular uptake and cytotoxicity of doxorubicin when loaded in antibody-conjugated (anti-nucleosome 2C5) poly(ethylene)glycol-phosphatidylethanolamine micelles in comparison to free drug and conjugated micelle system. Also, developed system was found to be efficient in drug-resistant ovarian cancer cells.

In addition to modifying the pharmacokinetics and biodistribution, and decreasing the systemic toxicity of drugs, polymeric micelles have been also found to stabilize their incorporated drug, as in the case of amphotericin B. Amphotericin B is a membrane disruptive drug used to treat systemic fungal diseases in its monomeric form. Its direct administration, however, is associated with self-aggregation, leading to loss of selectivity and systemic toxicity. Adams et al. have shown that when amphotericin B was incorporated into the core of PEG-b-poly(N-hexyl-laspartamide-stearic acid) micelles, it was stabilized in its monomeric form, preventing the non-selective hemolysis of mammalian red blood cells in vitro (Adams et al. 2003). Moreover, the drug-loaded polymeric micelle retained in vivo antifungal activity as compared to the standard clinical Fungizone® formulation (Barratt and Bretagne 2007). In another study, Ge et al. formulated polyplex micelle modified with PEG-polycation block copolymer for gene delivery. Furthermore, in order to enhance targeted delivery and to promote cellular uptake, cyclic RGD peptide was attached at distal end of PEG. It was found that circulation time was successfully increased through PEG conjugation. Thus, constructed cyclic RGD conjugated polyplex micelle with the prominent PEG shielding resulted in both improved accumulation at tumor site and provided efficient gene expression of antiangiogenic protein (sFlt-1) through intracellular trafficking (Ge et al. 2014).

3.2 Dendrimers

Dendrimers are highly branched, three-dimensional polymeric architecture with many arms originating from a central core (Gillies and Frechet 2005). They have an internal cavity, in which drug can be incorporated and also large number of functional groups present on the outer surface can be used to bind variety of drugs and other functional groups. Dendrimers can be prepared by iterative synthetic methodology, i.e., divergent synthesis or convergent synthesis (Nanjwade et al. 2009). The void area within dendrimer and its ease of modification makes it applicable for drug and gene delivery, imaging, and boron neutron capture therapy. Dendrimers generally have a uniform structure, with the prospective to create an isolated “active site” core area through chemical functionalization. Modification of branching may allow for incorporation of molecule within this system. For example, dendrimers may become water soluble when outer surface is functionalized with hydrophilic groups, such as carboxylic acids (Faraji and Wipf 2009). Cytotoxicity of dendrimers strongly depends on its surface modification and core material. For example, cytotoxicity can be reduced by changing the surface amine groups with hydroxyl groups (Wilczewska and Niemirowicz 2012). It was found by Singh et al. that doxorubicin dendrimers conjugated with biodegradable polymer were ten times less toxic than free doxorubicin toward colon carcinoma cells (Singh and Lillard 2009).

Among many different types of dendrimers, poly(amidoamine) (PAMAM) has been one of the commonly used systems and shown significant applications in pharmaceutical and biomedical fields (Madaan et al. 2014). Although it has many advantages in drug delivery, it has several issues which limit its clinical applications. PAMAM is cytotoxic and rapidly removed from systemic circulation due to presence of amine groups on the surface. This drawback is one of the key barriers for the clinical use of dendrimers. PAMAM surface has been modified using different techniques such as PEGylation, carbohydrate conjugation, acetylation, and amino acid or peptide conjugation for enhancing the biocompatibility and drug encapsulation as well as targeting ability. Various studies have been investigated w.r.t. the low cytotoxicity and enhanced transfection efficiency of modified PAMAM dendrimers (Bae et al. 2016; Luong et al. 2016). Yu et al. developed PAMAM dendrimers modified with histidine and arginine, which resulted in higher gene transfection efficiency when compared to unmodified PAMAM dendrimers. They also found that histidine unit attached with PAMAM dendrimers improved its proton buffering capacity in the range of 3.5–6. The histidine and arginine modified dendrimers also showed less cytotoxicity as compared to PEI modified dendrimers. This proves that PAMAM dendrimers modified with histidine and arginine may provide a promising gene carrier system (Yu et al. 2011). Dutta et al. studied toxicity profile of conventional and surface engineered fifth generation poly(propyleneimine) (PPI) dendrimers. They were evaluated after administering 2.5 mg/kg, 25 mg/kg, and 250 mg/kg doses via i.v. route to Wister rats. Hemolytic studies depicted a decrease in RBCs and hemoglobin content in conventional PPI dendrimers. Biochemical analysis also showed an increase in level of serum biochemical parameters indicative of the hepatotoxic effect of PPI. However, functionalized PPI dendrimers showed no signs of toxicity suggesting reduction of toxicity followed by functionalization (Dutta et al. 2008). In another study, fifth generation PAMAM dendrimer was functionalized with hydrophobic chains (12–16 C-alkyl chain) for delivery of plasmid DNA in mesenchymal stem cells. In vitro results demonstrated that functionalized dendrimers had low level of cytotoxicity due to presence of alkyl group on the outer surface. Further, it was found that smallest carbon chain of alkyl group showed higher efficiency for gene delivery (Santos et al. 2010).

3.3 Polymeric Nanoparticles

Polymeric NPs have great potential to enhance therapeutic benefit while reducing side effects of the free drug. These NPs are generally aqueous dispersions or dry powders obtained by freeze/spray drying. They overcome the stability and premature drug release issues associated with liposomes and emulsions and also facilitate prolonged drug release (Kumari et al. 2010). Polymeric NPs are prepared from PLA, PLGA, etc., possess excellent biodegradability and low toxicity, and consist of drug dispersed in an amorphous form within a polymer matrix; such particles could be prepared as nanosphere, wherein the drug is dispersed uniformly throughout the matrix of the particle or as nanocapsules. They have been studied for localized as well as controlled drug delivery. For example, carmustine-loaded polymeric implant disc made by polyanhydride polymer was used for the treatment of brain cancer. Such disc is placed into the brain to follow controlled release pattern while preventing tumor regrowth (Domb et al. 1999).

3.3.1 PLGA Nanoparticles

Among various polymeric NPs, PLGA-based NPs are particularly attractive in biomedical applications. PLGA is a frequently used biodegradable polymer as its hydrolytic degradation produces the original monomers, lactic acid, and glycolic acid which are side-products of various metabolic pathways. The degradation time of PLGA is controlled by the ratio of lactic acid and glycolic acid: the higher the content of glycolide units, the less is the time required. However, the copolymer with 50:50 monomers ratio reveals the fastest degradation. Because of its minimal toxicity and ability for providing greater control over the drug release profile, PLGA is suitable for drug delivery (Astete and Sabliov 2006). PLGA NPs are versatile among various polymeric NPs systems because of their biocompatibility for drug targeting at the cellular level (Mohamed and van der Walle 2008). Mathew et al. (2012) developed curcumin-loaded PLGA NPs conjugated with Tet-1 peptide for the treatment of Alzheimer’s disease. Due to antioxidant and anti-amyloid activity of curcumin, curcumin-loaded NPs were successfully used in the treatment of Alzheimer’s disease. In vitro cytotoxicity studies toward LAG cell lines demonstrated that Tet-1 conjugated or unconjugated NPs showed no cytotoxicity while exhibiting antioxidant property. Furthermore, cell uptake was performed by using GI-1 glioma cells for Tet-1 conjugated and unconjugated NPs. It was found that Tet-1-conjugated PLGA NPs showed multiple-fold increased uptake than the unconjugated NPs. Moreover, confocal microscopy studies showed that Tet-1 targeted NPs distributed more around the cell soma and nucleus compared to unconjugated NPs. These results demonstrated that Tet-1 conjugated PLGA NPs were proved to have great potential as therapeutic carrier of curcumin for Alzheimer’s disease treatment (Mathew et al. 2012). Yu et al. (2015) developed paclitaxel-loaded PLGA NPs complexed with PEI conjugated with Herceptin onto the surface of NPs through electrostatic interaction. In vitro cell line study toward HER-2 cancer cells showed that Herceptin-loaded polymeric NPs have high degree of cytotoxicity when compared with unmodified polymeric NPs. Furthermore, they also fabricated Herceptin-loaded polymeric NPs through chemical conjugation. It was found that chemically conjugated Herceptin polymeric NPs showed less cytotoxicity and cellular uptake efficiency in comparison with electrostatically bound Herceptin. These results demonstrated that electrostatic interaction provided an appropriate method for the synthesis of protein-loaded NPs (Yu et al. 2015).

3.3.2 PLA Nanoparticles

PLA is linear aliphatic polyester derived from lactic acid monomers and is widely used in micro- and nanoparticulate drug delivery systems because of its biocompatibility and biodegradation properties. It possesses the Generally Recognized as Safe (GRAS) status of the FDA (Kumari et al. 2010; Athanasiou et al. 1996). Among all biopolymers, PLA has been widely investigated in drug delivery systems since the 1980s due to its extensive biodegradability nature. For instance, Zhu et al. successfully developed D-alpha-tocopherol polyethylene glycol 1000 succinate (TPGS)-PLA NPs modified with polydopamine for the treatment of liver cancer. In this study, docetaxel was used as a model drug and galactosamine was conjugated on the surface of modified polymeric NPs to enhance the targeted delivery of drug to cancer cells. In vitro cellular uptake and cytotoxicity studies showed that galactosamine-conjugated polymeric NPs target HepG2 cells and inhibit the growth of cancer cells in comparison to Taxotere®. Furthermore, in vivo results demonstrated that drug-loaded polymeric NPs reduced the tumor size most extensively on hepatoma-bearing mice (Zhu et al. 2015). In another study, Tan et al. (2014) fabricated PLA-TPGS NPs to reduce the antagonist effect of docetaxel and tamoxifen. It was found that co-delivery of both anticancer drugs in PLA-TPGS NPs can significantly enhance the cytotoxicity toward MCF7 cancer cells. Moreover, decrease in IC50 in MCF7 cell line showed that polymeric NPs have the potential to enhance the synergistic effects by reducing antagonistic effect.

In gene delivery studies, DNA is coupled to NPs to create non-viral vectors protecting the DNA in body conditions. A size of around 100 nm and the positive charge of the particles facilitate the delivery and interactions with the cell membrane, respectively. Qian et al. (2014) developed different shaped PLA-polydimethylaminoethyl methacrylate (PDMAEMA) nanoarchitectures for the co-delivery of miR-21 and doxorubicin to treat glioma. It was found that star-shaped copolymer formed nanosized micelles and had positive charge and low cytotoxicity toward GES-1 cells in comparison to control PEI 25 kDa. Also, they had higher gene transfection efficiency (2.5 times) than the control PEI. The results depicted that star-shaped copolymers are promising candidate for the gene delivery and hydrophobic therapeutics.

3.3.3 Chitosan Nanoparticles

Chitosan NPs are also one of the most accepted polymer-based NPs used in drug delivery and targeting applications. Due to its hydroxyl and amine functional groups, the chemical modifications of chitosan have been extensively investigated in literature in order to improve its physicochemical properties, without changing its fundamental skeleton. The main goals of chemically modifying chitosan are to provide derivatives that are soluble at neutral and basic pH values, to control hydrophobic, cationic, and anionic properties as well as to attach various functional groups and ligands and optimize the process of drug release. Furthermore, the hydrophobic modification of cationic polymers has shown facilitated dissociation of polymer/DNA, enhancing the release of DNA to cytoplasm that would otherwise remain strongly bound through ionic interactions between phosphate groups of DNA and cationic units of polymers (Maximilien et al. 2015). The chemical modification of this cationic natural polymer, without changing its main properties, could be a huge approach for enhancing the efficiency of transfection. Recently, Rudzinski et al. (2016) formulated PEGylated chitosan NPs for delivery of siRNA with improved transfection efficacy in colon cancer cells. Confocal studies depicted high level of fluorescent tagged RNAs into cells showing better effect in treatment of colon cancer.

Hybrid lipid-based polymeric NPs provide several opportunities in drug delivery and gene delivery due to their physicochemical properties in terms of drug encapsulation, size as well as low immunogenicity, and no risk of transmission of infectious diseases (Pandita et al. 2015a). The versatility of this lipid-polymer hybrid NP platform allows for surface chemistry modifications. Ewe and Aiqner (2016) used cationic lipids in their hybrid NPs in order to form a DNA complex for gene delivery. In a study by Morales et al. (2009), the end group on the PEG that makes up the corona can be changed from a carboxyl group to an amine or a methoxy group in order to change the surface zeta potential. It was also shown that the surface chemistry of the hybrid NPs affects human plasma and serum absorption patterns by inducing different levels of compliment activation. The compliment and coagulant activation studies exhibited the potential for the lipid-polymer hybrid NPs to be a viable immunocompatible delivery option. Another type of surface chemistry modification is the addition of targeting ligands, which are used to increase cellular uptake and accumulation in the tumor sites. Different types of ligands are used to target hybrid NPs to cancer cells such as antibodies, proteins, small molecules, aptamers, and peptides (Wang and Thanou 2010). Zhao et al. (2015) synthesized lipid-polymer hybrid NPs to co-deliver HIF1α and gemcitabine for the treatment of pancreatic cancer. It was found that lipid-polymer hybrid NPs have higher stability and longer circulation time than polymer NPs. Furthermore, drug-loaded lipid-polymer hybrid NPs exhibited synergistic antitumor effect in vivo than drug-loaded polymer NPs.

3.4 Polymer-Drug Conjugates

Ringsdorf proposed the idea of polymer-drug conjugates in the mid-1970 that could enhance the delivery of an anticancer drug to a tumor. Polymer-drug conjugate is a drug delivery technology in which a drug is covalently attached to a polymeric carrier, mostly via biodegradable linker (Haag and Kratz 2006). These linkages might be cleaved to release the pharmacologically active moieties in a particular condition such as pH, temperature, and osmolality around the tumor. Polymers such as PEG, poly-styrene-maleic anhydride copolymer (SMA), N-(2-hydroxypropyl)-methacrylamide copolymer (HPMA), and poly(a, L-glutamic acid) (PG) have been investigated with different cytotoxic drugs (doxorubicin, paclitaxel, gemcitabine, camptothecin, irinotecan, cisplatin, etc.). More than 15 polymer-drug conjugates have entered in clinical trials till now (Table 6.2). The main benefits of polymer-drug conjugates compared to the free drug include EPR effect, reduced toxicity, improved solubility in biological fluids, ability to overpass some mechanisms of drug resistance, and increased half-life of the drug (Duncan 2006). For instance, Vasey et al. (1999) reported the conjugation of HPMA copolymer with doxorubicin (PK1). They found that PK1 demonstrated antitumor activity with significantly lower cardiotoxicity and alopecia when compared to unconjugated doxorubicin. Chytil et al. (2006) developed water-soluble polymer-drug conjugate for cancer therapy. This agent consisted of anticancer drug doxorubicin conjugated to a HPMA copolymer through the hydrolytically degradable hydrazone bond. The system was incubated at pH 7.4 and 5.0 in order to enable pH-controlled release of the drug. At pH 7.4 conjugates were found to be stable and sustain the drug release while at pH 5.0 burst release of drug was observed. Moreover, in vitro cell line study showed higher cytotoxicity of polymer-drug conjugate system when compared to the free drug suspension.

Table 6.2 Clinical status of polymer-based nanocarriers
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4 Inorganic Nanocarriers

Inorganic nanocarriers are emerging nanosystems utilizing inorganic nanomaterials to achieve medical breakthroughs in therapeutics and advanced diagnostics (Sekhon and Kamboj 2014; Huang et al. 2011; Malmsten 2013). These include metallic NPs, SiNPs, CNTs, QD, etc. (Fig. 6.3). Due to large size and anionic nature, these systems can reduce problems associated with chemical stability and poor gene transfection efficacy (Malmsten 2013). By varying size and surface composition, smart engineered nanocarriers can avoid the RES and can improve drug targeting to the specific area of the body. These nanocarriers have strong affinity to thiol, amine, and carboxyl groups. Such affinity enables surface-modified nanocarriers as promising material for targeting drugs and/or chemotherapeutics. For example, gold nanoparticles modified with albumin do not show any hemolytic response and are used as a carrier for drug release in systemic circulation (Khullar et al. 2012).

Fig. 6.3
figure 3

Structure of inorganic nanocarriers

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4.1 Metallic Nanoparticles

Numerous efforts have been made to fabricate various noble metal NPs and to investigate their optical properties. The use of gold (Au) and silver (Ag) NPs has been explored broadly in biological imaging and gene and drug delivery owing to their unique dimensions, tunable functionalities, and controlled drug release (Yang et al. 2012)

The strong plasmon resonance, surface-enhanced Raman scattering, and photothermal conversion of metallic AuNPs have been exploited for imaging and therapy of cancerous cells (Sekhon and Kamboj 2014). The scattering from a single AuNP or AgNPs has been found to be of many orders of magnitude than the signal from single fluorophores or QD. AuNPs are good candidates for drug and gene delivery. Andres and coworkers synthesized AuNPs and grafted them with thioctic acid-PEG-folic acid. The resulting folate-PEG-coated NPs were soluble in aqueous solution over pH range 2–12 and at electrolyte concentration of up to 0.5 M NaCl. They also demonstrated the specific uptake of folate-conjugated AuNPs by folate-receptor positive tumor cells (Dixit et al. 2006). Brown et al. (2010) functionalized gold NPs with a thiolated PEG capped with a carboxylate group for improved drug delivery of oxaliplatin anticancer drug (Brown et al. 2010). In an another attempt, doxorubicin-loaded PEGylated AuNPs were successfully delivered for overcoming multidrug resistance of MCF-7/ADR cancer cells (Wang et al. 2011). Recently, Manivasagan et al. (2016) synthesized fucoidan (Fu)-capped AuNPs for delivery of doxorubicin and photoacoustic imaging. Fu was used as a capping and reducing agent. The in vitro cytotoxicity effect of Fu, Fu AuNPs, and DOX-Fu AuNPs were studied on human breast cancer cells (MDA-MB-231). It was found that in vitro cytotoxicity of DOX-Fu AuNPs can be recognized to the greater uptake potential of Fu AuNPs, thus establishing the role of AuNPs as capable carriers for the delivery of anticancer drugs (Manivasagan et al. 2016). AuNPs have been used in gene therapy by using viral and non-viral vectors to transport foreign genes to somatic cells. Mirkin and coworkers demonstrated functionalized AuNPs with single-stranded thiolated oligonucleotide for gene therapy. They found that modified AuNPs have higher binding affinity for complementary nucleic acids and exhibited greater cellular uptake than unmodified NPs (Rosi et al. 2006). Niidome et al. (2006) synthesized spherical AuNPs to investigate the release of plasmid DNA through pulsed-laser-induced fragmentation and dissociation. The result indicated that laser power could release plasmid DNA without significant degradation. Braun et al. (2009) reported novel lipid-based gold NPs for delivery of siRNA that provided temporally and spatially controlled cellular delivery for gene silencing, with the help of direct endosomal release mechanism activated via pulsed laser treatment.

4.2 Silica Nanoparticles

Humans are introduced to silica, the most generous substance on earth surface and widely used in our daily life from various sources such as clays, rocks, and sand (Barik et al. 2008). Fumed silica, colloidal silica, etc. are industrial silica extensively used in microelectronics, optical communication, and food additives. SiNPs have attracted key attention for tumor-targeted drug delivery due to their enormous biomedical applications. There are two major types of SiNPs for tumor-targeted drug delivery, i.e., core/shell SiNPs (C/S-SiNPs) and mesoporous silica nanoparticles (MSNs) (Vallet-Regi et al. 2007; Vivero-Escoto et al. 2010; Malmsten 2013). But MSNs have served its groundbreaking potential in drug and gene delivery due to good physical and chemical stability, high loading capacity, zero premature release, and controlled release of drug molecules with proper rate of release to achieve an effective local concentration. MSNs possess unique features of tunable particle size/pore size/morphology, high surface area and pore volume, biocompatibility, and facile surface functionalization (Yang et al. 2012; Malmsten 2013; Pandita et al. 2015b). Recently, researchers have made several efforts to synthesize hybrid MSNs (AuNPs and magnetic NPs as functional components) loaded with drug molecules (Sekhon and Kamboj 2014). These multifunctional nanosystems have potential applications like targeted drug delivery, in vivo imaging, and theranostic applications (Pandita and Lather 2015). For example, researchers successfully loaded camptothecin in fluorescent MSNs and confirmed the release of drug to various types of human cancer cells to induce cell death (Lu et al. 2007). MSNs are promising vehicle for drug delivery, on account of their successful encapsulation and delivery to the target sites, for drugs such as quercetin, paclitaxel, and doxorubicin (Sarkar et al. 2016; Jia et al. 2013; Fu et al. 2016; Zhang et al. 2016b). Suwalski and coworkers developed a MSN-based gene transfer system that allows an acceleration of tendon regeneration by PDGF-B gene transfer. The MSNs were modified with amino and carboxyl group and in vitro and in vivo transfection efficiency were evaluated. The result showed that modified MSNs did not have signs of inflammation or necrosis in the tendon, kidney heart, and liver of rat during a period of 1.5 months (Suwalski et al. 2010). Roy et al. (2005) used organically modified MSNs to deliver the non-viral DNA sequence for gene delivery. Pan et al. (2012) synthesized peptide-MSN conjugates for targeting cell nuclei and delivering the anticancer drug doxorubicin into targeted nucleus, killing cancer cells with high efficiency. This study provided an effective strategy for design and development of cell-nuclear-targeted drug delivery systems.

4.3 Quantum Dots

QD are semiconductor nanocrystals whose optical and electronic properties are strongly size-dependent (Alivisatos 1996). Larger QD have a smaller bandgap and emit lower energy photons (toward the red side of the spectrum) and smaller QD have a larger bandgap and emit higher energy photons (toward the blue side of the spectrum). QD of various semiconductor compositions, such as ZnS, CdS, ZnSe, CdTe, and PbS, have size-tunable fluorescence emission that range between the UV and infrared range (Medintz et al. 2005). QD can replace organic dyes in many applications and offer several advantages over organic dyes, such as photo stability, long fluorescence lifetime, broad absorption, and narrow, tunable emission spectra (Chan et al. 2002). These properties make QD well suited for applications such as single molecule tracking (Dahan et al. 2003), time-gated imaging (Dahan et al. 2001), and multiplexed imaging where organic dyes would be of limited use due to their photo bleaching, short fluorescence lifetime, and narrow absorption spectra (Ghazani et al. 2006). Their photo stability and tunable emission spectra also enable long-term multicolor imaging, which can offer real-time insights into how cells and proteins interact (Jaiswal et al. 2003). They have also been used to track metastasis. The ability to tune their fluorescence into the near infrared window has also enabled their use in optically guided surgery and in vivo cancer imaging (Gao et al. 2004). The surface of QD can be coated or conjugated with polymers in order to enhance its biocompatibility and extend stability against biochemical reactions. Gao et al. (2004) successfully coated triblock amphiphilic copolymer over CdSe/ZnS QD which protected QD from hydrolysis and enzymatic degradation. They also found that in vivo imaging results of QD probes can be targeted to tumor site through active and passive mechanisms but active targeting is much efficient than passive targeting (Gao et al. 2004). In another study, Johari-Ahar et al. (2016) synthesized L-cysteine-capped CdSe QD conjugated with methotrexate (MTX) and evaluated their internalization and cytotoxicity in the KB cells. Based on cytotoxicity assessment results, drug-QD conjugate was found to efficiently internalize into cancer cells and eradicate MTX-resistant KB cells with IC50 of 12.0 μg/ml as compared to free drug molecules, i.e., 105.0 μg/ml. These findings proposed MTX-QD as promising material for targeted therapy of MTX-resistant cells (Johari-Ahar et al. 2016). Further, QD has been found as promising approache in gene delivery. For example, Yang et al. (2014) functionalized PEI on QD to study the effect of particle size and PEI coating on the efficiency of gene delivery into human mesenchymal stem cells. QD of several sizes (5, 10, 15, and 20 nm) were used in the study. It was found that PEI-coated QD exhibited high gene transfection efficacy as compared to unmodified QD. Particularly, QD with largest particle size (20 nm) exhibited much higher uptake capability and greater gene transfection efficacy.

4.4 Carbon Nanotubes

The rapid development of carbon nanotube (CNT)-based technology in many different fields rendered this material not as much unknown as few years ago for scientific community. CNTs are classified into single-walled carbon nanotubes (SWNTs) and multiwalled carbon nanotubes (MWNTs). These exhibit unique physicochemical properties like high aseptic ratio, high mechanical strength, ultralight weight, and high electrical and thermal conductivity (Yang et al. 2012). They have been demonstrated to be potential carriers for a wide variety of agents such as drugs, DNA, proteins, and genes. Functionalization of CNTs is receiving attention in biomedical applications. With the help of covalent functionalization of CNTs through side-walls and tips, these can be soluble in wide range of solvents (Madani et al. 2011). Also, functionalization of CNTs is responsible to determine its biocompatibility and cytotoxic effect. It was reported that the higher the degree of CNTs functionalization, the safer the material is, thus offering the potential exploitation of nanotubes for drug administration (Prato et al. 2008). Although the clearance mechanism of CNTs is still not clear, numerous studies have reported that accumulation of functionalized CNTs takes place in the RES after i.v. administration. Kang et al. (2010) reported that endocytosis of larger SWNTs (100–200 nm) takes place mainly through clathrin-mediated pathway, whereas smaller SWNTs having size 50–100 nm are internalized employing both clathrin- and caveolae-mediated pathway.

Although amine functionalized CNTs have good binding efficiency for DNA, their transfection efficiency is low (Singh et al. 2005). PEI was reported to be employed on the surface of CNTs to enhance DNA binding and cell uptake (Liu et al. 2005). Later, Liu et al. (2008) developed SWNTs functionalized with phospholipid branched PEG via a cleavable ester bond to deliver the paclitaxel drug at tumor site. They found that functionalized SWNTs owing to prolong blood circulation and tenfold higher tumor PTX uptake by SWNTs through EPR presented higher tumor suppression efficacy compared to clinical Taxol in a murine 4T1 breast cancer model. They also found that drug molecules were excreted via biliary pathway without causing toxic effect to normal cells. The clinical status of different inorganic nanocarriers is listed in Table 6.3.

Table 6.3 Clinical status of inorganic nanocarriers
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5 Conclusion

A variety of nanocarriers are being investigated for the drug and gene delivery. The size of these nanocarriers is responsible for the wide range of promising properties. The ideal nanocarrier can be used to achieve the suitable kinetic properties such as long circulation time, excellent biocompatibility, selective targeting to tumor sites, etc. Nanocarriers are developed to enhance the pharmacological effect and therapeutic properties of drugs. Several nanocarriers are being developed and under preclinical stages with few of them already being used in clinical cancer care like liposomal-based formulation Doxil® (doxorubicin), DepoCyt® (cytarabine), Myocet® (doxorubicin), etc. Nanoscale-based delivery strategies have begun to make a significant impact on global pharmaceutical market. NPs in gene therapy have also drawn significant attention as a potential method for treating both acute illnesses and chronic diseases.

The near future may cling to the appearance of new commercial nanocarrier-based products. However, this revolution of technology in terms of NPs is a challenge, since challenges are high in terms of investments, but due to increasing competition and industrial occurrence, the reimbursement can be greater.